Embed Size (px)
Citation preview
Investigation of portal vein blood flow in cirrhotic portal
hypertension using computer-based and physical modelling
methods
Svetla Bogomilova Petkova
For the degree of Doctor of Phylosophy
Swinburne University of Technology
Melbourne, Australia
2008
Abstract
Portal hypertension commonly accompanies liver cirrhosis and complicates the
condition by adding extra risk to the patient survival chances.
This thesis investigates the blood flow through the portal vein in portal
hypertension using a combination of physical experiments and computer
modelling and simulations. Tissue culturing techniques are discussed and a
novel design bioreactor is developed to accommodate the specific requirements
of such a system within the in vitro growth of a blood vessel on biodegradable
scaffold. That bioreactor is used for the physical experiments and Laser Doppler
Anemometry measurements and can be adopted for a range of tissue engineering
purposes. Computational Fluid Dynamics using commercially available software
to create non-Newtonian flow representations in idealised portal vein model
with and without additional obstructions is shown to have good agreement with
the experiments. The model provides a range of useful and otherwise difficult to
get information about the flow of whole blood and each blood cells group in
terms of velocity, particle paths, pressure, strain, wall shear stress and
combinations of those parameters.
The research presented in this work shows the impact of additional obstructions
within the portal vein on the flow, and the need to further investigate the
apparent impact the different blood cell types (their size and concentration) have
on the overall flow pattern. It is important to understand the factors impacting on
the flow as they are a determining factor in patient with liver cirrhosis and portal
hypertension survival chances.
i
Acknowledgements
This research was possible thanks to the work, dedication and help of the following people (and the list is not possible to be completed):
My lecturers and course co-ordinator during my Masters Degree and work colleagues thereafter, who taught me the basis of what I know now, and showed me that I can achieve anything if I work hard enough.
Dr. Enzo Palombo as my Co-ordinating supervisor who managed to stay positive and believe in me even when I didn’t.
Dr. Tony Barton, my second supervisor for his critical review of my work and constant questioning.
Peter Robb and the team in the Engineering workshop for helping me design and manufacture the stainless steel parts of the bioreactor.
Andrew Moore for his help with polymers and biomaterials and his “home-made” scaffolds. Andy in IRIS and my friend Tony Acquadro for their help learn how to operating the LDA.
Ian Birchall for introducing me to the great people in Melbourne Royal Hospital (Thanks to Prof. Robert Gibson), for giving me support and believing I can do the work.
Alamgir Hossain, a fellow research student, who helped with the development of the computer model and taught me how to utilise the available software, and for his friendship. Dr. Jamal Naser, for his supervision and advice related to the CFD component of this thesis. And for saying “it’s possible, yes, it’s do-able”.
My fellow postgraduate research students, my friends, who constantly told me to keep going and were there to support me (and remind me that live is waiting for me past the completion line).
The Swinburne Student Union and Swinburne University Postgraduate Association, not only for their support, but for giving me a different area to work on (student rights and representation) while writing my thesis. This goes especially to Sally Skinner, Terry Eyssense and Wayne Cupido for their guidance and help during difficult times.
To all my friends, who have been waiting for years to have a dinner with me or go to the movies, your sacrifice is appreciated and I’ll be making it up to you.
And lastly, but definitely not least, to my family, to all the people who mean the World to me and live thousand of miles away, thank you for allowing me the time to complete my research and supporting me from so far away.
ii
I, Svetla Bogomilova Petkova, declare that this thesis:
Contains no material which has been accepted for the award to the
candidate of any other degree or diploma;
To the best of my knowledge contains no material previously published
or written by another person except where due reference is made in the
text of the thesis; and
Where the work is based on joint publications, discloses the relative
contributions of the respective workers or authors.
There has been no professional editing on this thesis
The publications resulting from this work have been identified and their
prior publication has been acknowledged.
7th February 2008
iii
Table of Contents
CHAPTER 1 Introduction
1.1. Background 1
1.2. Aims of the research 3
1.2.1. Expected contribution of this research 4
1.2.2. Limitations of this research 4
1.3. Structure of the thesis 5
1.3.1. Thesis outline 5
1.3.2. Steps used in this research 7
1.4. Practical contributions to knowledge 7
CHAPTER 2 Literature Review
2.1. Anatomy and Physiology of the Liver – brief introduction 9
2.1.1. Introduction to the basics liver zoning 11
2.1.1.1. Structure of the liver 11
2.1.1.2 Structure of blood vessels 13
2.1.2. The effects of Cirrhosis on Liver tissue 15
2.1.3. Cirrhosis as a disease 16
2.1.4. Clinical Problems associated with Cirrhosis 16
2.1.5. Methods of dealing with Cirrhotic Liver 18
2.1.6. Direction of blood flow in Cirrhosis 19
2.1.7. Regeneration of the Liver 20
2.1.8. Factors for scar-production and for regeneration 21
2.2. Portal Hypertension 22
2.2.1. Nature of Portal Hypertension 22
2.2.1.1. Intrahepatic portal hypertension 24
2.2.1.2. Prehepatic portal hypertension 25
2.2.1.3. Posthepatic portal hypertension 26
2.2.2. Ways for overcoming portal hypertension or its complications 27
2.2.3. Effects of portal hypertension in some liver disease treatments 28
2.2.3.1. Partial liver transplantation 28
2.2.3.2. Orthotopic liver transplantation 29 iv
2.3. Blood flow through the Liver 29
2.3.1. Portal Circulation 30
2.3.2. Reverse flow 33
2.3.3. Spontaneous reverse flow and arguments against its existence 34
2.3.4. Streamline flow 36
2.3.5. Hepatic artery and portal vein blood flow relationship 37
2.4. Determining and regulation the Liver blood flow 38
2.5. Shunting 39
2.5.1. Nature of the shunts occurring during portal hypertension 40
2.5.2. Types of shunts depending on shunted blood volume 41
2.5.3. Transjugular Intrahepatic Portosystemic Shunt (TIPS) 45
2.5.3.1. Nature of TIPS – surgical procedures 46
2.5.3.2. Complications of portal hypertension treated with TIPS 47
2.6. Other treatments and methods for overcoming portal hypertension 51
2.6.1. Mechanical devices 51
2.6.2. Bioartificial Liver 62
2.6.2.1. Bioreactors – types, principles and some problems 63
2.6.2.2. Hollow fibre bioartificial liver 64
2.6.2.3. Fluidised Bed Bioartificial Liver 65
2.6.3. Non-shunt operations 65
2.6.4. Sclerotherapy 66
2.6.5. Balloon Tamponade 66
2.7. Medical conditions associated with Portal Hypertension 66
2.7.1. Complications of Liver Transplantation 67
2.7.1.1. Shunts 67
2.7.1.2. Stenosis 67
2.7.1.3. Embolization of the portal vein or one of its branches 68
2.7.2. Hepatopulmonary syndrome 69
2.7.3. Variceal bleeding 69
2.7.4. Hepatic hydrothorax 70
2.7.5. Portal hypertensive gastropathy 70
2.7.6. Porto-pulmonary hypertension 71
2.7.7. Other liver disease conditions 72
v
2.8. Vessel blockages and thrombosis 73
2.9. Cell Adhesion 77
2.9.1. Terminology 78
2.9.1.1. Tissue Engineering 78
2.9.1.2. Importance of the endothelial cell lining of blood vessels 80
2.9.2. History of cell adhesion and cell seeding 81
2.9.3. Methods of cell seeding and cell adhesion 82
2.9.3.1. Endothelial cell seeding 82
2.9.3.2. Cell differentiation 83
2.9.3.3. Possible improvement in endothelial cell growth 84
2.9.4. Difference between static and dynamic conditions for cell seeding
and cell adhesion 85
2.9.4.1. Electrostatic endothelial cell seeding method 87
2.9.4.2. Dynamic cell seeding technique 87
2.10. Scaffold requirements 88
2.11. Scaffolds and scaffold materials 90
2.11.1. Comparison of materials 92
2.11.2. Techniques for manufacturing scaffolds 97
2.11.3. Dacron prostheses 98
2.11.4. Non-woven scaffold 98
2.11.5. Modified ePTFE and PTFE 99
2.11.6. Biodegradable scaffold 99
2.11.7. Other types of scaffolds 100
2.12. Coating of biomaterials 101
2.12.1. Coating the material with a layer of endothelial cells 101
2.12.2. Coating with fibronectin and E-selectin 103
2.12.3. Carbon-deposited surface and Diamond-like Carbon coating 104
2.12.4. Coated with grafted adhesion peptides 105
2.12.5. Encapsulation of the graft 105
2.13. Why pulsatile flow is important 105
2.13.1. Waveforms and pulsatility 106
2.13.2. Endothelial cells – graft relationship 108
2.13.3. Effect of hemodynamics on endothelial cells 108
vi
2.13.4. Vessel compliance 109
2.14. Other methods and approaches for addressing the problems of cirrhosis
of the Liver and vessel transplant in general 110
2.15. Future work 112
2.16. Conclusions 113
CHAPTER 3 Commonly used methods and parameters in blood flow modelling and
thesis specific used theoretical and experimental methods
3.1 Introduction 114
3.1.1. Methods for measurement of portal blood flow 114
3.1.2. Generic flow measurements 117
3.1.2.1. Electromagnetic Flowmeters 117
3.1.2.2. Ultrasonic Methods 117
3.1.2.3. Electrical Impedance Techniques 119
3.1.2.4. Tracer Techniques 119
3.1.3. Flow measurements based on pressure gradients, flow in other
blood vessels, or numerical estimation. 119
3.1.3.1. Portal vein blood flow measurement based on pressure gradient
between portal and hepatic veins 121
3.1.3.2. Measurements based on pressure drop within the blood vessel 121
3.1.3.3. Volume flow measurements 122
3.1.3.4. Measurements of Portal Vascular Resistance 122
3.1.3.5. Measurements of the Hepatic and Portal Venous Pressure 123
3.1.3.6. Relationship between vessel diameter and velocity 124
3.1.3.7. Measurement of Portal blood flow 124
3.1.4. Doppler flowmetry 125
3.2. Laser Doppler Anemometer 125
3.2.1. Principle of Laser Doppler Anemometry 125
3.2.2. Models used for LDA 130
3.3. Computational Fluid Dynamics (CFD) Modelling 133
3.3.1. Non-Newtonian flow 137
3.3.2. Numerical simulations and modelling 137
3.3.3. Limitations of CFD and future work 139
vii
3.3.4. FLUENT model used in this thesis 140
3.4. Blood flow properties 141
3.4.1. Rheological properties of human blood 141
3.4.1.1. Properties of blood in patients with chronic liver disease 147
3.4.1.2. Non-Newtonian properties of blood 147
3.4.1.3. Blood viscosity 148
3.4.1.4. Blood cell behaviour as suspended particles in the blood flow 149
3.4.2. Newtonian flow 150
3.4.3. Factors governing portal vein hemodynamics 150
3.4.4. Specific factors impacting on branched vessels 151
3.4.5. Hemodynamics of vascular grafts 153
3.4.6. Impact of portal hypertension on vascular hemodynamics 153
3.5. Theoretical reasoning for the proposed model 157
3.5.1. Laminar flow of blood 157
3.5.2. Basic laws governing the cardiovascular system 158
3.5.3. Fluid mechanics definitions 160
3.5.4. Commonly used assumptions 160
3.5.5. Additional effects impacting the circulation 162
3.5.6. Wave propagation in the cardiovascular system 163
3.5.6.1. Pulsatile flow 164
3.5.6.2. Importance of hemodynamics on modelling of blood flow 166
3.5.7. Tissue culturing studies 166
3.5.7.1. Tissue culture methods 166
3.5.7.2. Preparation and test studies 167
3.6. LDA experimental set-up 169
3.6.1. Bioreactor and LDA 169
3.6.2. Fluid and model vessel 170
3.6.3. Pump and reservoirs 171
3.7. Ideas for future work 171
3.7.1. Heating of fluid 171
3.7.2. Developing the model vessel from different material 172
3.7.3. Blood flow modelling 172
3.8. Conclusions of the Chapter 172
viii
CHAPTER 4 Bioreactor
4.1. Introduction 174
4.1.1. The use and historical development of Bioreactors 174
4.2. Types of Bioreactors 176
4.2.1. Tissue culture static bioreactors 176
4.2.2. Extracorporeal bioreactor systems – some examples 177
4.3. Development of the Bioreactor used in this study 177
4.3.1. The initial idea and design 177
4.3.1.1. Parallel plate bioreactor 178
4.3.1.2. Pulsatile Bioreactor – where the idea came from 180
4.3.1.3. Another example of pulsatile bioreactor – aortic heart valve growth 182
4.3.2. First prototype of the Bioreactor 183
4.3.3. New, simplified Bioreactor 185
4.4. Requirements of a bioreactor for tissue culture of blood vessels 191
4.5. Advantages and disadvantages of the bioreactor 195
4.5.1. Advantages of the new bioreactor design 195
4.5.2. Disadvantages of the new design 197
4.6. Future work and optimization of the device 197
7. Conclusion 198
CHAPTER 5 Measurements, Simulations and Results
5.1. Geometry and Grid generation 200
5.1.1. Grid generation 200
5.1.2. Scaling 201
5.2. Model assumptions 201
5.2.1. Geometry 201
5.2.1.1. 3-D geometry 201
5.2.1.2. Size, diameter and branching 202
5.2.2. Flow 202
5.2.2.1. Common flow assumptions 202
5.2.2.2. Observations 203
5.3. Benefits and limitations 204
ix
5.4. Visualisation 205
5.5. Mathematics and parameters 209
5.5.1. Continuity and Momentum Equations 210
5.5.2. Viscosity equations 211
5.6. FLUENT models: simulation results 214
5.6.1. First model visualization 214
5.6.2. Comparison between models with and without obstructions 219
5.6.3. FLUENT comparisons of different velocities 224
5.6.3.1. Velocity magnitude 225
5.6.3.2. Visualization opportunities with FLUENT 226
5.6.4. Particle tracking 230
5.6.5. Newtonian verses non-Newtonian flow 233
5.6.6. Idea for portal vein shunt 239
5.6.6.1. Non-Newtonian flow visualization 240
5.6.6.2. Newtonian flow visualisation 243
5.7. Visualisation of LDA measurements 244
5.7.1. Measurements and different visualisation opportunities 244
5.7.2. Comparison between visualization using different vector lengths 250
5.7.3. Comparison between normal and obstructed models 252
5.7.4. The most appropriate vector length 254
5.8. Comparison between CFD and LDA models 256
5.9. Conclusions 262
CHAPTER 6
Conclusion 264
Future work 268
References 271
Appendix 1 i
Appendix 2 xi
Appendix 3 xvii
Appendix 4 xviii
Appendix 5 xxi
x
List of Figures CHAPTER 2 Literature Review
Figure 2.1. Anterior and Posterior views of the Human Liver 10
Figure 2.1.1.1.1.Structure of the liver 12
Figure 2.1.1.1.2. Blood vessel network within a hepatic lobule 13
Figure 2.3.1.1. Human blood circulation 30
Figure 2.3.1.2. Portal circulation and systemic circulation anastomosis 31
Figure 2.3.1.3.Portal circulation with most small intestines removed
and the liver turned upwards and backwards. 32
Figure 2.3.1.4.Normal Hemodynamics 33
Figure 2.3.1.5. Splanchnic steal theory 33
Figure 2.5.2.1. Distal splenorenal shunt 42
Figure 2.5.2.2. H-shunt 42
Figure 2.5.2.3. Portocaval shunt 42
Figure 2.5.3 TIPS placements principle 46
Figure 2.6.1.1.Medtronic pump 57
Figure 2.6.1.2. VFP pump designed by Yambe et al. (2000) 58
Figure 2.6.2.1. Extracorporeal BAL circuit schematic representation 62
CHAPTER 3 Methods Used in Modelling Experiments
Figure 3.2.1.1. Experimental setting of LDA – basic operational principles 127
Figure 3.2.1.2. The probe and the probe volume 127
Figure 3.2.1.3. Doppler frequency to velocity transfer function for a frequency
Shifted LDA system 128
Figure 3.2.2.1. Silicone cast of varicose vein (stage 1) 131
Figure 3.2.2.2. Rigid polyester resin model of the varicose vein (stage 2) 131
Figure 3.2.2.3. Elastic silicone rubber model based on the previous two models
(stage 3) 131
Figure 3.2.2.4. Representation of bi (tri) furcation 132
Figure 3.2.2.5. Comparison between the two glass models and scale 133
Figure 3.3. Wall shear stress and flow streamline patterns at different cross
xi
sections of the aorta using CFD analysis 135
Figure 3.3.2.1. Normal and Protrusion model vessel of prosthetics graft
connection to a blood vessel 138
Figure 3.3.4. Grid for (a) normal model; (b) blocked model 141
Figure 3.4.4. Rigid, simple model of equal dimension branching of a vessel 151
Figure 3.5.6. Relationship between pressure, area of vessels and the speed blood
moves with through them in the circulation 163
Figure 3.5.6.1.1. Pulsation of pressure in the circulation 164
Figure 3.5.6.1.2. Mean velocity and velocity fluctuations in the cardiovascular
system 165
Fig 3.6.1. Schematic representation of the experimental set-up 169
CHAPTER 4 Bioreactor
Figure 4.2.1. Multichamber pulsatile bioreactor (a) and experimental
set-up (b) 175
Figure 4.3.1.1.Parallel Plate Bioreactor 179
Figure 4.3.1.2.1 Pulsatile Bioreactor 181
Figure 4.3.1.2.2 Schematic representation of pulsatile Bioreactor 181
Figure 4.3.1.3. Pulsatile bioreactor for tissue engineered aortic heart valve 182
Figure 4.3.2.1. Diaphragm at neutral and above neutral position
and the pressure chamber 184
Figure 4.3.2.2. Schematic diagram of the prototype bioreactor 185
Figure 4.3.3.1. Simplified Bioreactor: front view with silicone tubing attached 186
Figure 4.3.3.2. Upper (left) and lower (right) rings with connected spikes 187
Figure 4.3.3.3. Inlets in front and outlets in the background 188
Figure 4.3.3.4. Glass lid of bioreactor with the gas inlet/outlet 188
Figure 4.3.3.5. Assembled Bioreactor with both rings without the lid 189
Figure 4.3.3.6. Technical drawing of the lower ring 190
Figure 4.4.1. Joining of lower and perfusion chambers 192
Figure 4.4.2. Close up of central area where the scaffold is attached 194
Figure 4.4.3. Inner cylinder and spikes 194
Figure 4.4.4. Bioreactor during LDA measurements 195
xii
CHAPTER 5 Measurements, Simulations and Results
Figure 5.4.1. LDA visualisation experiments of obstructed vessel in bioreactor 205
Figure 5.4.2. Cross-section of the inner diameter of an ideal glass vessel
and the outside square shape 206
Figure 5.4.3. LDA visualization experiments with normal (non-obstructed)
model in bioreactor 206
Figure 5.4.4. Laser beam through the glass model – side view 207
Figure 5.4.5. LDA experimental setup 207
Figure 5.4.6. Water tank with bioreactor submerged in it 208
Figure 5.5. Calculating k-n parameters for use in the Power Law equation 209
Figure 5.6.1.1. 3-D Grid of simplified blood vessel structure 214
Figure 5.6.1.2. Contours of static pressure (Pascal) in a vessel, assuming identical
outflow from both branches 215
Figure 5.6.1.3. Velocity vectors coloured by velocity magnitude (m/s)
in a vessel, assuming identical outflow from both branches 215
Figure 5.6.1.4. Contours of wall shear stress (Pascal) in a vessel 216
Figure 5.6.1.5. Contours of boundary cell distance in a vessel 216
Figure 5.6.1.6. Cutting planes parallel to the x-axis of velocity vectors
coloured by velocity magnitude 217
Figure 5.6.1.7. Cutting planes positioned within the model representing
contours of velocity magnitude 218
Figure 5.6.1.8. Scaled Residuals 218
Figure 5.6.2.1.(a) Contour of velocity magnitude on an x-y plane cutting
through the middle of the geometry (Z=0 plane) without obstructions 220
Figure 5.6.2.1.(b) Contour of velocity magnitude on an x-y plane cutting
through the middle of the geometry (Z=0) with obstructions 220
Figure 5.6.2.1.(c): Closer view of the contour of velocity magnitude
on Z=0 plane with obstructions 220
Figure 5.6.2.2.(a) Contour of static pressure on Z=0 plane without obstructions 221
Figure 5.6.2.2.(b) Contour of static pressure on Z=0 plane with obstructions 221
Figure 5.6.2.3.(a) Contour of strain rate on Z=0 plane cutting through
xiii
the middle of the geometry without obstructions 222
Figure 5.6.2.3.(b) Contour of strain rate on Z=0 plane with obstructions 222
Figure 5.6.2.4.(a) Contour of wall shear stress on Z=0 plane without obstructions 223
Figure 5.6.2.4.(b): Contour of wall shear stress on Z=0 plane with obstructions 223
Figure 5.6.3.(a). Contours of strain rate when velocity is set at 0.0015m/s 224
Figure 5.6.3.(b). Contours of strain rate when velocity is set at 0.0225m/s 224
Figure 5.6.3.1.1. Z=0 plane contours of velocity magnitude when velocity
is simulated at 0.07m/s 225
Figure 5.6.3.1.2. Z=0 plane velocity magnitude when velocity is simulated
at 0.015m/s 226
Figure 5.6.3.2.1. Velocity vectors at the wall with grid coloured by velocity
magnitude (m/s) when velocity is simulated at 0.07m/s 227
Figure 5.6.3.2.2. Velocity vectors at the wall without the grid coloured by
velocity magnitude (m/s) when velocity is simulated at 0.07m/s 227
Figure 5.6.3.2.3. Contours of the wall shear stress (Pascal) (at the wall) 228
Figure 5.6.3.2.4. Velocity vectors in the Z=0 plane coloured by Y velocity (m/s) 228
Figure 5.6.3.2.5. Velocity vectors in Z=0 plane coloured by static pressure 229
Figure 5.6.3.2.6. Histogram of frequency of velocity magnitude 229
Figure 5.6.3.2.7. Static pressure verses position in the model in the Z=0 plane 230
Figure 5.6.4.1. Leukocyte particle traces coloured by velocity fraction 231
Figure 5.6.4.2. Erythrocyte particle traces coloured by velocity fraction 231
Figure 5.6.4.3. Platelet particle traces coloured by velocity fraction 232
Figure 5.6.4.4. Plasma particle traces coloured by velocity fraction 232
Figure 5.6.4.5. Single stream particles 233
Figure 5.6.5.1. Contours of velocity magnitude (m/s) for a Newtonian flow 234
Figure 5.6.5.2.(a) Profiles of velocity magnitude (m/s) non-Newtonian auto scale 235
Figure 5.6.5.2.(b) Profiles of velocity magnitude (m/s) scaled 235
Figure 5.6.5.2.(c) Contours of velocity magnitude (m/s) for non-Newtonian 236
Figure 5.6.5.3. Contours of velocity (m/s) in Z=0 plane for Newtonian flow 236
Figure 5.6.5.4.(a) Profiles of velocity magnitude (m/s) for non-Newtonian flow 237
Figure 5.6.5.4.(b) Contours of velocity in Z=0 plane for non-Newtonian flow 237
xiv
Figure 5.6.5.5. Contours of wall shear stress (Pascal) for Newtonian flow 239
Figure 5.6.5.6. Contours of wall shear stress for non-Newtonian flow 239
Figure 5.6.6.1.1. Contours of velocity magnitude (m/s) 240
Figure 5.6.6.1.2. Contours of wall shear stress (Pascal) 241
Figure 5.6.6.1.3. Contours of wall shear stress (Pascal) 241
Figure 5.6.6.1.4. Contours of static pressure 242
Figure 5.6.6.2.1. Velocity vectors coloured by velocity magnitude (m/s)
in the default interior 243
Figure 5.6.6.2.2. Contours of velocity magnitude (m/s) at Z=0 plane 243
Figure 5.6.6.2.3. Contours of Static pressure (Pascal) in the default interior 244
Figure 5.7.1.1. LDA measurements of mean velocity in the normal vessel 246
Figure 5.7.1.2. Representation of mean velocity vectors only (vector length 300)
without the contours in a normal vessel (without obstructions) 247
Figure 5.7.1.3. Representation of the obstructed vessel with vector length of 300 247
Figure 5.7.1.4. Close up of the obstructed vessel with vector length of 300 248
Figure 5.7.1.5. A representation of rake of stream traces in a close up view
of the obstructed vessel with vector length of 300 249
Figure 5.7.1.6. A representation of rake of stream traces in a close up view of the
obstructed vessel with vector length of 1500 250
Figure 5.7.2.1. Normal vessel with vector length of 300 (same as Figure 5.7.2.2.) 251
Figure 5.7.2.2. Normal vessel with vector length of 500 252
Figure 5.7.3.1 Normal model with vector length of 700 253
Figure 5.7.3.2. Obstructed model with vector length of 700 253
Figure 5.7.4.1. Normal model with vector length of 300 254
Figure 5.7.4.2. Obstructed model with vector length of 1000 255
Figure 5.7.4.3. Obstructed model with vector length of 1500 255
Figure 5.8.1. CFD points from the inlet to the middle of the branching in
normal vessel 256
Figure 5.8.2. CFD points in the right branch in normal vessel 257
Figure 5.8.3. CFD points in the left branch in normal vessel 257
Figure 5.8.4. Points measured using LDA in normal vessel 258
xv
Figure 5.8.5. Height of vessels used in LDA and CFD simulations 259
Figure 5.8.6. Height points in the vessels from the CFD simulations
up until just after the trunk obstruction 259
Figure 5.8.7. Height points in the vessels from the CFD simulations
at the area of branching and just below the obstructions in the branches 260
Figure 5.8.8. Velocity in the Left and Right branches around
obstructions for the CFD simulated model 260
Figure 5.8.9. LDA measured points in the obstructed model below the Branching 261
Figure 5.8.10. LDA measured points in the obstructed model in the branching 261
xvi
List of Tables
CHAPTER 2 Literature Review
Table 2.5.2. Comparison between before and after distal splenorenal
shunting in 10 patients 44
Table 2.11.1.1. Comparison of different materials 93
Table 2.11.1.2. Comparison of commercially available suture materials 95
CHAPTER 3 Methods Used in Modelling Experiments
Table 3.3.4.1. Dimensions of the geometry used in this model 141
Table 3.3.4.2. Non-Newtonian power law parameters used in this study 141
Table 3.4.1.1. Composition of human blood 143
Table 3.4.1.2. Blood properties according to the literature 144
Table 3.4.1.3. Portal vein flow 145
Table 3.4.6.1. Comparison between control and Child-Pugh classified patients 154
Table 3.4.6.2. Duplex Doppler Ultrasound measurements of vessel diameter
and average velocity in 14 patients with alcoholic cirrhosis and
portal hypertension 155
Table 3.4.6.3. Duplex Ultrasonographic measurements in 22 control
and 29 PH patients 155
Table 3.4.6.4. Comparison between control and cirrhotic patients
using Doppler Ultrasonography data 156
Table 3.4.6.5. Difference between healthy individuals and patients
with portal hypertension 156
xvii
List of Abbreviations Transjungualar Intrahepatic Portosystemic Shunt (TIPS)
Laser Doppler Anemometry (LDA)
Computational Fluid Dynamics (CFD)
Portal hypertension (PH)
Portal vein (PV)
Transcutaneous Doppler Sonography (TDS)
Intravascular Doppler Sonography (IDS)
Endothelial Cells (EC)
Human Umbilical Vein Endothelial Cells (HUVEC)
Fused Deposition Modelling (FDM)
Rapid Prototype technique (RP)
Pellethane® (PEU)
NH4 plasma treated PEU (PEU-NH4)
H2O plasma treated PEU (PEU-H2O)
Fluorinated PEU (PEU-fluorine)
Polyethylene imine treated PEU (PEU-PEI)
Heparin treated PEU (PEU-heparin)
Polyethylene (PE)
H2O plasma treated PE (PE-H2O)
CF4 plasma treated PE (PE-CF4)
xviii
CHAPTER 1
Introduction
1.1. Background
There is expectancy that by 2010 the worldwide incidence of hepatitis C, a major
cause of liver disease, will probably exceed that of HIV (Gornam 2001). The most
popular treatment for liver disease still is transplantation, but with the number of
patients awaiting an operation increasing, the proportion of available donors
decreases. Growing liver tissue in vitro using “bioreactors” is a potential alternative
treatment. Such devices can be used for both growing new liver tissue for
transplantation or for drug testing. However, one of the setbacks of any in vitro
grown tissue is that there is no certainty that it can reproduce the function of the
original tissue under in vivo conditions.
Portal hypertension is one of the major complications in patients with diseases of
the liver, such as liver cirrhosis, veno-occlusive disease, idiopathic extrahepatic
portal vein obstruction and pre-hepatic portal idiopathic pathology. Portal
hypertension is a build up of pressure in the portal vein, usually just before it enters
the liver. Thus, there is a significant reduction of the blood flow to the liver, which
causes diminished blood supply to the organ and reduction of normal function.
When portal hypertension occurs, the most common solution to restoring the normal
blood flow is to use shunts. A shunt is a graft, which takes the blood from one part
of a blood vessel to another or from one blood vessel to another to bypass problem
areas in the vessel.
In the past, different types of shunts, such as side-to-side, end-to side and
transjungualar intrahepatic portosystemic shunt (TIPS), have been used. However,
in most cases they were used only as a temporary assist device and had to be
replaced after some time due to thrombosis, graft failure and platelet adhesion on
the graft. These are only some of the many reasons why shunts are usually used
only as a short-term solution to assist medication treatment. A more permanent
1
solution to these problems is to create a bypass or shunt which has minimally
thrombogenic properties, from a minimally platelet adhesive material with a high
degree of durability.
The material from which a shunt or bypass is made generally can be classified as:
synthetic (man-made), autograft (from another vessel of the same patient),
homograft (from another human donor) or animal modified vessel (treated to be
non-antigenic). As well as all the above there are natural shunts that are collateral
pathways of thin-walled new vessels, which the body “creates” to bypass the
diseased vessel. These occur spontaneously, are unpredictable, uncontrollable and
are at high risk of breaking thus causing internal bleeding in the patient. They are
not the type of shunts considered in this thesis.
Synthetic shunts are usually made from polyglycolic acid, polyglucolactic acid,
polylactic acid or polycaprolactone.
A homograft (or animal modified) blood vessel used as a bypass has problems of
incompatibility and rejection by the immune system of the recipient, resulting in the
patient having to take life long immune suppression drugs, thus being more exposed
to other diseases.
Nowadays, it is regular practice to create a bypass and shunt from another vessel of
the patient. There are many benefits to this method compared to the synthetic grafts
or other homografts including high biocompatibility, low risk of rejection and lower
tissue stress.
However, there are certain limitations to this practice of using other blood vessels
(usually venous autografts) – sometimes there are no suitable veins due to other
diseases causing occlusion of the veins (like thrombosis), high blood pressure or
even structural changes in the vessel wall; and the issue of vessel compliance has
been under investigation for the last two decades.
Thus, creating a graft in vitro, using the patient’s own tissue cells to form the new
vessel is a potential alternative solution.
2
In this research, methods for the creation of in vitro autografted blood vessels using
the patient’s own cells grown on biodegradable scaffold in a bioreactor to overcome
portal hypertension are proposed.
1.2. Aims of the research
This thesis comprises several parts, attempting to bring together a variety of current
research in the areas of liver and portal hypertension, in vitro cell growth and
scaffolding, design and building of necessary equipment, through to computer
modelling and simulations.
The overall aim of the work is to compose a model of portal vein blood flow in
portal hypertension, with emphasis of the flow dynamics in cirrhotic patients, to
potentially assist surgery assessment and provide information to the medical
practitioner.
One of the expected outcomes of the work is to investigate the possibility of
creating in vitro shunts for non-emergency operations by growing tissue on
biodegradable scaffold in a specially designed and novel type of bioreactor.
The computer simulations will be repeated as physical experiments using Laser
Doppler Anemometry (LDA) measurements utilising the new bioreactor, using a
vessel model with the same geometry as the one used in Computational Fluid
Dynamics (CFD) simulations and liquid with same viscosity as blood (as used in
the computer model). Good agreement between the computer model and physical
experiments is expected. As a result of this work a simple, easy to use, user friendly
and accurate model will be developed, which can be adapted to suit researcher’s
needs and individualised to each patient. That model should represent velocity,
velocity vectors and particle pathways for whole blood or each of the elements
found in blood in terms of: pressure; wall shear stress; strain rate; the impact
viscosity has on the flow; the changed flow when obstructions in the portal vein are
assumed or known; and any combination of the above. Blood will be treated as non-
Newtonian and cell types are to be described using their size and density. The
density, and size if needed, can be given as the exact values of the patient and the
model will individualise the flow visualisation to reflect the input parameters.
3
Although the 3-dimensional geometry created in this research will be simplified, it
can be re-drawn based on patient scans, thus giving a realistic view of the flow.
This simple geometry is the basis for the model, which needs customisation to be
useful in medical practice.
This work also outlines areas for future research and suggests improvements and
optimisations to all sections of the thesis. Clinical studies are highly recommended
for assessing the practical usefulness of the model.
1.2.1. Expected contribution of this research The aim of this research is to investigate ways to overcome portal hypertension and
provide a long-term solution via an autograft blood vessel and/or better
understanding of blood flow characteristics using computer modelling. The research
aims to show ways for growth of a new blood vessel in vitro on a biodegradable
scaffold under simulated flow conditions, using the patient’s own cells, in a
specially designed bioreactor. After the scaffold degrades, a new blood vessel will
result. This will eliminate the need to take another vessel from the patient to create
the shunt. The vessel will be biocompatible, resulting in less stress for the body.
The computer model is intended to further our understanding of hemodynamics in
the portal vein in cirrhotic portal hypertension and to allow for easy customization
of design of the vessel scaffold to reflect the specifics of each patient. It has to
allow for non-experts to run the modelling and understand the results. The intent is
to minimize the opportunity for misinterpretation of the results and to permit for
simulation of blockages within the portal vein.
1.2.2. Limitations of this research The implementation of the method of growing a blood vessel in vitro using cells
harvested from the patient in routine medical practice will have certain limitations.
For example, it is not suitable for emergency situations, as the technique will
require months for appropriate tissue growth, special equipment and trained
operators.
Tissue engineering and biodegradable scaffolding materials (including their
manufacturing) are still under development and there is no ‘perfect’ method to
achieve a replica of natural vessels. In this research, a difficulty that was
4
encountered was the manufacture of the scaffold, as either the material with the
desired characteristics was unsuitable for the currently available Rapid Prototype
manufacturing equipment, or the machines needed independent research and
development to be able to produce an appropriate scaffold from available materials.
The computer model has to be adjusted and adapted manually and accurate
parameters have to be specified by the operator. Currently, the shape of the vessel
(the grid generated with Gambit) needs to be re-drawn each time depending on the
scanned image of the vein in each patient. If the medical image can be converted
directly into the computer model (novel computer coding is required) this limitation
can be eliminated.
Most of the limitations can be addressed in future research, and so the major
limitation of this thesis is the lack of clinical testing and validation.
1.3. Structure of the thesis
1.3.1. Thesis outline The structure of the thesis is described in brief below.
In the first chapter a brief introduction to the problem and the methods currently
used for solutions are given.
The second chapter deals with the physiology of the liver, and the clinical problems
of cirrhosis and portal hypertension. Here, the most commonly used ways for
dealing with portal hypertension are discussed, and an emphasis on the blood flow
to the liver and the variety of techniques used to measure it are described. Types of
shunts currently used, as well as other treatment methods (including mechanical
devices, bioartificial liver support and bioreactors) are presented. Overviews of
complications and related medical conditions, as well as vessel obstruction, in
portal hypertension are presented.
Further, the literature review chapter gives general overviews of tissue culturing and
methods for cell adhesion to different scaffolding materials, and the variety of
coating materials available are presented. Special requirements for small grafts are
described as the branches of the portal vein and the possible new shunt model can
be smaller than 6mm (in this thesis the sub-branches only are smaller than 6mm
diameter). The importance of pulsatile flow for cell growth with respect to the
5
endothelial cell layer and vessel compliance are given. The experiments performed
for this part of the thesis are outlined and future work recommendations are made.
In the third chapter, methods for blood flow measurement in the portal vein are
described in light of their benefits and drawbacks. Laser Doppler Anemometry is
introduced and explained, as this was the method used for physical experimental
measurement in this work. The following sub-chapter on Computational Fluid
Mechanics using Gambit and FLUENT software is presented to enhance our
knowledge on the features of the computer model developed as part of this thesis
and provide the physics and mathematics background to the model. Rheological and
hemodynamic characteristics of the blood flow in healthy and cirrhotic patients and
the impact portal hypertension has on them are combined with general theories and
studies of the blood flow in the human cardiovascular system. The tissue culturing
experiments are presented in brief, and the experimental set-up for the physical
measurements is given. Some more ideas for future work are outlined.
Chapter Four describes the new bioreactor prototype and gives the rationale of the
optimisation work carried out. The operating principle and real-life images of the
device are presented. How this bioreactor meets the requirements for tissue-
culturing device is argued, and possible improvements are suggested. The
advantages of the novel design are explained and the reasons for creating a new
device are given.
The Results chapter (Chapter 5) deals mainly with the computer model and gives
details of geometry generation, flow properties used, the mathematics and
visualisation used in the model. Separate visualisations for velocity, pressure and
wall shear stress are given for both an obstructed (assuming three regions of
obstruction in the vessel) and normal simplified model of the portal vein.
Comparisons between results obtained by using assumptions for Newtonian and
non-Newtonian flow, and using a range of velocities, are presented. Multiphase
flows are simulated for blood plasma, erythrocytes, leukocytes and platelets by
solving separate non-Newtonian power law equations and these are presented as
velocity particle tracking. Visualisation of LDA measurements and the limitation of
that method are presented, followed by the Conclusions chapter and
recommendations for future work. This work highlights areas requiring further
6
research in both theory and practice and provides ideas for some future
investigations.
1.3.2. Steps used in this research
For the purpose of in vitro growth of a blood vessel, tissue-culturing methods are to
be used on a variety of materials under steady conditions. This will enable the
establishment of working protocols and evaluation of the methods currently used
and allow for possible areas of future research to be developed. Only basic
experiments were carried out, as this area fell outside of the scope of the research
undertaken. Nonetheless, it has been identified as an area of great interest and
importance.
Design and manufacturing of a novel bioreactor for in vitro tissue engineering will
provide optimum conditions for graft development and subsequent testing with the
device to find areas for optimisation.
Computer modelling (simulations and visualisations) with a vessel and flow similar
to the ones used in the physical experiments need to show good agreement with
results obtained with both methods. The model needs to be simple to use and to
provide information otherwise difficult to obtain via physical measurements.
The computer model has to be easily adaptable to represent the specific conditions
of individual patients, as they will have different physical parameters of the portal
vein (size and shape), blood composition (blood cell components) and flow
parameters (depending on the underlying reasons of the disease and their overall
health).
This thesis represents research carried out aiming to achieve all of the above,
namely: technical design and manufacturing of a novel bioreactor; computer
generation and verification of the vessels modelled: and the opportunity for tissue
culturing in sterile laboratory conditions.
1.4. Practical contributions to knowledge
The following areas are expected to have practical impact on research in this field
and to further our knowledge and understanding of those topics:
7
• The design and development of a new type easy-to-use low cost bioreactor for
in vitro growth, which can assist in providing an appropriate environment for
tissue graft culturing. It must have a minimal number of parts, be easy to
sterilise (by autoclaving), and be re-designed to provide versatility and
flexibility thus allowing individualisation of each graft geometry.
• Creating a computer model capable of simulating the flow specifics of
individual patients would provide the opportunity to study the blood flow
behaviour and possibly determine regions of higher risk of the development of
obstructions and blockages. The viability of the in vitro grown vessel could
then be ‘tested’ before manufacturing and tissue culturing by simply creating
new models or modifying existing models to be an exact replica of the desired
vessel.
• Providing a comprehensive review of liver functioning and blood flow in portal
hypertension, cell seeding methods and scaffolding materials, discussing the
physics of portal vein blood flow, and highlighting some of the key factors
which can assist in the development of the desired in vitro grown autograft to
be used as shunt or as replacement for the original vessel. These will make it
easier to understand the complexity of factors contributing to the problems in
the field.
• Modelling the components of blood (the cells) as separate phases will further
our understanding on the impact their density and size has on the blood flow
and portal vein in hypertension.
The practical aim of this research is to investigate ways to overcome portal
hypertension and propose a long-term solution via an autograft blood vessel and/or
better understanding of blood flow characteristics using computer modelling.
Investigations of various tissue culturing techniques and available biodegradable
materials is carried out and a novel bioreactor to host the graft is developed. The
computer model is intended to further our understanding of hemodynamics in the
portal vein in cirrhotic portal hypertension and to allow for easy customisation of
design of the vessel scaffold to reflect the specifics of each patient. This model is
easy to use and does not require computer programming, as data can be simply
inputted, by non-experts, once the model has been created.
8
CHAPTER 2
Literature Review
2.1. Anatomy and Physiology of the Liver – brief introduction The liver is the second-largest organ in the human body after the skin. It
is a spongy, reddish brown gland that lies just below the diaphragm in the
abdominal cavity. Lying in the upper right side of the abdomen, most of it is
protected by the ribs. The liver lies beneath the diaphragm in the abdominal
cavity, and above the right kidney (Miller and Leavell 1972). Weighing
approximately 1.5 kg, it pulses continuously as 1.5 litres of blood pass through
it every minute. There are reservoirs of blood in the liver called venous sinuses,
which can hold up to 3.5 litres for boosting blood volume in emergencies (The
Primary Biliary Cirrhosis Foundation 2000). It serves to metabolise
carbohydrates and store them as glycogen; metabolise lipids (fats, including
cholesterol and certain vitamins) and proteins; manufacture digestive fluid, bile;
produce blood-clotting factors; and destroy old, worn-out red blood cells. Two
large lobes, the right and the left, make up most of the liver while the smaller
quadrate and caudate lobes are attached to the right lobe. The lobes are made up
of lobules – six-sided cells arranged in sheets one cell thick – that are closely
arranged around blood vessels, bile ducts, lymph vessels, and nerves. Certain
reticuloendothelial cells (Kupffer cells) line these lobules and play a role in
immunity. Approximately three sides from each six-sided cell are in contact
with a blood vessel, and three are adjacent to a bile duct. The lobules are
grouped in clusters so that the bile manufactured by each lobule passes down a
common duct, which connects to larger ducts that lead to the common hepatic
duct. This duct joins with the cystic duct of the gallbladder and enters the
duodenum along with the pancreatic duct of Wirsung.
The histological composition of the liver is predominantly hepatocytes
(around 78%), non-parenchymal cells (including Kupffer cells and endothelial
cells – around 6%), with the remaining volume being sinusoidal lumen, biliary
channels and intracellular spaces (Puviani et al. 1998).
9
The liver has an unusual portal (venous) circulation that has two sets of
capillaries instead of one. Veins draining the upper intestinal tract unite to form
a large vein that then divides again, as an artery would, to form capillary-like
structures, called sinusoids within the liver; sinusoids reunite to form large
veins that return blood to the heart. As liver cells die, the fibrous tissue is
deposited around the small vessels in the liver. This fibrous tissue disturbs the
portal circulation of blood through the liver. The destruction of liver cells
impairs the liver’s ability to store nutrients and to detoxify chemicals produced
by the body or coming from outside.
Figure 2.1. Anterior and Posterior views of the Human Liver (Encyclopaedia Britannica 2005) The liver has a dual blood supply: the portal vein represents a low-pressure
system without significant pulsatility of flow; and the arterial system provides
high pressure and pulsatile flow. Normally, hepatic arterial flow amounts to
200ml/min corresponding to 20-30% of the total blood supply of the liver (Arey
1957; Flemming et al. 1983; Sherlock 1978; Stary et al. 1992; Strandell et al.
1973). Intrahepatic vascular resistance is increased 5-fold in cirrhosis (Moriyasu
et al. 1986), leading to an increase in portal pressure and opening of portacaval
shunts. In advanced cirrhosis, even reverse portal blood flow can occur (Hűbner
et al. 2000). With decreasing portal venous inflow to the liver and spontaneous
or therapeutic portosystemic shunts, the liver blood supply becomes largely
dependent on hepatic arterial perfusion.
10
2.1.1. Introduction to the basics liver zoning
According to Prof. Dr. MD. J. Reichen (1998) from the Department of Clinical
Pharmacology, University of Berne, Switzerland, and the handbook for students
at the same University, the liver’s weight is between 1200 – 1500 g, and the
blood passing through it is 25-30% from the heart minute volume, from which
1/4 - 1/3 is coming from the hepatic artery and 2/3 – 3/4 from the portal vein.
The blood circulation in the liver is divided into three zones. Zone 1 contains
the cells closer to the centre of the liver, next to the portal vein system.
Regeneration of the liver begins from this zone. The cells in different zones
have different functions. The importance of the zoning of the liver is seen, for
example, in acute right or left heart failure, where the centrilobular area is most
affected.
The following correlation (Reynolds et al. 1954) between the different pressures
measured or calculated for the liver blood flow can be used:
CSP = WHVP – FHVP (2.1.1.)
where CSP is corrected sinusoidal pressure, WHVP is the pressure of collapsing
of the hepatic veins and FHVP is the free pressure of the hepatic veins. The
increased resistance in cirrhosis is the reason for appearance of “backward
flow”.
2.1.1.1. Structure of the liver
As described above, the liver consists of lobules and is surrounded by a thick
capsule (Arey 1957).
11
Figure 2.1.1.1.1.Structure of the liver (Encyclopaedia Britannica 2005)
The liver is a highly vascular tissue. It receives around 25% of its blood from
the hepatic artery (Arey 1957; Balaz 2000; Flemming et al. 1983; Netter 1964;
Sherlock 1978; Stary et al. 1992; Strandell et al. 1973) and 75% from the portal
vein, which has a diameter of 1.09cm (Netter 1964). The liver blood inflow
from the hepatic artery is oxygenated, and from the portal vein is carrying
nutrients and drugs from the GI tract. The hepatic artery and portal vein fuse
within the liver and mix in the sinusoids and blood leaves the liver via the
hepatic vein (Balaz 2000). Sinusoids are microscopic spaces between rows of
liver cells. The hepatic artery and the portal vein divide into fine branches,
which supply blood to the fine bile ducts, which than drain into the sinusoids.
The liver derives its own supply of oxygenated blood from the hepatic artery,
which branches off the aorta. Blood leaving the liver is collected in the hepatic
veins, which join together into a single hepatic vein that empties the blood into
the inferior vena cava. From there it is passed back to the right side of the heart,
to be pumped to the lungs. The liver is composed of minute divisions called
lobules, separated from each other by connective tissue. These lobules are made
up of columns of cells surrounded by tiny channels known as canaliculi, into
which the bile secreted by the liver cells is released. These channels unite to
form progressively larger ducts, culminating in the hepatic duct. A central vein
12
is located within each lobule unit, and liver cells radiate outwards from the vein
in all directions. At the periphery of each lobule are 5 to 7 portal triads. Each
portal triad consists of a hepatic portal vein, a hepatic artery, and a bile duct.
The blood from the artery and vein flow into the sinusoids, and then it is
transported past Kupffer’s cells and into the central vein.
Figure 2.1.1.1.2. Blood vessel network within a hepatic lobule (Miller and Leavell 1972 p.419), where the hepatic vein is the outflow, and the branches of the hepatic artery and portal vein are the inflow.
2.1.1.2 Structure of blood vessels
Human blood vessels have a unique structure due to the specialised functions
they perform. The commonality between arteries and veins is that their wall
consists of three layers. The external layer is called adventitia, the middle is
media and the inner one is intima. These layers have different thicknesses and
their ratio to one another is different throughout the vascular system. The intima
is composed predominantly of endothelial cells. The wall of any blood vessel is
predominantly water (about 70%) in combination with elastin, collagen and
other fibres, smooth muscle and endothelial cells.
The portal vein has a very thin wall (Arey 1957) compared to other veins in the
human body.
Blood vessels are viscoelastic, non-linear, not homogenous nor isotropic. The
impact those properties have on the blood flow is discussed in a Chapter 3 of
this thesis.
13
The structure, geometry and mechanical properties of blood vessels change with
changes in the stress in the vascular system. This can be due to a variety of
factors, with organ disease, injury and surgery being some of the most common
ones.
The diameter of veins is dependent on the transmural pressure – at 1mmHg the
vein is almost collapsed and at 10mmHg the vein is round.
The most common variations of the portal vein include trifurcations, right
posterior branch arising from the main portal vein and right anterior branch
arising from the left portal vein (Gallego et al. 2002).
A prospective ultrasound study undertaken by Arti et al. (1992) determined the
prevalence of variants of the intrahepatic branching of the portal venous system
as following: of the 507 patients examined, 55 (10.8%) had trifurcation, 24
(4.7%) had a right posterior segmental branch arising from the main portal vein,
22 (4.3%) had a right anterior segmental branch originating from the left portal
vein, and one (0.2%) had absence of the horizontal segment of the left portal
vein. Not one patient had complete absence of the right portal vein in this series.
The remaining 405 (79.9%) patients had normal distribution of the portal
venous system, with some patients of the normal group having minor variations
in distribution.
Forty years ago, M.D. John H. Carter and co-workers (1961) published a paper
on their study of the changes in the hepatic blood vessels in cirrhosis of the
liver, which is still one of the best works in this area. During autopsy, they
gradually injected solutions of differently coloured vinyl plastic in acetone in
the hepatic artery, portal and hepatic veins. Following quantitative comparisons
they found that while the inflow tract and the outflow tract are approximately
equal in the normal liver, in the cirrhotic liver the outflow tract cross section
area is 55% of that of the inflow tract.
The thin-walled, unprotected, low-pressure hepatic veins are easily destroyed
and distorted by the disease process and hence show the greatest change. The
higher-pressure portal vein protected by the stroma of the portal triad is subject
to the same ravages but is more resistant and consequently exhibits less change.
14
2.1.2. The effects of Cirrhosis on Liver tissue
In Cirrhosis, damaged liver cells get replaced by fibrous tissue, and the
regeneration of liver cells does not follow the normal process but rather forms
nodules surrounded by that fibrous tissue (Worman 1995). Fibrous tissue causes
increase in resistance, leading to a decrease of blood flow to and through the
organ (Worman 1995). Due to this fibrous tissue portal hypertension is always
present in cirrhosis regardless of the cause of the disease.
Cirrhosis is defined as “a diffuse process characterised by fibrosis and the
conversion of normal liver architecture into structurally abnormal nodules”
(Anthony et al. 1978)
Among physicians there is agreement that cirrhosis is a generic term for hepatic
disease of varied etiology, for example, alcohol abuse, iron overload, drugs and
chronic active hepatitis. However, there is no universal accepted definition.
There are an agreed-upon characteristic of cirrhosis, however, that serves as
guidelines and indicators to aid diagnosis:
1. The architecture of the total liver is disorganised and altered by
interconnecting fibrous scars formed in response to hepatocytic injury and loss,
2. The fibrosis may take the form of delicate bands but may constitute broad
scars replacing multiple adjacent lobules,
3. Nodules (tiny collections of tissues) are created by the regenerative activity
and network of scars. The nodules vary in size, depending on causation, from
micronodules (less than 3mm in diameter) to macronodules (3mm to several
centimetres in diameter),
4. The parenchymal (functional elements of the liver architecture) is generally
disorganised within micronodules, that is, loss of central veins.
Different classifications can be made for the cirrhotic liver depending on the
cause, hepatic size, stage of the disease, etc.
Abnormal vascular (pertaining to blood vessels) connections develop in the
fibrous scars between the portal, arterial, and venous systems that, to an extent,
bypass the hepatic parenchyma. All forms of cirrhosis, whatever their origin,
are chronic progressive disorders, largely because the causation cannot be
controlled in most instances. Most of the fatalities over time result from liver
failure or one of the consequences of portal hypertension related to the
15
extensive scarring and nodularity. Cirrhosis distorts normal functional elements
of the liver, resulting in reduced perfusion of the diseased organ and portal
hypertension.
2.1.3. Cirrhosis as a disease
Patients with cirrhosis have reduced quality of life, high mortality risk and other
complications accompanying the disease. Cirrhosis is a worldwide health
problem. Cirrhosis can also lead to other systemic complications including
decreased production of blood clotting factors (leading to bleeding), changes in
the metabolism, immune system dysfunction and abnormalities in the brain. The
major causes of cirrhosis of the liver are alcohol abuse, hepatitis B, C and D
infection, autoimmune hepatitis, genetic abnormalities, inherited metabolic
diseases, drugs, and toxins. The most common understanding of cirrhosis is that
it is irreversible as a disease, although the liver is able to regenerate itself after
being injured or diseased (Arey 1957; Galambos 1979). However, there have
been recent studies based on clinical and laboratory data showing that cirrhosis
might be reversible (Iredale 2003).
Investigating the blood flow in order to understand and improve it, thus helping
the regeneration of the cirrhotic liver is one of the aims of this thesis and the
work undertaken.
Diagnosis of cirrhosis usually is made based on the symptoms of the underlying
disease (Netter 1957). Liver biopsy is done either to confirm such condition or
when the symptoms are not clear enough. It is a simple procedure as the liver is
located very close to the skin, involving a small specimen being taken with a
needle.
2.1.4. Clinical Problems associated with Cirrhosis
A serious outcome of cirrhosis is pressure on the blood vessels that flow
through the liver. Because the normal flow of the blood is slowed, pressure
builds in the portal vein, causing portal hypertension. Blood from the intestines
tries to find a way around the liver through new vessels, which have thin walls
and carry high pressure leading to high risk of breakage and bleeding (Bosh and
16
Garsia-Pagan 2000). Thus, lower blood flow and portal hypertension both lead
to progress of the disease and in turn prevents the efficiency of medical
treatments for other diseases (because the liver absorbs medications). By
overcoming portal hypertension and supplying the liver with the normal blood
flow, the regeneration process will be supported and the patient will gain more
time waiting for a transplant or even possible having liver function partially
normalised without the need of transplantation.
Intrahepatic vascular resistance is increased 5-fold in cirrhosis, leading to an
increase in portal pressure and opening of portacaval shunts (Moriyasu et al.
1986).
The sinusoid is fed by both the hepatic artery and portal venous flows.
The formation of the collateral pathways that accompany portal hypertension,
however, should be taken into account when considering portal perfusion. It is
frequently seen in portal hypertension shunts from the coronary vein into the
esophageal varices, from the splenic vein into the left renal vein, and from the
umbilical vein into the veins of the abdominal wall. The coronary vein usually
branches from near the confluence of the splenic and superior mesenteric veins.
About one third of the total blood perfusing the cirrhotic liver may bypass the
sinusoids, and hence functioning hepatic tissue, through these channels
(Galambos 1979).
Summary - in cirrhosis we have on one hand a lower blood flow, and on the
other portal hypertension, both leading to progress the disease and preventing
the efficiency of medical treatment for other diseases.
Solution – ideally the solution will incorporate overcoming portal hypertension
and supplying the liver with the normal blood flow which might support the
regeneration process and the patient will gain more time waiting for a
transplantation or even possible having the liver function normalised without
need of transplantation
Kupffer cells in cirrhosis are more sensitive to Endothelin-1 and platelet-
activating factor, both of which are separate causes for portal hypertension
(Yang et al. 2003).
Patients with liver dysfunction usually have gastrointestinal disturbances due to
obstructed portal vein blood flow (Thompson 1981). One of the functions of the
17
liver is to detoxify the blood of ammonia. In cirrhosis, due to damaged liver
tissue this function is obstructed, thus causing brain-damaging ammonia build-
up in the blood serum (Thompson 1981).
Decreased blood flow to the liver and obstruction of blood flow in the portal
vein and portal circulation can be responsible for complications like blood
backing up in the spleen causing it to enlarge or a backflow from the portal to
the systemic circulation leading to varicose veins in the stomach, esophageus
and rectum (which can rupture, bleed massively and even cause death).
Cirrhosis can also lead to kidney dysfunction and failure (Netter 1957).
2.1.5. Methods of dealing with Cirrhotic Liver
Traditionally, medicinal treatment of the underlying injury and the cause of the
injury are used to treat cirrhosis. While those treatments are useful and can
produce good results, they could be complementary to other methods (such as
surgery) for eliminating the disease and its complications.
So far there are two common ways for dealing with the complications of
cirrhosis – transplantation, which is very expensive and relies on finding a
suitable donor, and liver bypass (to avoid further complications such as
backflow and ruptured vessels). With liver bypass, there is still the problem of
lower blood flow through the liver. In addition toxins cannot be cleaned from
the blood because it does not pass through the liver. This makes the resorption
of medication and the blood filtration much lower than the normal range and
can cause many other damages to the body (such as renal or multiorgan failure
(Sauer et al. 2002) and brain damage (Thompson 1981)).
Bypassing the liver using shunts can occur in two ways – natural (new vessels
created within the body) or surgically created (practiced for more than half a
century and still not very well explored). Nevertheless, specialists agree that
shunts do not help the liver regenerate itself.
The regeneration of the liver starts from surviving cells which are provided with
the best blood supply and therefore obtain the most available oxygen and
nutrient source. Therefore, regeneration begins around the axial channels
(Galambos 1979).
18
Shunt operations, according to the British Liver Trust Information Service
(BritishLiverTrust), involve surgically joining two veins. Usually, the portal
vein and the inferior vena cava are joined. Shunting of blood is effective in
some patients for preventing recurrent bleeding. This operation is carried out,
usually in non-emergency conditions, in patients who have bled from varices
and whose liver function is still relatively good (BritishLiverTrust). One
disadvantage of shunts is a risk of impairment of brain function called
encephalopathy, as a result of toxic chemicals reaching the brain from the gut
because the blood has been diverted away from the liver cells, which would
normally detoxify the blood. That problem is pointed out as a disadvantage of
surgical shunts by the British Liver Trust (BritishLiverTrust).
2.1.6. Direction of blood flow in Cirrhosis
It is very difficult to measure the direction of blood flow in patients due to the
position of the portal vein. Only few studies have involved multiple patients,
usually they are done on individual cases.
A Doppler study carried out by Luigi Bolondi et al. (1990) showed the
following in respect of blood flow directions in patients with cirrhosis:
Only 7% of patients with hepatofugal flow, 1.1% of which in the portal trunk,
2.7% in the splenic vein and 3.2% in the superior mesenteric vein.
The authors (Bolondi et al. 1990) also found that the diameter of the portal vein
was higher by around 2mm in patients with hepatofugal flow compared to
patients with reversed flow.
An interesting finding (Bolondi et al. 1990; Darnault et al. 1989) shows a
decrease in arterial vascular resistance related to chronic and acute impairment
of the liver function. Using the Pulsatility Index (PI) showed significant
decrease of PI in the superior mesenteric artery in patients with cirrhosis and
acute hepatitis, but not in unrelated to cirrhosis portal vein thrombosis (Darnault
et al. 1989).
19
2.1.7. Regeneration of the Liver
There is controversy in the literature about the ability of the liver, and especially
the human liver, to regenerate itself after injury and disease. Some authors are
cautiously saying that usually the underlying liver damage is irreversible (Lai
1997) in patients with chronic liver disease.
Many researchers and literature point to improvement and regeneration of the
liver (Chamuleau 2002; Galambos 1979; Moser et al. 2000l Sauer et al. 2002).
In this respect this project has been developed and completed. According to the
Children’s Liver Disease Foundation (Children’sLiverDiseaseFoundation) some
90% of the liver can be cut away and, providing the remaining 10% is healthy,
the liver will grow back to its original size. There is still work to be done to
verify this, as will be seen in the discussion about small size liver grafts further
in this chapter. Such studies on humans are still far away due to their highly
invasive and uncertain outcome.
In 1953 Charles G. Child and associates (1953) reported experiments in dogs
designed to determine whether portal blood is essential for liver regeneration.
An operation was devised in dogs that accomplished complete diversion of the
portal stream and at the same time provided the liver with profuse supply of
systemic venous blood. According to this and previous studies the liver in
normal dogs regenerates rapidly and completely following partial removal. If,
however, its portal blood supply is compromised either by partial ligation of the
portal vein or side-to-side portacaval anastomosis, or diverted away, liver
regeneration is then inhibited or prevented. Since there is no evidence of any
substance in the portal blood, which might aid such regeneration, this might be
simply due to either lack of portal blood itself, or to simple reduction in the
afferent hepatic blood flow. The data obtained from the study by C.G. Child
(Child et al. 1953) showed that systemic venous blood is capable of supporting
liver regeneration, although not as effective as portal blood, hence the
conclusion is that diminished hepatic blood flow and not portal blood itself is
the reason for failure of liver regeneration. In their experiments regeneration of
the liver averaged 50 % (± S.D. 21) in dogs with portacaval transposition as
compared to 75% (± S.D. 27) in normal dogs subjected to a similar partial
hepatectomy.
20
However, this still has to be proven in humans, as it is not clear whether the
same principle will apply.
In cirrhosis, intrahepatic shunts develop and the sinusoids become capillarized
due to the deposition of collagen and cell necrosis, causing decreased drug
elimination and increased vascular resistance with portal hypertension (Cardoso
et al. 1994). The study carried out by Cardoso et al. (1994) in 13 perfused livers
from cirrhotic rats showed that after doubling the flow, intrahepatic resistance
decreased by 31%. The conclusions made by the authors were that increased
portal blood flow in cirrhotic rats induces a decrease in intrahepatic resistance
without changes in the intrahepatic shunting and improves drug elimination by
the liver without adverse effects on hepatocyte viability.
Here, as in the previous example, human studies need to take place to examine
whether this would be applicable to the human hemodynamics.
2.1.8. Factors for scar-production and for regeneration
A group at the University of Newcastle in the U.K. lead by Dr. Chris Day is
working to uncover the signals, which tell the scar-producing cells of the liver
to activate and produce scar tissue. They have shown the importance of
inhibitors for blocking these signals and the possible therapeutic approach to the
treatment of liver fibrosis.
Fibrosis is a term used to describe a build up of scar tissue as a result of
long-term liver disease prior to developing cirrhosis. Fibrosis results from
sustained wound healing in the liver in response to chronic or repeated injury, a
dynamic process of inflammation and repair (Iredale 2003). To the best of our
knowledge to date the two methods for treatment of fibrosis without liver
transplantation are either removing the cause of inflammation and giving
immunosuppressive drugs, or blocking the signals, which activate the hepatic
stellate cells and promote collagen secretion. An extensive review of the
possible therapeutic interventions in liver fibrosis can be found in the article by
Iredale (2003). These interventions will not be described in detail here as they
are out of the main scope of this thesis. However, the impact of the proposed
work on graft treatment of liver fibrosis would need further studies. Lead by Dr.
21
J. P. Iredale (2003) at the University of Southampton, a group supported by the
British Liver Trust (BritishLiverTrust) is currently working to discover the
factors, which regulate spontaneous recovery from liver fibrosis by
investigating enzymes degrading scar tissues and their interaction with drugs
and inhibitors. Their initial results suggest possible change in the scarring tissue
to normal under the influence of drug treatment. If this or any other similar
study is successful, cirrhosis may become curable. However, at this stage, to the
best of our knowledge there are no clinical trials in patients with cirrhosis.
2.2. Portal Hypertension Portal hypertension is defined as a sustained increase in the portal vein pressure,
usually as a result from obstruction of the blood flow within the portal
circulation. The scarring of the liver is also considered a cause for portal
hypertension (Encyclopaedia Britannica 2005). The normal pressure in the
portal vein is between 5 and 10mmHg. Most medical practitioners consider
pressure above 12mmHg as hypertension (Schiedermaier 2004).
Portal hypertension is usually seen in patients with liver diseases and less often
in patients without disease of the liver, and increases the risk of internal
bleeding in patients with this condition (Society Interventional Radiology
2004).
There is evidence that a pattern of enlarged paraumbilical vein can be used to
predict portal hypertension, i.e. can be used as an indicator for the presence of
portal hypertension (Dirchfield et al. 1992) and at the same time exclusion
criteria for presinusoidal cause of portal hypertension (Kane and Katz 1982).
The importance of preserving portal blood flow for maintaining the hepatic
function in patients with portal hypertension becomes apparent when
considering patients with non-selective shunting whose liver function
deteriorates and often leads to encephalopathy (Ozaki et al. 1988).
2.2.1. Nature of Portal Hypertension
Portal hypertension is the most common and probably most dangerous
complication of cirrhosis (Denk 2004) and other liver diseases
(BritishLiverTrust). Although the mechanisms triggering portal hypertension
22
are not yet defined, some studies have suggested possible and probable causes
of portal hypertension (PH). One of these studies gives a relationship between
the activation of hepatic stellate cells and PH by looking into the mechanisms
that increase the production of endothelin and consequently the increase in
intrahepatic sinusoidal resistance as factors contributing to PH (Iredale 2003).
Haynes et al. (1991) showed that in general veins are more sensitive to
endothelin than arteries. In such a case, advances in medications could provide
a solution to PH in the future. Up until today, to the best of our knowledge, such
drugs have not been created despite the research in this area.
Portal hypertension is caused by increased intrahepatic resistance in most cases,
but can also be due to increase of portal vein flow, or even insufficiency in
other veins participating in the liver inflow and outflow. The splenic venous
flow, for example, affected by large splenomegaly, can cause some degree of
portal hypertension due to the increase in the splenic blood flow.
Patients with portal hypertension have increased portal vein diameter by
approximately 30% and decreased portal vein flow velocity by over 40% (Haag
et al. 1999). Portal hypertension is considered present when the diameter of the
portal vein is larger than 1.25cm and/or the portal vein flow velocity is less than
21cm/sec (Haag et al. 1999). These figures can be used as a guide only and
individual measurements have to be conducted to diagnose PH. Other studies
give different values of the diameter and flow velocity of the portal vein, but the
range is not wide (for example, PV diameter of 14.4±2.4mm was reported in
(Rőssle et al. 1994)).
When PH is present a swelling and twisting of the portal vein can be observed
(varices) leading subsequently to hemorrhage (Zemel et al. 1991).
Depending on the location of the cause for PH it can be classified as
intrahepatic, prehepatic and posthepatic (Cwikiel 2006; Denk 2004).
Some authors (Thomson and Shaffer 2006) prefer classification based on the
cause relative to the sinuses or the site of increased resistance (Schiedermaier
2004), i.e. presinusoidal (which covers the prehepatic and some of the
intrahepatic causes), sinusoidal (covering several of the intrahepatic causes) and
postsinusoidal (covering some intrahepatic and all posthepatic causes).
23
An older classification made by Sherlock (1974) classified PH into two main
groups – presinusoidal and intrahepatic. The presinusoidal is further divided
into extrahepatic, in which the obstruction is in the main portal vein, and
presinusoidal, in which the obstruction is usually in the portal tracts.
In cirrhosis, the portal hypertension can also be classified as presinusoidal,
sinusoidal and postsinusoidal (Sherlock 1974; Thomson and Shaffer 2006).
2.2.1.1. Intrahepatic portal hypertension
When the liver itself is diseased, as in cirrhosis, it becomes a cause for portal
hypertension. Some of the commonly known causes of intrahepatic portal
hypertension are:
• Primary Biliary Cirrhosis
Among the diseases causing the intrahepatic presinusoidal portal hypertension
is primary biliary cirrhosis (Sherlock 1974). The intrasplenic pressure and
umbilical (portal) venous pressure are increased but the wedged hepatic venous
pressure is virtually normal. Usually the obstruction is in the portal veins, but
may be along the sinusoids in the space of Disse.
• Idiopathic portal hypertension
Patients with portal hypertension who cannot be classified into any of the
defined disease categories usually are referred to as idiopathic portal
hypertension patients. It is deemed not to be associated with cirrhosis,
extrahepatic portal vein occlusion, schistosomiasis or any other classified cause
(Ohashi et al. 1998). Some authors suggest that it could be due to thrombosis of
the extrahepatic portal vein that subsequently is recanalized in non-cirrhotic
portal hypertensive patients (Almoudarres et al. 1998). In most cases the portal
vein will have a thick sclerotic wall, stellate of fibrosis will be present in the
portal vein with the advance of the condition, and obstruction of the branches of
the portal triad will be visible. Different names have been given in the literature
to those conditions – hepato-portal sclerosis (Oikawa et al. 1998), (non-cirrhotic
(Schiedermaier 2004)) portal or periportal fibrosis, essential portal
hypertension, Mediterranean cirrhosis, or even cirrhosis of splenic origin.
Idiopathic portal hypertension is associated with satisfactory liver function.
24
• Hepatic schistosomiasis
The liver function is well preserved, and the cause for the hypertension is intra-
pre-sinusoidal vascular obstruction.
One of the major causes of intrahepatic portal hypertension is cirrhosis.
The portal vascular bed is distorted and diminished and the portal blood flow is
mechanically obstructed. About one third of the total blood perfusing the
cirrhotic liver may bypass the sinusoids through collateral venous channels
(Sherlock 1974). Micro-angiographic studies show arterioles entering the
venous channels surrounding the nodules instead of the sinusoids. A pathway
between hepatic arterial and portal venous branches certainly exists in the
cirrhotic liver, because retrograde flow can be shown, although rarely, in the
portal vein (Sherlock 1974). The portal venous obstruction in cirrhosis is
classified in general terms as intrahepatic (Sherlock 1974) or prehepatic
(depending on physical location) and involves increased resistance in the portal
zones, sinusoids, and hepatic veins. In addition to mechanical obstruction
(backward flow theory), the role of increased splanchnic blood flow in portal
hypertension must also be considered (forward flow theory). Vorobioff and
Groszmann (1983) have done studies in support of the forward flow theory in
rats, showing portal venous inflow, not resistance as reason for maintaining
elevated portal venous pressure. Total splenic blood flow is indeed increased
and splenic vascular resistance reduced in patients with cirrhosis and increases
further after end to side portacaval anastomosis although the portal venous
pressure has returned to normal (Sherlock 1974).
2.2.1.2. Prehepatic portal hypertension
Portal vein thrombosis is the major cause for prehepatic portal hypertension
(Denk 2004; Sarin et al. 2006). The causes for this thrombosis include cirrhosis,
tumours, surgical trauma, infections and pregnancy. The prehepatic portal
hypertension occurs before the blood enters the sinusoids of the liver, and most
often is associated with the development of porto-systemic collateral
anastomoses. Depending on their flow rate, porto-systemic anastomoses can be
divided in the following groups (suggestions only):
25
a) Low to moderate flow rate (including gastro-esophageal, hemorrhoidal
and retroperitoneal)
b) Spontaneous and possibly higher flow rates (including spleno-renal,
umbilical, etc.)
Obstruction of the portal vein outside the liver is usually seen in a region close
to a large number of collateral vessels. In most cases the liver is normal and its
architecture is not affected by the obstruction. The presentation of this condition
is portal hypertension. Sometimes the splenic and mesenteric veins can also be
obstructed, contributing to the pressure build up in the portal vein. Extrahepatic
portal vein obstruction can also be a secondary complication of cirrhosis. Thus,
any obstructions to the portal vein have been viewed as relevant to our study
and models will be proposed for patients with portal hypertension, with and
without liver diseases.
Extrahepatic portal vein obstruction in children is a common cause of portal
hypertension (De Ville et al.1999), and in most cases the liver is normal and
bypassing the obstruction can restore the hepatic portal flow and decompress
the portal hypertension.
Partial portal vein occlusion, after portal vein ligation, shows the appearance of
uniform collaterals, resulting in recanalization of the vein (Krupski et al. 2002).
2.2.1.3. Posthepatic portal hypertension
This type of hypertension is observed in cases where the blood flow in the
hepatic veins after exiting the liver is obstructed. The two main conditions are
Budd-Chiari syndrome and veno-occlusive disease.
Veno-occlusive disease obstructs the hepatic outflow and is caused by occlusion
of the central venules and small branches of the hepatic vein. It is a variation of
the Budd-Chiari syndrome, which is characterised by obstruction of the large
hepatic veins caused mainly by hepatocellular carcinoma and bacterial
infection.
Veno-occlusive disease of the liver is a complication of high-dose
chemotherapy and autologous or allogeneic bone marrow or peripheral blood
stem cell transplantation (Zenz et al. 2001). Veno-occlusive disease not only
involves the liver, but is also associated with renal, cardiac and respiratory
26
failure, often requiring intensive care and mechanical ventilation. Depending on
the severity of the occlusion of hepatic veins, the arterial blood flow is partially
or completely drained via the terminal portal branches. If most of the terminal
hepatic veins are occluded, the arterial perfusion depends mainly on the
capacity of the portal outflow, which is defined by the splanchnic collaterals
developing during the early phase of the disease. Such patients may not survive
until sufficient collaterals have been developed and die of hepatic failure or
necrosis of the gut. This is where different types of shunts could play a life-
saving role. In some of the patients with posthepatic obstruction portal vein
thrombosis is also present. A successful transjungualar intrahepatic
portosystemic shunt (TIPS) treatment for patients with veno-occlusive disease
has been discussed later in this chapter in section 2.5.3.2.
2.2.2. Ways for overcoming portal hypertension or its complications
Below, a brief classification of most commonly used procedures for decreasing
portal hypertension and/or increasing blood flow through the portal vein is
presented:
a) Shunts – splenorenal shunt; TIPS; small-diameter H-grafts, etc. are
discussed later in this chapter.
Portal blood preservation is one of the most important requirements for
long-term success of any shunt therapy.
b) Medication - β-blockers for treatment of esophageal varices (Dib et al.
2006; Garcia-Tsao et al. 2004); Glucagon injection for increasing the
collateral blood flow in the left gastric vein (response decreases with the
increase of the grade of portal hypertension (Marsutani et al. 2003);
Vasoconstrictive drugs are used to stop bleeding (Harry and Wendon
2002; Lata et al. 2003)
c) Transplantation – full or partial and from living donor or organ donor
d) Pumps- small devices pumping blood through the portal vein into the
liver
27
e) Gene therapy: myr-Akt gene therapy, for example is aimed to restoring
Akt activation and NO production in cirrhotic liver, thus may help for
the treatment of portal hypertension (Morales-Ruiz et al. 2003)
f) Balloon-occluded retrograde transvenous obliteration is used when there
are large gastric varices with spontaneous splenorenal shunt (Miyamoto
et al. 2003) as a short-term solution.
g) Endovascular embolization of the hepatolienal vessels (either splenic
artery, left gastric artery or collateral pathways), which has shown a high
mortality rate of 29.8% (out of 329 patients) due to esophageal bleeding
(Karimov 2003).
h) Endoscopic variceal sclerotherapy and ligation (Amitrano et al. 2002;
Rikkers et al. 1987) is usually used alone or in combination with
medication or balloon-tamponade
Most of those methods are discussed in more detail later in this chapter.
2.2.3. Effects of portal hypertension in some liver disease treatments
2.2.3.1. Partial liver transplantation
Partial liver transplantation has been developed as a method to assist more
patients receiving liver transplants using one organ to help two or more patients.
The World wide shortage of organ donors and the long waiting lists have called
for this novel form of liver transplantation. So far the minimum volume of
donor liver required for a successful transplantation has been considered 30% of
the standard liver volume of the recipient (Smyrniotis et al. 2002).
Portal hypertension is the usual cause of small graft failure in partial liver
transplantation (Asakura et al. 2003). Control and management of portal vein
pressure has shown to be crucial for the success of liver transplantation,
especially for partial liver transplantation (Asakura et al. 2003).
This is one of the reasons why this thesis is looking into ways of decreasing
portal vein pressure at the same time as maintaining the blood inflow to the
liver by investigating the mechanisms of the flow.
In split liver transplantation the portal flow is redirected through relatively
small-for-size grafts. Common understanding in the literature is that excessive
28
portal blood flow is responsible for graft injury. Many studies have
demonstrated the relationship between portal vein and hepatic artery flow
showing that the increase in portal vein flow leads to decrease in hepatic artery
flow. The diminished arterial flow, which is a cause for hepatic graft thrombosis
and ultimately small-for-size grafts failure, could be potentially prevented by
decreasing the portal hypertension by modifying the portal vein flow
(Smyrniotis et al. 2002). Part of the work in this thesis is related to better
understanding the portal blood flow and thus, creating the ability to modify it.
2.2.3.2. Orthotopic liver transplantation
Not very common, extrahepatic portal hypertension after orthotopic liver
transplantation is usually caused by portal vein stenosis or is due to ligation of
portosystemic shunts (Malassagne et al. 1998). It has been shown in dogs (De
Jonge et al. 2003) that acute ligation of the portal vein can lead to portal
hypertension after partial orthotopic liver transplantation. De Jonge et al. (2003)
recommended banding to divide the portal blood flow between the host liver
and the graft as a better procedure, and concluded that free-flow is not to be
recommended in such patients.
Another rare, but possible complication is portal vein thrombosis
(Bakathavasalam et al. 2001). In such cases re-transplantation and cavo-portal
shunts might be useful to augment portal blood flow.
Portal hypertension combined with ascites, variceal bleeding, esophageal
varices or splenomegaly is making it more difficult for diagnostic methods to
measure portal vein flow. Portal vein stenosis and thrombosis may be
responsible for the failure of the hepatic allograft (Glanemann et al. 2001).
Sometimes spontaneous portal decompression via formation of venous
collaterals can occur. Such spontaneous shunts have been observed in 15.9% of
patients with non-cirrhotic portal fibrosis (Dhiman et al. 2002) and are assumed
to protect these patients from variceal bleeding.
29
2.3. Blood flow through the Liver It was mentioned above that the blood flow through the liver is complex and
depends on numerous factors. In this section a better understanding of the
mechanisms and principles of portal circulation is presented.
2.3.1. Portal Circulation
The portal circulation is a differentiated part of the systemic circulation
(Encyclopaedia Britannica 2005). A certain amount of blood from the intestine
is collected into the portal vein and carried to the liver. There it enters into the
open spaces called sinusoids, where it comes into direct contact with the liver
cells. In the liver important changes occur in the blood, which is carrying the
products of digestion of food recently absorbed through the intestinal
capillaries. The blood is collected a second time into veins, where it again joins
the general circulation through the right atrium (Funk&Wagnalls Multimedia
Encyclopaedia).
Figure 2.3.1.1. Human blood circulation (Miller and Leavell 1972)
30
It is known that there is a hemodynamic interaction between the hepatic arterial
and portal venous vascular beds such that an increase in blood flow through one
circuit leads to an increased inflow resistance in the other circuit, tending to
maintain a constant blood flow through the liver. This effect has been termed
“reciprocity” between the hepatic artery and the portal vein (Chatila et al.
2000).
Peripheral vasodilatation initiates the hyperdynamic circulation in cirrhosis
(Chatila et al. 2000).
Figure 2.3.1.2. Portal circulation and systemic circulation anastomosis (Miller and Leavell 1972 p.420)
31
Figure 2.3.1.3.Portal circulation with most small intestines removed and the liver turned upwards and backwards. The splenic vein is clearly visible (Miller and Leavell 1972 p.417)
The importance of normal blood flow can be seen in all cases of liver disease.
Restoring the blood flow in patients after tumour removal from the main branch
of the portal vein or its bifurcation is essential for postoperative recovery and
long-term survival (Ramesh et al. 2003).
Studies have suggested that portal venous flow is essential for maintaining
normal coagulation (Mack et al. 2003).
In 1995 Newby and Hayes (2002) proposed a hypothesis about splanchnic steal.
They viewed the splanchnic circulation as trying to compensate for the
decreased portal flow not by increasing liver perfusion, but via incremental
shunting of portal blood via porto-systemic collateral anastomoses, thus
creating a ‘steal’ phenomenon. They presented two steals – arterial (from the
systemic circulation into the splanchnic arterial system) and venous (from the
portal vein inflow to the liver into the porto-systemic collaterals). In advanced
liver disease the venous steal can become extreme, thus reversed portal vein
flow may occur.
32
Splanchnic Arteries Hepatic Artery Portal Vein Hepatic Vein
Aorta Vena Cava
Figure 2.3.1.4.Normal Hemodynamics (Newby and Hayes 2002)
Aorta Splanchnic Arteries Hepatic Artery Portal Vein Hepatic Vein Porto-systemic Collaterals Vena Cava
Figure 2.3.1.5. Splanchnic steal theory (Newby and Hayes 2002)
2.3.2. Reverse flow
In a healthy liver there is no (or it is very minimal) backflow from the portal to
the systemic circulations (Worman 1995).
According to a study of 228 patients carried out by Gaiani et al. (1991),
reversed flow in the portal venous system was detected in the portal vein in 7
patients (3.1%), and their study indicated that the actual prevalence of reversed
flow in the portal, splenic, and superior mesenteric veins in a nonselected
cirrhotic population was 8.3%.
Reversal in portal venous flow was found in 8 out of 72 consecutive patients
(approximately 11%) studied for evaluation of portal hypertension and cirrhosis
33
(Tasu et al. 2002), in 17% of patients with portal hypertension and liver disease
(Ozaki et al. 1988), and in only 5.3% from 118 patients in another study
(Dirchfield et al. 1992). Based on those studies, it can be assumed that
approximately 11% of patients with portal hypertension (SD ± 6) have reversed
portal vein flow.
In 2001 Zenz et al. (2001) diagnosed veno-occlusive disease by duplex-
sonography showing reversed flow direction in intrahepatic portal branches and
the extrahepatic portal vein.
The profile of liver vasculature is affected by cirrhosis and portal hypertension.
Some authors have shown the presence of reversed flow in patients with
portosystemic shunts and veno-occlusive disease and the reversal of the portal
venous flow with the advance in portal hypertension (Kok et al. 1999) and
advanced liver diseases (Nerem 1991).
Reversed flow has been found to occur with a higher rate in patients affected by
alcoholic cirrhosis by Luigi Bolondi et al. (1990).
Reverse flow in the portal circulation has also been reported in patients with a
rare anomaly called congenital hepatoportal arteriovenous fistula (Agarwala et
al. 2000).
In case of post-sinusoidal obstruction, reversal in the intrahepatic portal flow
could be observed.
2.3.3. Spontaneous reverse flow and arguments against its existence
A discussion of reversed flow must mention the huge literature review and
study carried out by Moreno et al. (1975). They introduced a physical analysis,
based on first principles, concerning manometric studies, which demonstrated
that the occluded portal pressure could not be used to construct a hydraulic
gradient for portal flow. They found that actual measurements of magnitude and
direction of portal flow provided impressive evidence against the occurrence of
spontaneous reversal of portal flow in cirrhosis.
“Spontaneous reversal of portal flow” implies that the portal vein delivers
hepatic blood into the splanchnic bed instead of delivering splanchnic blood
into the liver as normal (Moreno et al. 1975). Under these conditions, the blood
34
entering the liver through the hepatic artery would find it easier to exit the organ
through the portal vein than through the normal route of the hepatic veins. The
diverted hepatic blood would then fight its way against the incoming splanchnic
flow and eventually reach the right heart via the collateral network.
Apparently, the need for a concept of spontaneous reversal of portal flow was
created by some unexpected results observed after side-to-side portacaval
anastomosis. Although the lower limb of this shunt decompresses the
splanchnic bed, the upper one drains the hepatic blood into the inferior vena
cava.
Over three decades ago, Moreno et al. (1975) made a conclusion still widely
accepted, that “there was no justification for the claim of spontaneous reversal
of portal flow in cirrhosis reported in the literature if the claim was based solely
on the presence of hepatic occluded portal pressure which is higher than either
the free portal or occluded splanchnic pressure”.
In a group of 23 patients with cirrhosis, Reynolds (1955) compared the values
of the sinusoidal pressure using hepatic vein wedged measurements and of the
portal pressure measured simultaneously through the recanalized umbilical vein.
He found that the sinusoidal and portal pressures were almost identical, a fact
that shows the very small resistance existing between the portal vein and the
sinusoids. The results of the important studies of Longmire and associates
indicated that reversal of flow does take place in the hepatic limb of a side-to-
side portacaval shunt, which does not mean that in cirrhosis there is a
spontaneous reversal of portal blood flow. The results from a study involving
294 patients, 273 with cirrhosis and 21 controls (Moreno et al. 1975) showed
reversal flow only in patients after side-to-side portocaval shunt. The
measurement of this group corresponded very well with the ones made by
Sovak and associates (1999) using totally different techniques.
Reversal of portal flow causes diminished flow to the lower cells thus
insufficient oxygen and nutrients supply, which may explain the frequent
encephalopathy following side-to-side portocaval shunts (Sherlock 1978).
However, it seems very unlikely that such reversal of flow occurs with any
frequency. Evidence marshalled by Moreno and his colleagues (1975) casts
35
doubt on its existence as a spontaneous phenomenon. It may be a consequence
of any side-to-side portal systemic shunting operation (Sherlock 1978).
In reviewing the literature Kok et al. (1999) concluded, “Portal venous blood
flow becomes reversed with advanced portal hypertension”
2.3.4. Streamline flow
There is another important issue regarding hepatic blood flow, which has
arguments for and against – streamline flow. Whether blood draining from the
mesenteric and splenic vascular beds is mixed within the portal vein or is
selectively distributed to different lobes of the liver has been the subject of
controversy for a century. Gary F. Gates and Earl K. Dore (1973) produced
arguments in favour of the existence of streamline flow in the human portal vein
(Gates and Dore 1973). Their study in 12 patients without liver diseases
demonstrated streamline flow in the human portal system after injection of
radiolabelled gold into various mesenteric veins. According to the researchers,
portal vein blood is directed predominantly to the right lobe, particularly from
subdivisions of the superior mesenteric vein. In support of this is the fact that
since the right lobe is six times larger than the left (Gray 1959 p.1294-1305), if
the hepatic blood was homogenous, a simular proportion of main lesions would
be expected. However, some vessels prefer perfusion through the right lobe 9
times greater than through the left, which can be seen by the clearance of certain
infections (Gates and Dore 1973).
On the other hand there is a study, carried out three years earlier in
unanesthetized individuals, which re-examined the lobar distribution
(Groszmann et al. 1971). The physical model consists of two reservoirs each
with one outlet and two inlets, one of the latter coming from the divided portal
vein into left and right branches. This study was performed on normal and
cirrhotic patients and a consistent pattern of steaming could not be identified in
either group. Also, altering body position in one participant did not affect the
distribution of mesenteric blood flow. Variation in lobar perfusion may occur in
cirrhosis because of differences in the degree of scarring, vascular distortion,
and portosystemic shunting. The arguments both for and against the existence
36
or otherwise of streamline in the portal blood flow are strong. Therefore, more
studies are needed to evaluate the portal blood flow, which is quite complicated.
In 2002 Gallix et al. (2002) demonstrated streamlining of splanchnic blood in
the portal vein of fifteen normal subjects using MR data.
Both micro sphere and radio labelled tracer studies suggest that there is no gross
difference between the lobes of the liver in the proportion of either arterial or
portal venous blood received (Richardson and Withrington 1981).
A suggestion, based on experimental results using a gamma camera and
scanning over 10 seconds (Sherlock 1978) is that crossing over of the blood
stream can occur in the human portal vein, which supports the view that the
flow is streamlined rather than turbulent.
In the model of blood flow developed in this thesis the streamline of the flow
was assumed and no turbulence was present at the inlet.
2.3.5. Hepatic artery and portal vein blood flow relationship
Hepatic artery occlusion reduces elevated portal venous pressure and this
procedure has been used in the treatment of severe portal hypertension in
humans. Different quantitative studies have shown that occlusion of one inflow
to the liver usually reduces the calculated vascular resistance of the other circuit
by about 20%, and this occurs for both the hepatic artery and the portal vein
(Richardson and Withrington 1981). In humans, hepatic artery occlusion
reduced portal pressure in portal hypertension by about 15% and portacaval
anastomosis increases hepatic arterial blood flow by 6-400%. Quantitatively it
is clear that though interactions do occur between the hepatic artery and the
portal vein, they are inadequate to compensate fully for marked reductions in or
obstructions to one of the inlets (Richardson and Withrington 1981). Of course
any changes in the outflow resistance from the liver will have an impact on the
inflow resistance, i.e. the portal vein and hepatic artery resistances.
During portacaval shunting procedures any compromises in the hepatic arterial
inflow yield in poor prognosis in cirrhotic patients (Richardson and Withrington
1981). Not only would such change increase the portal vein inflow but it can
37
also increase inflow to the intestines and other vascular beds through
arteriovenous shunting.
In the case of liver transplantation arterialisation of the portal vein has been
used for short-term perfusion. Re-canalisation of the portal vein due to
thrombotic occlusion of the portal vein is related to diminished inflow of blood
to the liver. Usually the first treatment is thrombectomy and if it fails then the
use of large collaterals or grafts to the superior mesenteric vein are performed.
Insufficient portal inflow in most studies has been resolved either using portal-
to-caval anastomosis or permanently arterialising the portal vein (Barakat
2003). Arterialisation of the portal vein is regarded as a rescue option for a de-
arterialised liver after the other treatments have failed (Grazi et al. 2003).
Georg H. Hübner and this group (2000) have chosen a thread model of the
hepatic artery with the velocity range of the thread (10-180 cm/s) largely
covering the flow velocities observed in in vivo investigations (6-120 cm/s).
The diameters of the right hepatic artery ranged between 4.3 and 12.4mm, and
with both methods of measurement (transcutaneous and intravascular Doppler
sonography, i.e. IDS and TDS) turbulence were to be expected behind
bifurcations at distance 1.3 times the vessel diameter (Hübner et al. 2000).
2.4. Determining and regulation the Liver blood flow There are three principal determinants of liver blood flow (Richardson and
Withrington 1981): “the hepatic arterial vascular resistance which at constant
arterial pressure governs the hepatic arterial blood flow; the vascular resistance
of the intestine which governs the inflow of blood into the portal vein; and the
intrahepatic portal vascular resistance”. The connections between the portal
venous and hepatic arterial branches entering the sinusoids is one of the factors
enabling the blood flow from a high pressure (arterial) to the low pressure
(portal) systems. In humans, total blood flow is about 800-1200 mL/min
(Richardson and Withrington 1981) of which the hepatic artery supplies roughly
one-third. A hepatic arterial blood flow of 350 mL/min at an arterial pressure of
90 mmHg in a 70kg man (liver weight at 2% body weight = 1400g) gives a
hepatic arterial resistance of about 4mmHg×ml-1×min×100g.
38
There are differences in the liver blood flow volume in the literature, with some
general agreement of approximately 200-300 mL per minute per 100grams of
liver tissue or between 800 and 1000 mL per minute (Miller and Leavell 1972).
In experimental studies different researchers obtained different values for the
hepatic arterial and the total liver blood flow, respectively between 83±10 and
559 mL/min for the hepatic arterial and between 131±24 and 1229±230 mL/min
for the total blood flow (Richardson and Withrington 1981). Clinical levels of
anaesthesia reduce total liver blood flow depending on type of anaesthetic used.
Tygstrup et al. (1962) showed the following properties of the blood flow in
humans: pressure of the hepatic artery of 98mmHg; blood flow from the hepatic
artery of 559ml× min-1 ×100g-1; portal venous pressure of 9mmHg (or about
8mmHg according to (Miller and Leavell 1972)) and hepatic venous (or inferior
vena caval) pressure of 5mmHg.
Preoperative percutaneous transhepatic portal vein embolization is used to
improve the outcome of surgery for hepatocellular carcinoma Kubo et al. 2002,
and done on the right portal vein increases the hepatic functional reserve of the
left lobe as well as its volume.
The regulation of liver blood flow by mechanisms independent of external
innervation or vasoactive agents of extrahepatic origin may be considered in
three ways (Richardson and Withrington 1981): ‘(1) regulation of hepatic
arterial blood flow, (2) regulation of portal venous blood flow, and (3) the inter-
relationships between the hepatic arterial and portal venous inflow circuits’. It
may be that hepatic portal venous pressure and not blood flow is the regulated
variable (Richardson and Withrington 1981) since maintenance of a constant
portal pressure would tend to maintain a normal pressure profile across the
hepatic sinusoids and would minimize changes in outflow resistance from the
intestinal and splenic circulations.
2.5. Shunting Creation of shunts is one of the most popular treatments for overcoming portal
hypertension after medication and non-invasive methods have failed. Other
treatment methods are discussed later in this chapter.
39
In all kinds of surgical shunt procedures the sudden decompression of the
splanchnic circulation induces a blood volume shift into the systemic bed,
which may cause impaired cardiac function and hemodynamics. There is also,
as discussed earlier, the possibility of reverse flow, which can lead to liver
failure even after a successful shunt surgery. The shunts used in practice are
either intrahepatic or extrahepatic. Intrahepatic shunts include the arterioportal
shunt between the portal vein and the hepatic artery and TIPS. Extrahepatic
shunts are, for example, H-graft shunts between the mesenteric vein and the
vena cava and distal splenorenal shunts.
Another way of classifying shunts would be in terms of the amount of blood
flow re-directed from the portal vein of the liver to another blood vessel
(Collins and Sarfeh 1995). There are total shunts and partial shunts. Total
shunts are, for example, non-selective decompressive shunts, where all the
blood bypasses the liver, thus the liver loses its normal blood supply and at the
same time cannot detoxify the blood which goes straight to the systemic
circulation. The effect of this has been discussed earlier in this thesis. Total and
partial shunts are explained in more details below.
The ideal shunt would preserve portal perfusion, minimally alter portal
hemodynamics, and have low risk of causing encephalopathy and liver
dysfunction (Collins and Sarfeh 1998).
2.5.1. Nature of the shunts occurring during portal hypertension
In the portal hypertensive liver, the formation of shunts between the hepatic
artery, the portal vein, and the hepatic vein can often be seen. These shunts can
affect the portal vascular resistance and the effective blood flow of the liver.
Portosystemic shunts are a common complication in patients with portal
hypertension (Dib et al. 2006; Grace et al. 1996; Sekido et al. 2002), especially
as extrahepatic collaterals. The collaterals can be divided into two major
categories: ascending and descending (Eguchi 1986). The ascending collaterals
mainly involve the gastric coronary vein and usually result in rupture of
esophageal varices (Sekido et al. 2002), and the descending collaterals, such as
splenorenal shunt, often cause refractory hepatic encephalopathy.
40
Studies in cirrhotic rats (Tsuchiya et al. 2003) show clearly the presence of
spontaneous portosystemic shunts. Their purpose is to alleviate the hypertension
by redirecting blood through vessels with lower resistance.
Intrahepatic shunts on the other hand, are rarely reported in cirrhotic patients.
One such report of an individual case by Takayama et al. (2001), showed two
portacaval shunts – one from the left portal vein and the other from the
bifurcation of the portal vein occurring.
In trying to bypass obstruction in the portal circulation in portal hypertension
porto-systemic collaterals develop (Dib et al. 2006) carrying high risk of
bleeding (with mortality over 50% from each bleeding episode) (Smith and
Kampine 1984).
The formation of the collateral pathways that accompany portal hypertension
should be taken into account when considering portal perfusion. It is frequently
seen in portal hypertension shunts from the coronary vein into the esophageal
varices, from the splenic vein near the splenic hilum into the left renal vein, and
from the recanalized umbilical vein into the veins of the abdominal wall
(Moriyasu et al. 1986).
According to Ohnishi et al. (1987) it seems that in patients with cirrhosis, the
development of intrahepatic arteriovenous shunts is not as great as that of
portal-systemic shunts, which were found in their study to be considerable and
variable in degree.
2.5.2. Types of shunts depending on shunted blood volume
Portal vein shunts are either total shunts (side-to-side and end-to-side portacaval
shunts) diverting all portal flow from the liver, or partial shunts (small diameter
portacaval H-graft, TIPS etc.).
41
KIDNEYS
Figure 2.5.2.1. Distal splenorenal shunt (Malagó et al. 1998)
LIVER SHUNT
K I D N E Y S
Figure 2.5.2.2. H-shunt (Malagó et al. 1998)
LIVER KIDNEYS
Figure 2.5.2.3. Portocaval shunt (Malagó et al. 1998)
42
Porto-systemic shunts between the portal vein and inferior vena cava were
introduced in 1945 to relieve pressure in the portal vein (Kofidis et al. 2003).
The American Liver Foundation (American Liver Foundation 2004) reports that
¼ of patients receiving such shunts have uncontrollable bleeding and either die
or require emergency surgery. Even more alarming is that the mortality risk in
emergency operations is between 20% and 50%. Reversal of portal vein flow is
possible in patients after side-to-side porto-caval shunts. There could be many
reasons for this phenomenon, some of which include enlargement in the
diameter of the hepatic vein and atrophy of the liver.
Total shunts are more effective in preventing hemorrhage than medical therapy,
but show increase incidence of encephalopathy and liver failure (Collins and
Sarfeh 1998).
Selective shunts, for example distal splenorenal shunts, are not used in alcoholic
cirrhotic patients even though they maintain good portal perfusion.
Partial shunts, such as small-diameter H-grafts, also maintain good hepatic
perfusion, but are not for use in patients waiting for liver transplantation
because they violate the upper right quadrant (Rhee 1993). Small diameter
portacaval H-grafts usually have a diameter of 6, 8 or 10mm (Collins 1998).
The graft can be made from different materials, with polytetrafluoroethylene H-
graft with collateral ablation showing durability and protection against variceal
re-bleeding (Collins 1998). Ideally, this shunt preserves the portal flow by
minimally altering portal hemodynamics. Even though small H-graft portacaval
shunts provide partial portal decompression, the values of the reduction of
portal pressure or the portal-to-inferior vena cava pressure gradient cannot be
used to predict the outcome of the shunting operation (Rosemurgy et al. 2002).
In the later study, 10% of the patients died within 30 days, and within 3 years an
additional 35% died predominantly due to liver failure.
Randomised comparison of H-graft shunts and TIPS show that the later results
in higher incidences of re-bleeding, death and liver failure (Collins and Sarfeh
1998).
43
TIPS can function as a partial shunt if the blood stream’s resistance is high
enough to maintain portal perfusion to the liver (Collins and Sarfeh 1998).
Shunts are either selective or non-selective. Non-selective ones are for example
end-to-side portocaval shunts, small diameter H grafts, proximal splenorenal
shunt, mesocaval interposition shunt and mesocaval C graft (Lai 1997).
Examples of selective shunts are the splenocaval shunt and the distal
splenorenal shunt via the splenic vein and short gastric vein into the renal vein.
One of the problems with distal splenorenal shunts is the possibility in the long-
term for collaterals to develop to decompress the portal pressure, which stays
higher after those shunts when compared to total shunts (Luca et al. 1999).
Distal splenorenal shunts have higher operative mortality, but lower rate of
encephalopathy later compared to porto-systemic shunts (Kofidis et al. 2003).
Some problems with distal splenorenal shunts are the possibility of over 50%
decrease in mean portal blood flow velocity and volume, and the occurrence of
reversal flow. In 1988 the following data was reported (Ozaki 1988) using
Duplex Ultrasonography:
Parameter Before shunting After shunting
Portal vein diameter (mm) 11.13 ± 0.63 10.33 ± 0.55
Portal velocity (cm/s) 9.79 ± 1.35 4.89 ± 1.31
Portal blood volume
(ml/min)
643 ± 152 247 ± 68
Table 2.5.2. Comparison between before and after distal splenorenal shunting in 10 patients (Ozaki 1988)
In many studies selective distal splenorenal shunts are shown to effectively
decompress the spleen and gastroesophageal varices, but to maintain portal
hypertension (Grace et al. 1996; Henderson et al. 1992; Jin and Rikkers 1991;
Rikkers et al. 1987).
Comparisons between end-to-side portacaval shunts and distal splenorenal
shunts show that the first one, despite normalising the portal pressure worsens
the peripheral and pulmonary vasodilatation, while the second one caused no
pulmonary and less peripheral vasodilatation, thus maintaining higher portal
pressure (Luca et al. 1999).
44
Decompression of the portal vein by surgical shunts is a possible prevention
treatment for the formation of ascites due to portal hypertension, but there is no
improvement in the survival rate (Ochs et al. 1995) as the mortality has been
reported to be between 5 and 39%.
Total shunts control bleeding well, but the liver bypass of blood increases the
chance of encephalopathy and liver failure. Partial and selective shunts have a
similar degree of bleeding control, but lower incidence of liver failure and
encephalopathy (Henderson 1995). Devascularization has a higher risk of re-
bleeding, but does not alter the portal blood perfusion or the liver function.
Comparing partial portacaval shunts to direct side-to-side portacaval shunts
shows that the first one better preserves the long-term liver function and
minimises postoperative encephalopathy in patients with cirrhosis and variceal
bleeding (Capussotti et al. 2000).
Percutaneous inferior vena cava-to-portal vein shunt (PIPS) is created through
the caudate liver lobe by a transhepatic puncture through the inferior vena cava
and the portal vein (Quinn et al. 2002), where an endograft
(polytetrafluoroethylene sutured to a Palmaz stent) is placed using a jugular
approach. This shunt has the same principle as TIPS and usually has a pressure
gradient between the portal vein and the inferior vena cava between 1 and
9mmHg (with a mean of 5) (Quinn et al. 2002).
Another classification of shunt procedures (Malagó et al. 1998) can be made
based on the shunt location. Central shunts can be divided into total and partial
shunts, while peripheral shunts can be classified into selective and non-selective
shunts (Malagó et al. 1998).
Other parameters to differentiate the type of shunts available are the material,
surface treatment and elasticity of the shunt.
2.5.3. Transjugular Intrahepatic Portosystemic Shunt (TIPS)
TIPS are a side-to-side portocaval shunts used for the treatment of the
complications of portal hypertension. They have similarities to both the total
and the partial surgical shunts (Grace 1993). Transjugular intrahepatic
portosystemic stent-shunts involve the establishment of a portosystemic shunt
by the transjugular insertion of an expandable metallic stent between the hepatic
45
and portal veins entirely within the liver parenchyma. TIPS effectively decrease
portal hypertension by connecting the portal and hepatic vein with an
expandable metal stent (Svoboda et al. 1997). In the last 20 years this shunt has
become the favoured treatment of many complications of portal hypertension.
TIPS were first described by Rösch in 1968, and Calapinto developed this
method as a technique in humans in the 1980’s (Mid-America Interventional
Radiological Society). Shunt failure initially diminished the potential success of
TIPS, but with advances in biomaterials and pharmacy, they became a focus for
intensive research.
Where: 1.Outflow - hepatic vein 2. Inflow – portal vein 3. TIPS And the guiding thread in black The TIPS insertion is described below
Figure 2.5.3 TIPS placements principle (Rőssle et al. 1994)
2.5.3.1. Nature of TIPS – surgical procedures
TIPS are performed in many hospitals around the World. The shunt is placed in
a non-operative way, while the patient is under local anesthesia. Usually a
guided catheter is positioned in the right or middle hepatic vein, and a needle is
used to make a puncture in the hepatic vein wall and into an intrahepatic branch
of the portal vein (Mid-America Interventional Radiological Society). Then a
guide wire is introduced and connection is achieved via balloon dilatation of the
parenchymal tract. The last step for establishing the shunt is the implantation of
46
metallic vascular stents and their extension onto the two veins, followed by
redilation of the shunt with an angioplasty balloon (appropriate to the patient
size) for reduction of portal vein pressure. TIPS are an effective bridge to liver
transplantation. Functioning as a side-to-side shunt, some of the complications
of TIPS procedure are hepatic encephalopathy and occasional liver failure. TIPS
are not recommended for preoperative portal decompression solely to facilitate
liver transplantation (Brown 1997; Rosado and Kamath 2003). Usually TIPS
have a diameter less than 10 mm, but this can vary depending on the specific
requirements.
The most common complication during TIPS placement is puncture through the
liver capsule, which, if not resealed quickly, can have a fatal outcome for the
patient (Mid-America Interventional Radiological Society). As the catheter is
most commonly introduced via the right jugular vein, it traverses the right
atrium, and thus can cause cardiac rhythm disturbances during the operation.
There is also another possible problem with the catheter – it can buckle in the
atrium (due to the initial placement of the balloon through the parenchyma) and
prolapse into the right ventricle, thus can produce ectopy, ventricular
tachycardia or ventricular fibrillation (Mid-America Interventional Radiological
Society). Nevertheless, the technical success of placing TIPS is above 90%
(Grace et al. 1996).
TIPS are reported as effective in the treatment of recurrent bleeding due to
variceal hemorrhage, refractory ascites (although not well proven), hepatic
hydrothorax, Budd-Chiari syndrome, treatment of hepatorenal failure,
hepatopulmonary syndrome, veno-occlusive disease and bleeding from portal
hypertensive gastropathy (Boyer 2003; Brown 1997; Rosado and Kamath
2003).
2.5.3.2. Complications of portal hypertension treated with TIPS –
advantages and disadvantages of their use
Some studies report on the successful management of portal hypertension with
TIPS placement for long-term results. Although it is commonly agreed that
TIPS has been a major advance in the treatment of portal hypertension, some
cautiousness should apply due to the mortality rate and other complications.
47
TIPS have been reported to correct portal hypertension, thus to be more
appropriate as a treatment for variceal bleeding than endoscopic therapy
(Collins and Sarfeh 1998). The two methods have similar survival rates,
although encephalopathy is 30% higher after TIPS (Collins and Sarfeh 1998).
Other reports show limited success for portal hypertension decompression with
TIPS due to early thrombosis (12%), stenosis (41%) and high re-bleeding rate
(Becker 1996). In patients with good liver function elective operation might be
more beneficial than TIPS (Becker 1996), and the latter is more beneficial for
patients with poor liver function, active bleeding or liver transplant candidates.
In patients with refractory ascites (study with 65 patients (Thuluvath et al.
2003)) TIPS are associated with unpredictable and high rates of mortality and
morbidity.
In some cases although the reduction in the portosystemic pressure gradient was
significant (58% - from 24±6 mm Hg to 10±4 mm Hg) 13 out of 29 patients had
shunt insufficiency (Rőssle et al. 2000). As most of those patients had alcoholic
liver disease it is not certain that similar results will be observed in non-
alcoholic patients (Lake 2000). Some studies suggest reduction in portal vein
pressure by 30% and of portosystemic pressure gradient of more than 50% after
TIPS placement (Hidajat et al. 2000). Those percentages are smaller before and
after TIPS revisions.
In its severe form, the portal vein is used as an outflow tract for the arterial
hepatic perfusion (Zent et al. 2001). A portosystemic side-to-side shunt, e.g.
TIPS, may facilitate portal outflow thus increasing hepatic (i.e. arterial)
perfusion.
Surgical side-to-side shunts and TIPS are used for treatment of Budd-Chiari
syndrome (occlusion of central hepatic veins) as they facilitate portal outflow
and thus increase hepatic (arterial) blood supply and improve hepatic function
(Zent et al. 2001).
The limited number of controlled trials on the comparison between TIPS and
other forms of therapy for portal hypertension (Boyer 2003) does not provide
sufficient confidence to conclude that TIPS is the best, or even a better solution.
48
Contradictions in results in different case studies may be due to the individual
patients and not to the procedure as such.
TIPS are a common and widely used procedure. Some of the problems with
TIPS are reported in different case studies. Details are given in appendix 6 of
this thesis. Below are some examples of disadvantages of TIPS procedure:
Mortality rate at 30 days of 10% (12.5% reported in (Patel et al. 2001), and
22% in (Dagenais et al. 1994); over 60% in (Ochs et al. 1995) where out of 50
patients 2 died in the hospital and 29 during follow-up) and 15% re-bleeding
caused by thrombosis of the shunt (Svoboda et al. 1997); shunt stenosis and
occlusion in 33 % of patients (Rőssle et al. 1994); some authors even report
occlusion or stenosis of the shunt by neo-intimal hyperplasia narrowing of the
lumen in 20-80% of all patients during the first 6-12 months after the procedure
(Svoboda et al. 1997).
In long-term follow up TIPS have been shown to be associated with high rates
of shunt stenosis and thrombosis (Becker and Reed 1996; Lind et al. 1994).
Shunt dysfunction and hepatic encephalopathy are the major limits for the
success of TIPS (Grace et al. 1996; Rosado and Kamath 2003). Regular and
long-term surveillance of the shunt is required to prevent complications of
portal hypertension due to shunt stenosis. In many cases multiple re-
interventions are required due to shunt stenoses (Pozler et al. 2003).
When rigid or semi-rigid stents are implanted, due to the change in compliance
in the area where the stent joins the vessel, occlusion of the treated vessel can
occur (Puel et al. 1988; Wright et al. 1985).
When compared to endoscopic variceal ligation for re-bleeding prevention in
cirrhotic patients, TIPS do not show improved survival rate two years after
shunt placement in a randomised study (Pomier-Layrargues et al. 2001) in
patients with moderate to severe liver failure.
There are certain complications associated with the placement of TIPS itself,
including liver capsule perforation, intraperitoneal hemorrhage, portal vein
thrombosis and renal failure. The main two problems after the insertion of the
shunt are hepatic encephalopathy (in 20-30 % of patients) and stent occlusion or
stenosis (in 15-50%) (LaBerge et al. 1993; Lai 1997; Nazarian et al. 1994;
Rőssle et al. 1994). There is always the risk, although small of stent dislocation
49
following TIPS placement. This procedure, although relatively simple, requires
precise guidance by a radiologist and X-ray tracing, thus exposure to X-rays for
an extended period of time (American Liver Foundation 2004; Society
Interventional Radiology 2004).
In a recent report (Mancuso et al. 2003) investigating patients with Budd-Chiari
syndrome treated with TIPS – 8 of the 15 patients had hepatic failure.
Higher rates of stent occlusion and early re-bleeding are seen after emergency
TIPS placement compared to elective TIPS (Gerbes et al. 1998; Kofidis et al.
2003).
Due to the high occlusion rate and the possibility of blockages, shunt patients
are required to be under intensive medical supervision for a long period of time
(Malagó et al. 1998). The British Liver Trust is also mentioning the temporary
character of TIPS procedure (BritishLiverTrust) in terms of benefits, and the
need for long-term monitoring and evaluation. TIPS can be effectively used as
bridging treatment to liver transplantation in end-stage liver disease (Collins
and Sarfeh 1995).
An advantage of using TIPS is the ability to enlarge the diameter of the shunt
after placement via a catheter procedure (Zemel et al. 1991) and as a major
advantage in TIPSS (TIPS-stent) (Malagó et al. 1998; Rőssle et al. 1994). This
can result in more gradual portal pressure decompression, and thus might help
avoid problems associated with sudden pressure decrease in the portal vein. In
addition to changing stent diameter, the reduction of portal vein pressure can be
influenced by the number of stents and their total length (Hidajat et al. 2000).
TIPS do not require general anaesthetic and has been shown to reduce ascites
(Kofidis et al. 2003; Lake 2000). The transjugular shunt reduces the porto-
systemic pressure gradient and the arterial resistive index, and increases the
stagnant portal vein flow to normal indicating an increase in the arterial
perfusion of the liver (Zenz et al. 2001). In some cases the reported reduction of
the pressure gradient is over 200% (from 36 to 11Hgmm (Zemel et al. 1991)).
Reduction of the porto-systemic pressure gradient from 24.3mmHg to
9.3mmHg (around 2.6 times) is an average decrease as reported in the literature
(Patel et al. 2001).
50
The direct mortality from TIPS complications is as low as 5%, but as mentioned
before, the approximately 25% of patients developing encephalopathy is a
drawback (Kofidis et al. 2003).
The success of TIPS depends to a great extent to the skill of the physician
performing the procedure; hence the results a less experienced surgeon achieves
might not be as good (Lake 2000).
TIPS placement is quite expensive, with costs sometimes above 50 000USD,
not taking into account stent revision if needed (Lake 2000). This is another
reason why TIPS should be used only after other treatment methods have failed.
2.6. Other treatments and methods for overcoming portal
hypertension The two most popular techniques for overcoming portal hypertension and
treating late stage liver diseases are shunts and liver transplantation. Liver
transplantation is the best treatment for end stage liver disease (Collins and
Sarfeh 1995). Because of the shortage in donor organs, alternatives such as
partial liver transplantation have been developed.
Approximately 25% of all patients receiving transplantation have variceal
bleeding as a complication of their end-stage disease (Grace et al. 1996).
Here other methods for either treatment or extending the life of the patient are
discussed in some detail. The list is not explicit as new technologies emerge
daily, and some techniques are still not well tested.
2.6.1. Mechanical devices
So far no mechanical devices capable of performing all functions of any given
human organ have been developed and therefore these can only be used as a
temporary measure until a more permanent solution can be employed. In the last
decade there have been many researchers advocating the benefits of
micropumps for liver perfusion. The rationale behind trying to create a
mechanical device is to increase portal flow and so reduce portal venous
pressure. Those promising results determined the direction of our initial
research. Firstly, we looked into the types of pumps for possible implantation in
51
human. Then we looked at the materials from which a pump could be made.
Lastly, we investigated the possible ways to drive, monitor and maintain an
implanted mechanical device. Below are some of the ideas we found most
promising, but the list is far from being exclusive. At the end of this section the
reasons for abandoning this as an approach to overcome portal hypertension in
cirrhotic patients are given.
The following is a brief discussion of the types of pumps and micropumps
developed for improving liver perfusion, highlighting their advantages and
disadvantages. These devices could be useful for the short-term, but are not a
good option for long-term treatment.
Types of pumps:
• In 1991 a study (Habib et al. 1991) with a ‘Portac’ balloon pump in pigs
with portal hypertension reduced splanchnic portal pressure and
increased portal flow. Apart from the promising results the researchers
have outlined a potential problem, which does not appear to have been
addressed by previous studies using mechanical devices. That problem
is the possibility that increased portal flow might lead to increased
intrahepatic portal-venous shunting rather than increased sinusoidal
circulation, which could lead to liver failure. Another common problem
is the drive of such a pump. Here the authors have used compressed air,
claiming that it could be replaced by fluid or electromagnetic force. As
this was an early study no means have been employed to address
biocompatibility (although recognised as a problem) and long-term
patency.
• Electric driven impeller pump fixed to the vessel wall for preventing
backflow (Jiao et al. 2000; Marseille et al. 1988). In a pig model, even
though the pump performed really well (~ 50% increase in portal
inflow), thrombus formation could be observed around the impeller.
Given that thrombi developed rapidly (the experiments lasted around 2
hours), this pump may be suitable only for short-term applications,
perhaps complementing other surgical assist devices.
• Totally Implantable Assist Device (Huff 1997) - although its application
is quite different from the one of liver perfusion, is a long-term, small,
52
totally implantable, pulsatile pump with no external venting that
includes an integral frictionless hydrodynamic bearing to significantly
reduce long-term wear, hemolysis and thrombosis. The device has a
single rigid moving part (a hollow piston) that works in conjunction with
two FDA approved mechanical heart valves (gate blood flow in a single
direction which functions as safety check valves for fail-safe operation),
providing a pulsatile assist for an ailing ventricle. For this device a novel
multi-motion motor was developed, that rotates and translates the piston
simultaneously for pumping action and to maintain the hydrodynamic
blood bearing. The techniques used in the development of the vascular
assist device (VAD) mechanism are 3-D modelling, Finite Element
Analysis and Control System Simulation tool sets for optimal
performance. So far this device is still under development, but is
expected to be on the market soon. It could be a good starting point for
the future development of a micropump for liver perfusion.
Micropumps are used for a number of applications and have a different work
principle and drive. The only thing in common is the micro-size of the devices.
• Piezoelectric micropumps use a piezoelectric disk to drive the device
(piezoelectric materials, which act very well – operate with high force
and speed, and return to a neutral position when un-powered). This is
one of the most used actuator types in recent years (Bardell et al. 1997;
Dept of Mechanical Engineering and Dept of Electrical Engineering;
Ederer et al. 1990-2000 multiple; Matsumoto et al. 1999). The motion
arises from dimensional changes generated in certain crystalline
materials when subjected to an electric field or to an electric charge.
Piezoelectric materials respond very quickly to changes in voltages
(materials of this type are SiO2; lead zirconate titanate (PZT); lithium
niobate; polyvinylidene fluoride (PVDF)).
With this drive there are many different designs of pumps – membrane,
diaphragm, valve, chamber, passive ball valve, nozzle chamber or
combinations of these.
Micropumps with no moving valve parts (Galambos 1979) are driven by
a piezoelectric disk bound to the pump membrane. They are positive
53
displacement pumps, operating at low Reynolds numbers and operate at
750μl/min with maximum heads of 4.75m of water. The development of
a linear system model gives the ability to optimise the pump
performance. Other work has been carried out with similar pumps
(Bardell et al. 1997; Foster et al. 1995; Jang et al. 1999; Jang et al.
2000; Micro Infusion System 2000). The driving element, 5mm diameter
piezoelectric disk is centred over the pump chamber and bonded to the
outer site of the cover plate. The valve operates solely by the differential
pressure characteristics in each flow direction, which are caused by the
flow through it. The pump has higher volumetric efficiency (diodicity)
and easily accommodates 20μm diameter particles
• Thermal driven actuators (thermo-mechanical, phase change and shape
memory require cooling, not suitable for implantation in the human
body) are, so far, used for fluid pumping in applications not involving
implanted systems (Grosjean and Tai 1999; Matsumoto et al. 1999).
• Electromagnetic and electrostatic actuators (fields will arise and
disappear rapidly). The IMALP (Implantable Microsystem for
Augmented Liver Perfusion) project is an example of this type of
actuator for implantation in the human body (Marseille et al. 1988;
Versweyveld 1997) (and electrohydraulic - Yambe et al. 2005).
• A centrifugal pump which increases the negative pressure and is placed
in the venous line (vena cava) was used and studied in Lausanne,
Switzerland to optimise the pump driven venous return for minimally
invasive open heart surgery (Tevaearai et al. 1999). Centrifugal pumps
have been used as heart assist devices for many years, possess good
characteristics and are well studied. The rotary blood pump was
developed 20 years ago after the clinical demonstration of the non-
pulsatile flow in 1949. The non-pulsatile rotary pump is a very useful
ventricular assisted device. Recent research indicated that the diseased
heart could recover if allowed to rest for at least one year, and this is
where the rotary pump could provide a temporary solution. Other
advantages of the blood pumps are that they are simple in design (no
valve), of small size, efficient, inexpensive, can operate at high rpm,
54
have no on/off system and do not introduce any psychological
abnormality. At present a team led by Dr. Nose has formulated a
strategy to develop a totally permanent implantable rotary pump –
expected to be available soon (Nose et al. 2000). Currently available
products based on this type of pump are: 2 day pump (eg the
Nikkiso/Fairway pump) with application in cardiopulmonary bypass; 2
week pump (eg the Kyocera Gyro C1E3 pump) with application in
ECMA (Extra Corporeal Membrane Oxygenation) and PCPS
(Percutaneous Cardio-pulmonary Support); Long term pump (eg. the
DeBakey LVAD) with application as a bridge to a recovery device.
Centrifugal blood pumps for medium-term implantation (under 6
months) showed thrombus formation and blood kinking in vivo tests (in
sheep), although performed well on other counts (Goldstein et al. 1992).
• The following two types of devices are used for fluids, but are not
suitable for blood –electro-rheological actuator creates a charge in the
flow, and electro-hydrodynamic actuator needs the fluid to be polarised.
• Catheter pumps are used worldwide. Their applications in drug delivery
and as assist devices are well explored. One of the products on the
market in this group is the P.A.S. PORT® Implantable Peripheral Access
System, developed by SIMS Deltec, Inc., St. Paul, MN55112 U.S.A.
used as long-term central venous access device for delivery of
chemotherapeutic drugs, antibiotics, pain medications, nutritional
solutions, and other intravenous therapies. It is designed to allow a less
traumatic implantation procedure and to be placed in the arm. It is
preferred cosmetically by many patients, is cost-effective, provides
convenient access without having to undress, and the septum access is
designed for needle retention and stability of the fluids. The portal is
made from titanium, has a height of 10mm, weight of 5.6g, and the size
of the base is 26.7×16.5mm. The catheter is made from PolyFlow
Polyurethane, has an inside diameter of 1mm, an outside diameter of
1.9mm, and the length is 76cm, which provides access to places remote
from the arm implanted system. Another catheter pump is the enabler
circulatory support system, which expels blood from the left or right
55
ventricular cavity and provides pulsatile flow in the ascending aorta or
pulmonary artery (Nishimura et al. 1999). It is driven by a bedside
installed pulsatile driving console. The system contributes to
stabilization of the circulation in the presence of cardiac unloading. In
heart failure it actually supports the circulation by increasing cardiac
output and perfusion pressure. As a catheter pump the SubQ pump,
designed by Kent Scientific Corporation, CT, U.S.A. is a good example
of an implantable continuous infusion pump, which consists of a
hydromer catheter, has a silicone disk spring that contacts at a constant
rate forcing the infusion liquid out the flow restrictor, and is filled
through the top with serum. So far it has been used in small animals and
has a low continuous flow rate. The Totally Implantable Drug Delivery
System (TI-DDS), developed by Micro Infusions Systeme GmbH,
consists of a pump, port, drug reservoir and catheter. With this device
the patient can release a preset quantity of 5μl or 10μl of the drug by
finger pressure, as required. It is entirely passive, does not possess a
battery or electronic components, can be anchored under the skin, on the
surfaces of a bone or integrated into tissue. A flexible catheter leads
from the TI-DDS to the destination of the drug. It has good
biocompatibility and the materials used are pure titanium and pure
silicon.
• Continuous-flow blood pumps were developed very early by the
pioneers of the heart-lung machines used for cardiopulmonary bypass. A
review of the history of continuous-flow blood pumps was undertaken
by Don B. Olsen (2000) and described the historical development of this
device starting from DeBakey, Gibbons and Wesolowski. The
importance of plasma-free haemoglobin (Ottenberg 1938) is mentioned
as pointing to the meaning of the critical 150mg% level of the plasma-
free haemoglobin and its exit through the urine – the conjugation in the
liver to haptoglobin, which is cleaned through the reticuloendothelial
system to levels up to 150 mg% (reported by Latham). A miniaturized
axial pump mounted on the tip of the catheter that could be passed from
the femoral artery into the left ventricular chamber as an assist device
56
was described by Wampler et al. (1988). The impeller on this pump
rotated between 25,000 and 35,000 rpm, to yield an output of 3.5L per
minute. The red cell destruction was very low, it produces 10g of free
haemoglobin per day at 3 L/min of flow (one blood cell for every 56,000
red blood cells that passed through the pump would have to lyse
according to the calculation Dr. Wampler made regarding the
Hemopump). The DeBakey VAD is a very small axial pump to be used
as a left ventricular assist device that has powerful magnets set in the
tips of the inducer impeller vanes that minimize the air gap between the
rotor and the stator of the driving brushless-DC motor. Other pumps
with similar size and configuration have been developed by other
groups.
For example Medtronic Bio-Medicus centrifugal pump as reported by
Clark et al. (1996) is shown on the figure 2.6.1.below.
Figure 2.6.1.1.Medtronic pump (Clark et al. 1996)
57
A magnetic suspended centrifugal pump is under development in the
Utah Artificial Heart Institute. Dr. Antaki and the Pittsburgh group have
developed the Streamliner axial flow pump with the impeller suspended
in a magnetic bearing (Bolognesi et al. 2001). Don B. Olsten (2000)
proposed classifications for the blood pumps designed to be used as total
artificial hearts or cardiac replacement devices, as well as a variety of
ventricular assist devices. He divides them into 3 generations: First
generation pumps are the positive displacement or pulsatile pumps (the
CardioWest total artificial heart, the Thoratec biventricular assist
devices, the TCI HeartMate I VAD, the Novacor VAD, the HeartSaver
by Worldheart, and the Pierce Lion Heart, as well as others that are
under development); Second generation are the rotary pumps with
contact bearings and/or seals (the MicroMed DeBakey VAD, the Jarvik
2000 VAD, and the Nimbus-TCI HeartMate II VAD, as well as others
under development) and Third generation pump are rotary pumps with
only magnetic bearings, or rotary pumps without mechanical or touching
bearings (the Terumo VAD, the MedQuest Heartquest VAD, the
University of Pittsburgh Streamliner VAD, and others).
• Vibrating flow pump used as a ventricular assist device, developed for
increasing the blood flow to the brain, where the flow oscillates at
frequency 10-50Hz due to its central tube has shown by Yambe et al.
(2000). The pump is shown on the figure below, and has one jellyfish
valve, four coins and four permanent magnets providing the shake of the
vibrating centre tube.
Figure 2.6.1.2. VFP pump designed by Yambe et al. (2000)
58
This totally implantable vibrating flow pump for left heart assist device (pump
outside chest cavity), where oscillating flow is achieved using magnets and
coins to shake the central tube. Jellyfish valve is mounted at the outside of the
central tube.
Miniaturisation of the pumps is a major benefit for all types of machines, but
the introduction of these small devices into patients to assist in the treatment
and monitor progress for different diseases is a major hurdle. As said earlier our
initial project was to develop and optimise such a pump for blood supply to the
liver in cirrhosis. The device above could be used as a starting point in the
future development of such a pump.
So far there is just one group, developing a pump for this application. This is the
joint group of companies and Universities from Belgium, Germany and U.K.
working on the IMALP project under the ESPRIT 4 program. Their project
presents an electric brushless micromotor driven micropump implanted in the
portal vein, with a flow of 2L/min and backpressure of 50mmHg. The pump
consists of an impeller, which has a rotational speed of 24,000rpm. Its diameter
is 7mm and length 32mm. The pump and the motor are physically separated and
the torque is transmitted by a magnetic coupling. Their goal is for a long-term
working micropump. This project relates to the first one described below in the
drive and monitoring section.
Materials for micropumps:
The materials from which the micropump will be manufactured are as important
as the design of the device. Recently, many studies have been conducted on
biocompatibility of materials and new materials have been developed.
The outcome of coating of cardiopulmonary bypass devices is reduced clotting
and significantly improved the biocompatibility of artificial surfaces exposed to
blood (Tevaearai et al. 1999). The process of Trillium™ coating involves
polyethylene oxide, sulphonate groups and heparin. Coating may be one
solution to prevent adverse effects induced by contact of blood elements with
foreign surfaces.
As an example, catheter thrombogenicity has been studied by A. L. Bailly et al.
(Bailly 1999) in 31 patients using a 50 cm catheter test sampler. The tested
59
catheters were mainly of polyethylene (PE), Pebax® or polyamide (PA)
sterilised and ready for clinical use. The results showed that the most
thrombogenic material was the smoothest and that there was no correlation
between surface chemical composition and thrombogenesis. However, catheters
that were based on PE appeared less thrombogenic than PA catheters in their
study. Some positive results in attempts to calculate the amount of adherent
thrombus to a rotary pump (paracorporeal left ventricular assist device in
calves) from pump flow rate, motor speed, activated clotting time and pumping
days were achieved by Nakata et al. (2000), although during the trial the study
subjects had to have continuous heparin admission to prevent faster clotting.
Research into the variety of materials, combinations of materials, coating, and
development of new materials is gaining momentum, although so far no perfect
material has been presented. Apart from the biological requirements, such as
low thrombogenesis, biocompatibility and low blood cell damage, a number of
mechanical properties need to be observed (stress, strain, velocity and
compatibility between the materials used in the device). And once all those
questions are answered, the next one to address is how to drive, maintain and
modify the already implantable pump. Some short-term solutions have been
found, mainly in the use of batteries, but not one can provide reliable, long-term
support.
Drive and monitoring:
This is a crucial stage in the development of implantable mechanical devices.
Answers to the questions such as: how to provide power for a long time; how to
eliminate the need for the device to be connected to extracorporeal machinery;
how to monitor and how to carry out maintenance and repair work are required.
In relation to the Belgium/Germany/U.K IMALP project, the Department of
Computing & Electrical Engineering at Heriot-Watt University, Scotland, the
Micro Engineering Group developed a novel project – research into energy
transfer for means of powering implantable systems. Their focus was in
developing a microwave antenna link that was capable of both power and data
transfer through subcutaneous tissue in a safe manner. This microwave link
could provide the necessary bandwidth required by the telemetric
60
communications for the control and sending of signals between the implantable
system power supply and also the external monitoring system.
Guidnant Ltd, Belgium, is involved in the same IMALP project, based on
powering the pump by a lithium rechargeable battery and monitoring using
microwaves. This project is residing under the ESPRIT program (Marseille et
al. 1988; Versweyveld 1997) and Guidnant is part of the same project.
Other power alternatives involve optical cables, ultrasound and electromagnetic
waves. With the predicted advances in all areas of science and technology in the
next decade it may be possible to have all the above components developed into
an implantable micropump ready for in vivo testing.
Although micropumps are a solution for overcoming portal hypertension in
patients with cirrhosis there are some problems with this approach, as outlined
below:
• The materials from which devices could be built have to comply with a
variety of biological and mechanical requirements. So far there is no one
material that could comply with all of them.
• The design of the pump has to cause minimal damage to blood cells and
vessel walls, cause no thrombus formation, enable long-term function
and be able to be manufactured easily and inexpensively.
• The drive, monitoring and maintenance of the system needs to be easy,
durable, quick, not requiring hospitalisation or constant medical
observation, and if extracorporeal, not cause infections due to skin and
deeper tissue penetration and damage.
• And finally, even though a pump would increase liver perfusion and
decrease portal venous pressure, it will not help prevent shunting from
collaterals, nor is there evidence it could improve or maintain hepatocyte
function. There are studies suggesting that intrahepatic shunting and
pressure might build with the increase in portal perfusion. If the pump
could slowly increase its speed, change the flow if desired, and be less
invasive to the vessel wall and damaging to blood cells it could provide
a possible solution for overcoming portal hypertension in patients with
cirrhosis.
61
2.6.2. Bioartificial Liver
The worldwide shortage of liver donors, the growing number of people
requiring liver transplantation and the success of hemodialysis for patients with
non-functioning kidneys predetermined the development of artificial
extracorporeal liver devices. It needs to be noted that these devices provide only
temporary support to the patients.
Currently, a number of bioartificial liver devices have been proposed and
studied. In this part of the thesis only a brief overview of these is given, with
more details provided on the ones that, in my opinion, have a greater impact on
the emerging technologies.
Bioartificial liver devices incorporate living cells (predominantly hepatocytes)
and are expected to perform most of the functions of the living organ. To date,
all devices of this type are extracorporal, i.e. not implantable.
There are two types of bioreactors currently being developed. The first type
includes bioreactors as devices for growing tissues, predominantly blood
vessels and heart valves, in vitro for study of flow behaviour or for
implantations. That type will be discussed in chapter 4 of this thesis in detail.
The other bioreactors are designed as hosts for tissues, for use as long-term
extracorporeal devices to assist or replace a malfunctioning body organ. These
types of bioreactors are discussed here and some examples of promising work
and future challenges are presented.
Figure 2.6.2.1. Extracorporeal BAL circuit schematic representation (Chaib et al. 2005; Chamuleau 2002)
62
This circuit separates the plasma from the blood and provides heating of the
plasma while passing through the circuit.
In general, bioreactors need to comply with a wide range of requirements. From
an engineering point of view the scale, material and assembly of the bioreactor
are amongst the most important issues. On the other hand, from a medical and
biological point of view issues like sterility, growth conditions, operational
temperature and the inner surface of the device are important. The task of
creating a bioreactor is therefore very complex and requires a cross-disciplinary
team of experts. Thus, up to date these are expensive, unique and not widely
available as life-saving devices. Niklason et al. (Mitchell and Niklason 2003;
Niklason et al. 2001; Niklason et al. 1999) have reviewed this area as well as
describing their own work on the topic of bioreactors and tissue grafting. Some
of their work is discussed in other parts of this thesis. They urge a better
understanding of all aspects of the process for developing and operating a
bioreactor for the purpose of tissue engineering.
2.6.2.1. Bioreactors – types, principles and some problems
The following list of currently available or under development bioreactors
includes examples only and is not exhaustive. The one thing these bioreactors
have in common is that they all are extracorporeal liver support methods.
a) CellModule multicompartment bioreactor uses primary human liver
cells with small capillaries with interwoven membranes. Additions to
this system are the DetoxModule for the removal of albumin-bound
toxins and the DialysisModule for veno-venous hemofiltration (Sauer et
al. 2002). Cells attached to a nonwoven polyester fabric acting as a
matrix allowing for direct contact between plasma and cells are
commonly used (Chamuleau 2002; Chamuleau 2003; Naruse et al. 2001.
b) Porcine hepatocytes extracorporeal bioartificial liver (Mischiati et al.
2003) is used to overcome the shortage of available human livers,
although there is the problem with immune intolerance and diseases
transmitted via animal cells or tissues. An important challenge relates to
storage and delivery of these ‘ready-to-use’ bioreactors, with
suggestions refrigeration might be suitable. A flat-plate bioartificial liver
63
device with an internal membrane oxygenator and porcine hepatocytes
has shown promising results in rat studies (Shito et al. 2003). So far
there has not been evidence of any immunological problems or
endogenous disease transmission, but at the same time, improvement in
survival rates has also not been definitely established (Chamuleau
2002).
c) Bioartificial liver support system using genetically modified hepatocytes
(Chen et al. 1997; Kawashita et al. 2000; Neuzil et al. 1993; Shatford et
al. 1992) employs collagen coated microcarriers to allow larger number
of hepatocytes to be cultured in a smaller area (like extra fiber space).
Those methods have been successfully tested in large animals with the
use of different genetically modified cells.
d) Liver cells encapsulated in gels are under investigation for both direct
cell transplantation in the peritoneal cavity and for use in extracorporeal
BAL (Chamuleau 2002).
2.6.2.2. Hollow fibre bioartificial liver
The blood perfused hollow fibre cartridge bioartificial liver consists of
hepatocytes seeded in the extra-capillary space. The supply of oxygen to the
cultured cells is very important as hepatocytes consume high quantities of
oxygen to facilitate their metabolic functions.
Those BAL have been used successfully as artificial kidneys, where either the
blood or only the plasma passes through the capillary system of semipermeable
membranes on the other side of which the active cells are attached (Chamuleau
2002).
A number of requirements of the hollow fibre system (Hay et al. 2000),
including the media and cell density need to be taken into account and carefully
studied before further progress can be achieved. In terms of modelling P.D. Hay
et al. (2000) have presented a more complex model than the previously studied
one by introducing transmembrane convective flux, and the practical use of this
device is still to be evaluated. Their model predicts the inappropriateness of the
use of the hollow fibre cartridge as a bioartificial liver.
64
2.6.2.3. Fluidised Bed Bioartificial Liver
As can be seen from the number of studies involving extracorporeal bioartificial
liver support, the race for creating the most effective design is still underway.
Fluidised Bed Bioartificial Liver featuring a “bioreactor for extracorporeal liver
supply containing alginate beads in a fluidised bed regimen” was proposed in
2000 (Legallais et al.). In that device, given as an example, empty alginate
beads and saline solution served as solid and liquid phases for the first phase of
their experiment with the tendency of these to be replaced later with hepatocyte
and blood plasma. The materials used for the cylinder and the caps are
polycarbonate, nylon and polyethylene, chosen for their biocompatibility and
possibility for autoclaving. This model was developed by Legallais et al. (2000)
for in vivo testing in pigs and therefore calculated for the specific parameters
(flow rate and vessel diameter) of that animal. In vitro studies performed by the
authors using a roller pump, tank, bioreactor and a safety filter, showed the
effectiveness of this method for extracorporeal liver support.
2.6.3. Non-shunt operations
Non-shunt operations include the Sugiura procedure (transthoracic esophageal
transection), transabdominal esophageal transection and Hassab procedure. In
some countries like Japan (Ohashi et al. 1998) those procedures have been
found useful in preventing bleeding from esophageal varices in patients with
idiopathic portal hypertension. Recurrent varices after non-shunt operation have
been reported with a rate of 3.8% within 8 years going up to 8.9% by the the
15th year (Ohashi et al. 1998), and the rate of recurrent bleeding being up to
5.1% within 10 years and reaching 9.8% till the 15th year. In patients with non-
alcoholic cirrhosis (the predominant cases in Bulgaria) the modified Sugiura
procedure has been used with success in both emergency and planned treatment
for patients with hemorrhage from esophageal varices (Merzhanov and
Damianov 1989).
The Modified Sugiura procedure of transabdominal extensive esophagogastric
devascularization with esophageal or gastric-stapled transection has shown a
survival rate of 88% after 5 years of management of patients with non-cirrhotic
65
variceal bleeding even in emergency procedures in patients considered not
suitable for surgical shunts (Mathur et al. 1999). In those operations, patients
with portal and splenic vein thrombosis were not excluded, which gives those
patients which are not considered for most shunt procedures good control of
variceal bleeding in the short and medium terms (Shah et al. 1999).
2.6.4. Sclerotherapy
Sclerotherapy is widely used today as one of the first treatments for portal
hypertension and usually only if it fails are shunt procedures recommended. It
can be used in preventing variceal re-bleeding (Neuhaus and Blumhardt 1991).
2.6.5. Balloon Tamponade
The use of balloon stents was introduced over three decades ago for the
treatment of atherosclerosis. Even though the stent reduces the restenosis of the
blood vessel, the problem with the stenosis of the stent itself still has not been
resolved. Balloon mounted stents are now in use for the coronary artery, still
facing the same problems with long-term efficiency and stent blockages.
There are many reports on the successful use of the balloon tamponade
technique in patients with variceal hemorrhage. There are certain time limits
related to this procedure, with a common use for up to 12 hours but not
exceeding 24 hours. Also reports suggest that half of the patients re-bleed
within 24 hours, and up to 20% of the patients could have fatal complications
after the procedure (Chojkier and Conn 1980). This is a very short-term solution
and so far has not been modified in a way to allow for continuous long-term
use.
2.7. Medical conditions associated with Portal Hypertension Portal hypertension in cirrhosis commonly results in the development of
complications including variceal hemorrhage, ascites, hepatorenal syndrome,
hepatic encephalopathy and spontaneous bacterial peritonitis (Lata 2003).
For most of these conditions transjugular intrahepatic portosystemic shunt
(TIPS) is the last resort treatment after medications and other procedures have
66
failed. The nature, benefits and disadvantages of TIPS procedure were
discussed together with other shunt procedures in part 5 of this chapter.
Extrahepatic portal vein obstruction was shown to progress in a long-term
follow up study (Ogawa et al. 2002) due to portal hypertension. The collaterals
that develop in this condition helped to maintain the hepatic flow and there was
no progression of the intrahepatic portal vein obstruction.
2.7.1. Complications of Liver Transplantation
Chronic rejection in allograft liver transplants leads to graft failure within a
couple of years post-operation. Such rejection is usually preceded by acute
cellular rejection (Petrovic 2003). So far only immunosuppression is employed
to prevent graft loss. As it is discussed in several parts of this thesis there are
many risks associated with the use of immunosuppressive medications. Some
patients develop portal hypertension after receiving a liver transplant. There are
suggestions that in such cases the use of TIPS should be considered, even
though the mortality rate at 30 days is around 25%, and 30% of the recipients
undergo another transplantation (Amesur et al. 1999).
Some of the shunts currently in use violate the hemodynamics in a way that
makes it not advisable to proceed with transplantation. An example of such a
shunt is the small-diameter H-graft (Rhee and Sarfeh 1993).
2.7.1.1. Shunts
Patients receiving liver transplantation often have portosystemic shunts due to
portal hypertension (Nosaka et al. 2003). There is much discussion on the use of
ligation of such shunts during the transplant operation. This process is invasive
and needs to be carried out very carefully to avoid varicose bleeding, but helps
increase the portal blood flow to the liver. Some other complications in partial-
liver transplantation were discussed earlier in this chapter (2.2.3.1.).
2.7.1.2. Stenosis
Amongst the most common causes of morbidity and mortality after liver
transplantation are hepatic artery thrombosis, portal vein thrombosis and
inferior vena cava thrombosis.
67
Stenosis of the coronary artery bypass (or shunt) has been the object of many
studies in recent years. They have shown the effect blockages have on the
outcome of the procedure and the importance of constant monitoring after the
procedure. In the modelling part of this thesis different models of the portal vein
with blockages are presented together with their impact on the blood flow
pattern when compared to non-blocked vessels with the same parameters.
Studies of coronary bypass have shown that restenosis of the bypass
anastomosis together with the release of growth factors contribute to early graft
failure (Armstrong et al. 2000; Collart et al. 2000: Reicher 1998; Song et al.
2000).
Symptomatic portal vein stenosis is not one of the common complications after
liver transplantation (about 1-2% of patients) and is usually successfully treated
with portal vein angioplasty via either the percutaneous or the mesenteric vein,
or, as recently reported, using transjugular intrahepatic access for introduction
of a balloon catheter (Glanemann et al. 2001). In the literature there have been
reports that the majority of cases of such stenosis are in patients requiring
intraoperative reconstruction of the portal vein or who had preoperative portal
vein thrombosis. Stenosis might be related to the discrepancy between the donor
and recipient vessel size. Some of the other contributing factors are decreased
portal blood flow due to spontaneous collateral formations or previous
splenectomy, hyper-coagulation or severe allograft edema.
Stenosis of the portal vein, or the shunt, can cause a wide range of wall shear
stresses (Hinds et al. 2001) and modelling may need to examine axisymmetric
flow, with changes in the flow pattern.
2.7.1.3. Embolization of the portal vein or one of its branches
Before hepatectomy usually embolization of the portal vein, or one of its
branches in case of partial hepatectomy, is performed. In such cases when
arterial embolization is performed before portal vein embolization the effects of
atrophy are significant. Pre-operative embolization reduces the risk of post-
operative hepatic failure after major liver resection (Taraszov et al. 2002).
68
2.7.2. Hepatopulmonary syndrome
Hepatopulmonary syndrome can occur in cirrhotic and non-cirrhotic portal
hypertension. It is suggested that portal hypertension is the predominant
etiopathogenic factor related to hepatopulmonary syndrome (Kaymakoglu et al.
2003).
2.7.3. Variceal bleeding
Portal hypertension is assumed to be amongst the main causes for variceal
hemorrhage due to the increase of portal venous pressure above 12mmHg
(Grace 1996; Lai 1997). The risk of re-bleeding increases in patients with
decompensated cirrhosis amongst other factors. Unfortunately there are studies
showing that up to a third of patients with varices bleed at least once carrying
up to 30% mortality risk at each bleeding (Grace 1990; Lai 1997). Moreover,
there are reports that about 60% of the re-bleeding patiens will die within 1 year
of the last bleeding (Harry and Wendon 2002; Lai 1997). The risk of re-
bleeding has been reported to be as high as between 50 and 80% (Becker and
Reed 1996).
The long-term usefulness of treatments like TIPS still has to be studied in
controlled trials. Some factors, like decompensated cirrhosis, large varices and
hepatocellular carcinoma, increase the risk of recurrent variceal hemorrhage.
Variceal hemorrhage complicates cirrhosis in 50% of patients (Harry and
Wendon 2002). Variceal bleeding results in high morbidity and mortality. The
first treatment is usually endoscopic and pharmacological, but if they fail
balloon tamponade, sclerotherapy, ligation, TIPS or surgery are the only
alternatives. Variceal bleeding has also been reported in a patient after orthotic
liver transplantation, with was combined with portal vein stenosis (Glanemann
et al. 2001).
The rate of growth of varices in patients with cirrhosis is proportional to the
severity of liver disease (Grace et al. 1996).
Portal flow is preserved using distal splenorenal shunts, and the rate of re-
bleeding and encephalopathy are reduced when compared to central shunts.
69
Esophageal varices due to opening of porto-systemic collaterals for portal vein
decompression are a common complication in portal hypertension. The main
risk is bleeding from the esophageal varices as this has a high mortality rate.
Idiopathic portal hypertension (IPH) has different complications depending on
the geographical region where it has been studied. For example, management of
esophageal varices in patients with idiopathic portal hypertension in Japan
commonly involves non-shunt operations, where the prevention of bleeding
from esophageal varices is a priority and the incidence of re-bleeding is very
low (Ohashi et al. 1998).
Surgical interventions during acute variceal bleeding result in up to 70%
mortality (Krahenbuhl et al. 1999). If the patient is waiting for liver
transplantation TIPS is the preferred option, although the long-term outcome of
surgical shunts is much better than of TIPS (Krahenbuhl et al. 1999). Some
authors strongly argue in favour of partial shunts like H-grafts for long-term
effective treatment in patients with variceal hemorrhage due to portal
hypertension (Collins et al. 1994). In those trials in the early 1990s, small-
diameter grafts were gaining popularity due to their good management of
variceal bleeding and low rates of hepatic encephalopathy.
2.7.4. Hepatic hydrothorax
In the absence of cardiopulmonary disease in patients with cirrhosis and portal
hypertension a pleural effusion can develop, called hepatic hydrothorax
(Chamutal et al. 2004). The usual treatment for this disease includes medical
therapy, liver transplantation, or a combination of TIPS and thoracoscopic
repair of defects of the diaphragm (this procedure has a high morbidity and
mortality rate due to the nature of the condition). Only 4-6% of patients with
cirrhosis, and up to 10% with decompensated cirrhosis, have this condition, but
most of those patients do have portal hypertension and require liver
transplantation (Chamutal et al. 2004).
2.7.5. Portal hypertensive gastropathy
Portal hypertensive gastropathy is suggested to be related to portal hypertension
although portal hypertension is not the only factor responsible for this condition
70
(Grace et al. 1996; Merkel et al. 2003). The most common description of the
condition is in relation to the morphological alterations of gastric mucosa in
patients with liver cirrhosis. Some studies suggested that this condition is a
common complication in cirrhotic patients (Primignani et al. 2000). For the
treatment of acutely bleeding portal hypertensive gastropathy, emergency
portacaval shunts seem to be beneficial (Grace et al. 1996).
One of the largest studies involving 373 patients with cirrhosis was the one in
which all patients seen at 7 hospitals during June and July 1992 were included
and followed up with clinical and endoscopic examinations every 6 months for
up to 3 years in Italy (Primignani et al. 2000). This study concluded that portal
hypertensive gastropathy (PHG) was common in patients with cirrhosis, and its
prevalence paralleled the severity of portal hypertension. Gastropathy can
progress from mild to severe and vice versa or even disappear completely.
Bleeding from this lesion is relatively uncommon and rarely severe.
Sclerotherapy of esophageal varices does not seem to influence the natural
history of this condition.
2.7.6. Porto-pulmonary hypertension
Porto-pulmonary hypertension is described as a hemodynamic constellation of
elevated pulmonary arterial pressure, increased pulmonary vascular resistance
and normal pulmonary capillary wedge pressure occurring in patients with
portal hypertension (Sulica et al. 2004).
Studies have shown that pulmonary hypertension of various degrees is
responsible for liver intraoperative and immediate post-transplantation death
due to intractable right ventricular failure (Kaymakoglu et al. 2003; Sulica et al.
2004).
It is widely acceptable practice to have the absence of pulmonary hypertension
as one of the most common pre-requisites, together with good liver function,
when considering patients for the suitability of shunt treatment.
71
2.7.7. Other liver disease conditions
Hepatic encephalopathy may complicate nearly all types of liver diseases. It can
be reversed in some cases, and can lead to death in others. Usually the treatment
of the condition causing the encephalopathy is the only way to deal with this
complication.
Congenital absence of the portal vein with a systemic shunt is a rare
malformation. To date around 30 cases have been reported in the literature
(Appel et al. 2003; Mitchel et al. 2000; Niwa et al. 2002; Northrup et al. 2002)
and the origin of the malformation is still unknown. In the case described by
Appel et al. (2003) and based on the currently available literature, an
assumption could be made that this condition is caused by a thrombotic
occlusion of the extrahepatic portal vein, due to the existence of normal
intrahepatic bile ducts in the liver, as demonstrated by liver biopsy. The
intrahepatic branching of the developing portal vein is a prerequisite for the
formation of intrahepatic bile ducts during embryogenesis and so agenesis of
the portal vein can be excluded as a cause of this disease (Appel et al. 2003).
Distal splenorenal selective shunts have shown the lowest rate of
encephalopathy (3.5%) when compared to TIPS (29%) and total shunts (16%)
(Becker and Reed 1996). The portal vein develops in the 5-10-weeks of embryo
development (Northrup et al. 2002). Absence of portal vein is usually observed
in either absence of joining of splenic vein and superior mesenteric vein, or as
the joint of those two veins to the inferior vena cava instead of entering the liver
(Northrup et al. 2002).
Absence of bifurcation of the portal vein is a rare condition and was first
described by Couinaud C in 1957. Some reports have followed (Cheynel et al.
2001; Couinauld 1993) most in agreement that this abnormality could be found
in 1.5-1.9% of patients undergoing liver surgery. It differs from other similar
conditions in which one of the branches of the portal vein is missing.
Ascites is common in patients with decompensated cirrhosis. It involves fluid
accumulation in the peritoneal cavity, and its presence in cirrhosis has poor
prognosis.
Porto-pulmonary hypertension is another uncommon complication of portal
hypertension.
72
Portal vein aneurysm and arterioportal fistula are rare, and in the case reported
by Nam KJ et al. (2003) turbulent flow was observed. In the case of
arteriovenous fistula, increased portal vein flow may lead to prehepatic portal
hypertension. Portal aneurysms are usually found in the main portal vein
branches. As it is not a very well documented condition the treatment is still to
be established.
Isolated intractable bleeding from anorectal varices is a rare complication of
portal hypertension usually treated with shunting procedures.
A number of hepatic diseases have been associated with portal hypertension
during their progression. Some of them have been briefly noted above, and for
illustration purposes a few others are mentioned here. Conditions like pure
hepatic steatosis, hemochromatosis, Wilson’s disease, sarcoidosis, chronic
hepatitis, acute viral or alcoholic hepatitis, partial nodular transformation and
some types of hepatic tumours, are all in some degree associated with portal
hypertension.
2.8. Vessel blockages and thrombosis Thrombus is a clot formed inside a blood vessel, and the condition of forming
clots is called thrombosis (Miller and Leavell 1972).
Liver cirrhosis is the most common cause of thrombosis of the portal vein
(Bolondi et al. 1990).
Portal vein disease could be caused either by stains at the site (thrombi
formation) or by propagation of portal thrombosis from another site (Tanaka
and Wanless 1998).
Studies show that, on average, every third patient with extrahepatic portal vein
obstruction has portal and splenic vein thrombosis (Shah et al. 2003). This
contributes to predominantly left-sided collateralisation. Large clots in the
portal vein could be spread from the intestines and spleen as they carry blood to
the liver. An increase of pressure in intestinal wall capillaries by more than
15mmHg above normal can cause sudden death of the patient. On the other
hand, when the pressure in the hepatic vein entering the vena cava increases by
3-7mmHg, the build up of ascites in the abdominal cavity becomes a new
problem.
73
The most common cause of portal vein thrombosis is surgical intervention for
reduction of the portal vein pressure. Thrombosis is most common in patients
with idiopathic portal hypertension (about ¼ of cases), followed by
splenectomy (13%) and cirrhosis (only 1.8% of the cases) (Eguchi et al. 1991;
Shinoka 2002)). Portal vein thrombosis is rare in cirrhotic patients without
related neoplasm (Amitrano et al. 2002)
In children, extrahepatic portal vein thrombosis is associated with abnormal
circulating anticoagulants and procoagulants (Mack et al. 2003). In those cases
it is suggested that restoring the portal flow is crucial for maintaining normal
coagulation.
Portal vein thrombosis is a common complication of splenectomy in patients
with splenomegaly (Eguchi et al. 1991), but the underlying factors for the
development of portal vein thrombosis are still being investigated (Senderos et
al. 2001).
In patients with transcatheter arterial chemoembolization-induced bile duct
injuries some reports have shown over 90% of the patients having narrowing or
obliteration of the adjacent intrahepatic portal vein branches (Yu et al. 2001).
In many studies conducted to evaluate different shunt procedures, patients with
portal vein thrombosis have been excluded (Collins 1998: Rőssle et al. 2000).
This is to illustrate the complex character of thrombosis and its implications on
portal hemodynamics. On the other hand, non-shunt operations have been
performed on patients with portal and splenic vein thrombosis (Shah et al.
1999).
The mechanism of thrombogenesis has been described in the literature and in
basic terms occurs as follows (Fung 1993):
a) Endothelial layer is injured and exposes collagen
b) The collagen interacts with the glycoprotein on the platelet membrane
and other factors within the blood plasma
c) The platelet adheres to the endothelium via fibrinogen and other factors
d) Other platelets circulating in the plasma become attracted to the same
place, thus enlarging the aggregate
An even simpler definition for thrombogenesis can be found in (Strackee and
Westerhof 1993), whereby the “damage to the endothelial layer and the
74
subsequent exposure of the sub-endothelial layer to the blood lead to the
adhesion of platelets to the latter, thus the formation of thrombi”.
The formation and stability of a thrombus is mainly due to thrombin, which
generates fibrin from fibrinogen. Thrombosis is a dynamic process, and clot can
be dissolved as easily as formed. There are medications specially designed to
prevent the formation of thrombi, with aspirin being the most commonly used.
Apart from the threat of thrombosis to human life, it is also essential for
preserving it through stopping internal bleeding. The case of portal vein
thrombosis is no different to that of general thrombogenesis in the human
vascular system.
Portal vein thrombosis is an uncommon cause for presinusoidal portal
hypertension, which can be caused by one of three broad mechanisms:
spontaneous thrombosis when thrombosis develops in the absence of
mechanical obstruction; mechanical obstruction caused by vascular injury and
scarring or invasion; or extrinsic constriction by adjacent tumour or
inflammatory process (Uflacker 2003). Although recanalization of the portal
vein is a used and useful technique, it is associated with risk of intimal or
vascular trauma to the portal vein, which then can resolve in recurrent
thrombosis (Uflacker 2003).
Portal vein thrombosis has been reported in approximately 7% of consecutive
patients studied for evaluation of portal hypertension and cirrhosis (Tasu et al.
2002), and in 3.4% in another consecutive study (Ozaki et al. 1988).
Some of the non randomised trials with TIPS addressing the issues of stent
patency, variceal re-bleeding and encephalopathy, showed that, within the first
year of shunt placement, thrombosis or critical stenosis occurred in
approximately 50% of the patients (Collins and Sarfeh 1998; Grace et al. 1996).
This has shown to be a major cause for the TIPS re-bleeding rates of 15-30%
(Grace et al. 1996).
Another alarming study (Rőssle et al. 1994) showed that 25% of the patients
whose stents narrowed after TIPS placement or had stent obstructions (33% of
all 100 patients) developed degenerative brain disease.
75
A resection study of 15 livers (Tanaka and Wanless 1998) shows that severe
portal vein disease is more frequent in patients with prior shunt surgery, and the
authors found no evidence of shunt surgery preventing portal vein thrombosis.
Their study is not conclusive and could have been impacted by the small
number of livers, or there might have been bias in assigning the patients for the
shunt procedure.
Even in patients with idiopathic portal hypertension undergoing non-shunt
operations, 1 out of 46 patients died due to portal thrombosis and 4 due to
hepatic failure (Ohashi et al. 1998). Thrombosis was seen as the cause for
mortality attributed to the operation.
With the advance in medicine and pharmacology, pylephlebitis, a
thrombophlebitis condition involving the portal vein and its intrahepatic
branches, is not as common as it use to be. It manifests as obscuring of the
portal triad architecture and is usually caused by infection.
Thrombosis can affect either the liver outflow by obstructing the hepatic vein,
or the inflow via the portal vein or hepatic artery.
The two most important conditions causing obstruction of the hepatic outflow
are veno-occlusive disease and Budd-Chiari Syndrome. The first was briefly
mentioned above (section 2.2.1.3. of this chapter), and involves mainly the
smaller intrahepatic venules. The second usually has an underlying reason of
either obstruction in the inferior vena cava or in the larger branches of the
hepatic artery leaving the liver.
Vessel blockages have been shown to be common in patients with Budd-Chiari
syndrome (BCS). Tanaka et al. (1998) evaluated 15 resected livers to determine
the distribution of vascular obstruction. The authors noted a correlation between
the presence of portal venous disease and the form of cirrhosis in BCS. They
found portal venous disease in all livers, and the grade of the intimal fibrosis
varying in different sections within the same liver. A smaller percentage of
narrowing in the large portal vein has the same grade as a large percentage in
the medium and small portal veins (for example, the highest grade 3 is assigned
when the narrowing is above 20% in the large PV, and above 75% in the
medium and small portal veins) (Tanaka and Wanless 1998).
76
Even though rare, there have been reports of portal vein thrombosis following
laparoscopic splenectomy (Eguchi et al. 1991). Up to date there are theories of
the etiology and treatment of portal vein thrombosis, but more studies need to
be conducted to be able to determine the most suitable approach in patients with
thrombosis.
In children with normal liver the occurrence of portal vein occlusion is a
common cause of hypertension. In those cases it is expected that the risk of
bleeding will decrease with age, thus usually the treatment is non-invasive.
In patients with obstructive jaundice the possibility of portal vein thrombosis
has to be considered (Lin et al. 1996).
In a report of 600 pediatric liver transplants in 325 patients (Buell et al. 2002)
with late post-transplant portal vein (38 patients) or vena cava stenosis (12
patients) or thrombosis required further treatment. Predominantly, the factor
responsible for this was considered to be the cryo-preserved vein for portal
conduits.
Exclusion criteria for portal vein pulsatility measurements are portal vein
thrombosis and reversed portal vein flow, amongst other criteria (Barakat
2002). Thrombosis is an exclusion factor in treatment and evaluation of portal
hypertension in the most up-to-date studies (Lake 2000).
2.9. Cell Adhesion The aim of this section of the literature review is to review the importance and
methods of cell seeding and adhesion, the variety of scaffolding materials, and
the methods for growing a blood vessel in vitro. This will involve a discussion
on the various available materials and currently used methods, concluding with
the seeding method, scaffold material and conditions for tissue growth selected
for use in future study. This review provides information and suggests possible
steps to be undertaken as the next step following this study. The importance of
pulsatile flow for culturing the new vessel and other current methods for liver
disease treatment are briefly presented at the end of the chapter. Although some
tissue culturing experiments have been carried out during the time of
completing this project, they were neither comprehensive nor part of the main
77
aim of this study. This review however was used in the creation of the
Bioreactor for in vitro tissue culturing.
2.9.1. Terminology
When discussing adhesion in this thesis I am refereeing to cell-adhesion for the
purpose of tissue growth and improvement of biocompatibility of biomaterials.
There is, however, another meaning used in surgery, which I will describe
briefly below. Adhesions are fibrous bands that connect tissue surfaces that are
normally separate. Adhesion formation is a natural consequence of surgery,
resulting when tissue repairs itself following incision, catheterisation, suturing
or other means of trauma. At the places where a surgeon has to cut, handle, or
otherwise manage internal body parts, tissue, which should normally remain
separate, will sometimes become “stuck” together by scar tissue, defined as
adhesions (Dark 2003). The word "adhesion" comes from the Latin "adhaerere"
meaning "to stick to or cling to" (MedicineNet.com 2003). I will not use this
meaning when talking about adhesion in this and the following chapters.
Cell adhesion is affected by a number of factors, which need to be studied and
evaluated in each case study. Examples of such factors are flow patterns,
pulsatile flow, biological activity at the wall, shear stress, pressure and stenosis.
Local hemodynamics affects the spatial distribution of adherent cells via inertial
forces, gravitational forces and stress (Fung 1993; Hacking et al. 1996; Hinds et
al. 2001; Kobashi and Matsuda 2000; Mori 1989).
2.9.1.1. Tissue Engineering
Tissue engineering is a relatively new interdisciplinary field, which applies the
principles of biology and engineering to develop possible substitutes to restore,
maintain or improve the function of tissues or organs. It can be viewed as a
form of therapy, which differs from the standard therapies in that the engineered
tissue or organ becomes integrated within the patient, giving a potentially
permanent and specific cure of the disease state (Chamuleau 2002; Chamuleau
2003; Langer and Vacanti 1993).
Generally, the tissue engineering approaches can be divided in to three groups:
78
• Design and growth of human tissues outside of the body for subsequent
implantation to repair or replace diseased tissues - the most popular
example for this approach would be the skin graft, which has been used
for over 10 years for treatment of burns.
• The implantation of either cell-containing or cell-free devices that
encourage the regeneration of functional tissues relying on the
purification and large-scale production of appropriate signal molecules
(such as growth factors) to assist in tissue regeneration. Novel three-
dimensional polymers have been developed, to which cells attach and
grow to reconstitute tissues. A popular example of this category is the
biomaterial matrix used to promote bone re-growth for periodontal
disease.
• The development of external or internal devices containing human
tissues designed to replace the function of diseased internal tissues - this
approach involves isolating cells from the body (using techniques such
as stem cell therapy), placing them on or within structural matrices, and
implanting the system inside the body or using the system outside the
body. Some examples are repair of bone, muscle, tendon, and cartilage,
cell-lined vascular grafts and artificial livers and include the
extracorporeal liver support (Sauer et al. 2002), BAL (Kawashita et al.
2000; Legallais et al. 2000; Mischiati et al. 2003; Shito et al. 2003),
isolated hepatocytes (Puviani et al. 1998), etc.
Areas of concentrated research efforts include: cell isolation and cell
substitution; tissue-inducing substances; and cell placement on or within
matrices (extracorporeal or for implantation) (Langer and Vacanti 1993).
In cases like tumour removal in the portal vein or its branches a section of the
hepatic vein is used to patch-up the portal vein and restore blood flow through
the liver (Prakash et al. 2003). This is an example of tissue engineering, but is
not an area this thesis has looked into due to the mainly surgical aspect of the
problem.
79
2.9.1.2. Importance of the endothelial cell lining of blood vessels
The importance of endothelial cell lining of blood vessels have been studied for
many decades with emphasis on their biological function, density, orientation
towards the flow and mechanical properties. Most studies have concentrated on
the endothelial cell (EC) inner layer of the aorta and main arteries. In the
arteries the interaction between EC and smooth muscle cells is very important
for the biomechanical behaviour of the vessel and for its reaction to the pulsatile
blood flow. Major studies over the last decade reviewing the intima (inner
layer) of arteries and its relationship with atherosclerosis have been carried out
by the American Heart Associations Committee on Vascular Lesions show the
important role EC have on the overall structure and behaviour of the arteries. It
is worth mentioning that in their first study published in 1992 (Stary et al.), an
explanation on the different vessel layer structure and thickness in regions of bi-
or trifurcations for arteries was given, which can lead us to conclude that there
is a need for studies to determine how veins adapt their structure in such
regions. In the absence of any such data, in this thesis we have assumed unified
vessel structure in all regions of the created model.
Another fact to remember from that study is that under normal conditions the
endothelium does not support the adherence of large numbers of leukocytes,
platelets or the formation of thrombi. There are many reasons for this but the
work presented here will only deal with the most important. A small review on
some of those factors is made further in this chapter, but its purpose is to
explain why endothelial lining of a graft could be beneficial to the outcome of
the shunt procedure. Investigations on the topic will need to be a subject of
further studies.
The thin (0.1-0.5μm) single layer of endothelial cells on the blood vessel wall
contact surface has predominantly two functions: preventing the adhesion of
particles and cells to the wall, and selective permeability of substances.
Although this layer has no effect on the elastic properties of the blood vessel it
responds to physical (such as shear stress and pulsatility) and to chemical
stimuli (Mori 1989, Chapter 9). The “endothelium interacts with the basement
membrane in one of the following ways” – adhesion, spreading, migration and
proliferation (Mori 1989, Chapter 10).
80
Endothelial cells change their shape, size, orientation and intercellular contacts
depending on specific conditions, i.e. they elongate and re-orient their
cytoskeletons in the direction of the flow as a normal response to prolonged
shear stress (Bruden et al. 2001; Nerem et al. 1998; Shiomi et al. 2000).
2.9.2. History of cell adhesion and cell seeding
Cell adhesion and tissue culture are techniques with a very wide range of
applications, such as bone, artery and organ tissue growth, tissue reconstruction
and regeneration and the grafting of prostheses to improve their qualities.
Although work on endothelial cell seeding of vascular prostheses was first
published in 1978, no clinical breakthrough had been achieved before the early
1990’s (Zilla 1991). Clinical data on single-staged procedures using freshly
harvested autologous venous or microvascular endothelial cells (to the graft at
time of implantation) are scarce and controversial. The alternative approach –
the application of culture techniques – has the disadvantage of being restricted
to major centres. Moreover, this in vitro endothelialization is confined to
elective cases because of the delay caused by cell cultivation. Nevertheless,
initial clinical trials with this two-staged technique are encouraging and indicate
that the creation of an endothelium on the inner surface of prosthetic grafts is
feasible in humans.
The replacement of arteries with purely synthetic vascular prostheses often
leads to the failure of such reconstructions when small-diameter or low-flow
locations are concerned, due in part to the thrombogenecity of the internal graft
surface (Bordenave et al. 1999). In order to improve long-term patency of these
grafts, the concept of endothelial cell seeding has been suggested because this
metabolically active endothelial surface plays a major role in preventing in vivo
blood thrombosis and because vascular grafts placed in humans do not
spontaneously form an endothelial monolayer whereas they do in some animal
models (Mori 1989; Niklason et al. 2001; Niklason et al. 1999).
Research has been concentrated in studying the interaction between the
prosthetic graft (scaffold) and the tissue culture. Some of the most significant
studies in relation to the liver include:
81
• Cell adhesion to fibronectin – pre-coated, smooth and textured silicon
• Dynamic cell seeding – for bone tissue in high aspect ratio rotating
bioreactor using polymeric scaffolds (prospective use for the high
density hepatic cell culture)
• Bioartificial Liver support system
• Heterotypic cell co-culture (several cell types grown simultaneously)
2.9.3. Methods of cell seeding and cell adhesion
The conditions, under which a graft is seeded with endothelial cells in situ
(including the axial force) has a huge impact on the graft’s performance post-
transplantation (Charara et al. 1999).
The composite structure resulting from the combination of biologically active
cells to prosthetic materials creates more biocompatible vascular substitutes. To
achieve endothelialization of synthetic vascular grafts, “one-stage” procedures
(cell seeding on the graft at time of implantation) were first developed which
are now replaced by “two-stage” procedures (in vitro cell seeding and growth
followed by implantation). Demonstration of the superiority of the two-stage
method in terms of significantly increased patency of the graft is shown in
(Bordenave et al. 1999) where successful endothelialization of grafts was
observed.
2.9.3.1. Endothelial cell seeding
Endothelial cell seeding is carried out in order to improve long-term patency of
purely synthetic vascular prostheses, to improve biocompatibility and decrease
thrombogenecity.
Long-term patency of artificial vascular grafts for hemodialysis access and for
bypass or interposition in small calibre arteries is limited due to neointimal
hyperplasia and associated graft thrombosis (Ballermann and Ott 1995;
Greenwald and Berry 2000). Given the anticoagulant and vasodilator properties
of endothelial cells, these problems could be partially overcome if grafts were
seeded with an adherent monolayer of differentiated endothelial cells prior to
implantation.
82
Endothelialization of a vascular graft in clinical settings can be achieved either
by high cell seeding density or by creating surfaces on which endothelial cells
adhere and grow to a confluent layer (Sipehia et al. 1996).
In most in vitro studies the vascular endothelial cells are cultured on either glass
or plastic cover slips and then subjected to flow conditions. However, this
method is not adequate for studying the dynamics of cell adhesion on solid
surfaces and therefore Brian Lin (2000) developed a flow system for testing the
cell responses to shear as well as a method for studying the various
morphological and physiological effects of flow on endothelial cells cultured on
solid biomaterials. It is known that endothelial cells form monolayers in vivo
which line the vascular wall and serves as a selective barrier between flowing
blood components and the vessel walls (Nerem et al. 1998). The endothelium
provides a non-thrombogenic surface, which allow for the diffusion of nutrients
and gases (Nerem et al. 1998).
Dynamic cell seeding involving cell attachment to microcarrier scaffolds
during rotating culture, showed evidence of a lower rate and extent of
proliferation compared to control cells cultured under non-rotating conditions
(Botchwey et al. 2001).
2.9.3.2. Cell differentiation
Cell-cell communication between multiple cell types in tissues is essential to
maintain differentiated cell function. Organ regeneration and tissue engineered
constructs require coordinated cell communication to produce and maintain
differential functions of several cell types simultaneously. Below is an example
of the complexity and novelty in grafting two different cell types to achieve
preliminary success in organ culturing. In 2001 Yamato Masayuki et al.
(Masayuki et al.) utilized patterned surfaces to produce a successful heterotypic
cell co-culture. Using Bovine plasma fibronectin in Dullbecco’s phosphate-
buffered saline (PBS) solution at 20oC and adding hepatocytes initially adherent
at 37oC for 24h and consequently having their temperature reduced to 20oC
following which endothelial cell suspended in culture medium at 37oC were
83
added to the patterned dishes in which the hepatocytes remained adhered to the
ungrafted surface area.
In the case of the liver, which is extremely complex structurally and
functionally, most researchers are concentrating on some aspects and cannot
replicate the anatomy and physiology of the human organ. Even though it looks
like replicating blood vessels is much simpler than creating an organ, here also
we have at least two types of cells performing different functions and their
interactions are still an object of many studies.
According to Ramm (2000), myofibroblasts are the primary cells responsible
for increased matrix deposition in hepatic fibrosis. In chronic cholestatic liver
injury, one of the earliest events in the development of hepatic fibrosis is the
activation of hepatic stellate cells and portal fibroblasts to cells with a
“myofibroblasts-like” phenotype, which are largely responsible for the
increased deposition of extracellular matrix components, including collagen,
observed in hepatic fibrosis
2.9.3.3. Possible improvement in endothelial cell growth
Covering the luminal surface of a vascular prosthesis with endothelial cells is a
process that may require the presence of growth factors (GFs) and extracellular
matrix support. Endothelialization could be improved by combining both GFs
and an extracellular matrix analogue. Sirois et al. (1993) carried out
experiments with human umbilical vein endothelial cells to determine which of
a number of different biological substrates made of type I or IV collagens,
gelatine, fibronectin, fibrin, laminin, chondroitin sulphate, heparin or hyaluronic
acid could be used to support endothelial cell culture. Apart from the different
surfaces endothelial cell growth supplement (ECGS) was incorporated in (for
group 1) or overlaid on (for group2) the substrates; or present in medium (for
group 3); or absent (for group 4). Growth was relatively stable for the first 48
hours, but later in groups 1, 2 and 4, cell death was observed on all the
substrates except for fibronectin. In group 3 where the ECGS was present in the
medium, growth increased and confluence was reached within 5-8 days on all
the substrates except for gelatine and type I collagen. Those experiments
84
suggest that continually delivered growth supplement in a fresh soluble form
seemed to be the appropriate condition to obtain an endothelial cell lining.
Synthetic vascular grafts do not spontaneously endothelialize in humans and
require some form of anticoagulation to mainly patency. Bhat et al. (1998)
studied and reviewed various methods of EC seeding by pre-seeding synthetic
graft materials such as expanded polytetrafluoroethylene (ePTFE) and
polyethylene terephthalate (PET) with endothelial cells (ES). The results
indicate that a heterogeneous ligand treatment of graft surfaces using avidin-
biotin and Fn-integrin attachment mechanisms increased cell seeding efficiency,
initial cell retention and cellular spreading.
Bruder et al. (2001) studied the phenomenon of low mortality rate in southwest
France by determining changes resulting from the interaction of endothelial cell
with resveratrol (a component of wine). Resveratrol treatment leads to increased
adherence of BPAEC (bovine pulmonary artery endothelial cells) under
simulated arterial flow conditions and the cells were evenly distributed
throughout the area of the cover slip exposed to flow. When compared to the
control sample, a significant percentage of resveratrol-treated BPAEC remained
attached to the plastic cover slip after 2 min and 5 min flow challenge (no cells
were left attached in the control after 5 minutes of flow conditions testing).
2.9.4. Difference between static and dynamic conditions for cell
seeding and cell adhesion
One of the major focuses of recent studies in cell adhesion has been the big
difference between cell adhesion in vitro and in vivo (or in vitro under
conditions mimicking the in vivo conditions). It has become apparent that the
seeding technique has a great role in the success of the cell adhesion and thus,
the tissue growth. Brief comparisons of the currently used approaches will be
discussed in this section.
There are currently a number of methods used for cell seeding. Some of the
most popular ones are:
85
• Passive cell seeding, where the scaffold is placed in media with
suspended cells. This method uses gravity as a means of cell adhesion,
and the number of cells adhering is lower than in dynamic cell seeding.
Dynamic cell seeding, where there are a couple of different approaches -
either the scaffold is not moving and is fixed so that media and
suspended cells flow through it either at a constant rate or pulsatile flow,
or the scaffold is moving through the media with cell suspension either
rotating, being shaken or in other ways floating through the media.
Endothelial cells (EC) covering the surface of a prosthetic material which
comes into contact with the blood could potentially enhance the non-
thrombogenicity of the surface. In order to create such a surface, the EC must
become attached to the surface, spread and ultimately form a monolayer. Jarrell
et al. (1991) came to the conclusion that a new method of attachment of EC –
by filtering EC onto the graft luminal surface (dynamic seeding) – had a 2 to 5-
fold increase in EC attachment when compared to gravity forced cell
deposition.
One of the difficulties facing the development of a bioartificial liver is that
hepatocytes show little ability to proliferate under the usual culture conditions,
and an adult liver normally contains >1011 parenchymal hepatocytes in humans.
Hepatocytes tend to lose their metabolic functions rapidly within a few hours in
suspension culture and thus cannot resist a long immunization process lasting
several hours.
Cell seeding is one of the key procedures in the construction of tissue-
engineered organs. Yang et al. (2001) used a packed-bed reactor utilizing
porous poly vinyl formal (PVF) resin as a 3-D scaffold to achieve high-density
cultures of hepatocytes (above 1x107 cells/1cm3 substrate) and long-term
maintenance of metabolic function to create a bioartificial liver. The cell
seeding method they used was centrifugal cell immobilization which achieved
improvement of efficiency was improved to about 70% after a serial centrifugal
cell immobilization process.
86
2.9.4.1. Electrostatic endothelial cell seeding method
A feasibility study was conducted (Bowlin and Rittgers 1997) using an
electrostatic endothelial cell (HUVEC – human umbilical vein endothelial cells)
seeding technique, a static pool apparatus, a voltage source, and a parallel plate
capacitor. The HUVEC concentration and seeding times were constant at 560
000 HUVEC/ml and 30 min, respectively. The results indicated that the total
number of adhered endothelial cells was 2.4 times higher when the ePTFE had
an induced positive surface charge than without such a charge and the number
of flattened (matured) cells were 8.1 times higher compared to controls. Those
results indicate that electrostatic interaction is an important factor in both the
endothelial cell adhesion and spreading processes and suggested that the
electrostatic seeding technique may lead to an increased patency of small
diameter (<6mm) vascular prostheses (Bowlin and Rittgers 1997).
2.9.4.2. Dynamic cell seeding technique
Dynamic cell seeding of tubular scaffolds can be achieved with higher density
when the scaffold actively moves through the media, via rotation for example.
If a magnet is attached to the ends of the scaffold or is inserted in the tube and a
magnetic field is applied using the same principle as when stirring chemicals,
more cells attach due to the higher collision rate between cells and scaffold.
Surely, any other method for moving the scaffold through the media will
produce similar results. However, careful evaluation of the collision force needs
to take place before this type of dynamic seeding is used, as high force might
damage the cells, affecting their viability and long-term functional and adhesion
abilities.
An alternative method of dynamic cell seeding would be to have the scaffold
fixed in place while the media and suspended cells are flowing through the
scaffold. Here again we have a collision force, but it is one-way collision as
only the cells are moving (in the previous methods the scaffold was moving in
random directions and thus creating a variety of collision forces and angles).
Thus, we believe that this method for dynamic cell seeding is less damaging to
the cells and thus have created a bioreactor, where the scaffold can be fixed and
the media with the cell suspension can be pumped (both continuously and
87
pulsatile) through the scaffold. Details on the construction of the bioreactor and
its function are presented in the Bioreactor chapter later in this thesis.
An investigation on hepatocyte transplantation using biodegradable polymer
matrices as an alternative treatment to end-stage liver disease was made by Kim
et al. (2000). In such procedure one of the major limitations has been the
insufficient survival of an adequate mass of transplanted cells. They
investigated a novel method of dynamic seeding and culture of hepatocytes in a
flow perfusion system. The conclusion of that study was that hepatocytes can be
dynamically seeded onto biodegradable polymers and survive with a high rate
of albumin synthesis in the flow perfusion culture system.
Dynamic cell seeding may however not be suitable for all applications. An
example of such case is the novel approach to grow in vitro mineralised bone
tissue utilizing lighter-than-water, polymeric scaffolds in a high aspect ratio
rotating bioreactor (Botchwey et al. 2001). Dynamic cell seeding, used in this
study, and cell attachment to micro-carrier scaffolds during rotating culture
showed a lower rate and extent of proliferation than those cultured on non-
rotating controls.
In any case, a flow perfusion system could be developed and used for the
dynamic or static cell seeding of 3-D resins and vessel grafts. With this in mind
the Bioreactor was developed as part of this thesis.
2.10. Scaffold requirements Ideally, a scaffold should have the following characteristics:
• Be highly porous with an interconnected pore network for cell growth
and to allow the flow transport of nutrients and metabolic waste;
• Be biocompatible and bioresorbable with controllable degradation and
resorption rates to match tissue replacement;
• Have suitable surface chemistry for cell attachment, proliferation, and
differentiation; and
• Have mechanical properties to match those of the tissue at the site of
implantation.
Apart from these universal requirements there are specific requirements
depending on the type of tissue to be grown. Scaffolds for heart valve have to
88
take into account the high rate of calcification, while the scaffolds for arteries
are more exposed to thrombosis. The changes that occur in the graft before and
after implantation have to be monitored.
Using a scaffold made out of a totally biodegradable material will ideally leave
us with a 100% cell tissue vessel at time of implantation, thus, minimizing the
problems which normally occur following implantation.
The following are some of the most desirable characteristics of in vitro seeded
blood vessels prior to implantation:
• Endothelialised blood-contacting surface
• Cellular potential for extracellular matrix synthesis, remodelling and
repair
• Appropriate heterogeneity, anisotropy and amount of extracellular
matrix
• Stable geometry but potential growth with the patient
• Stable mechanical properties
• Absence of harmful immunological and other inflammatory processes
• Resistance to tissue overgrowth
• Resistance to infection
• Chemical inertness and lack of hemolysis
• Easy and permanent insertion
• Minimal thrombogenicity
A major limitation of any in vitro grown tissues is the resultant stiffness of the
cultured construction (Shinoka 2002).
The creation of small size grafts is an area of concentrated research. Grafts
smaller than 6mm in diameter have specific requirements of the scaffold
material, physiological conditions and time for cultivating, cells differentiation
etc.
Most studies on small size grafts have been directed to arterial vessels cultured
with mammalian cells. The material used to build the scaffold needs to be
strong enough not to dissolve before the new vessel has been formed, but once
89
the new structure is completed, it needs to degrade. This is a challenge not just
for the small grafts. There are suggestions that increasing the cultivation time to
8 weeks can improve the vessel morphology and mechanical characteristics
(Niklason et al. 2001).
Creation of vascular grafts that mimic the native vessel in terms of mechanical
responses, elasticity and endothelialization is the focus of many studies
(Greenwald and Berry 2000; Mitchell and Niklason 2003). A study by
Nicklason et al. (1999) provides an example of the successful creation of small-
diameter grafts using biodegradable polyglycolic acid scaffolds with chemically
modified surface (with sodium hydroxide), under pulsatile and non-pulsatile
conditions in vitro (in a bioreactor) for 8 weeks.
2.11. Scaffolds and scaffold materials The term scaffold is used to describe the graft, made of biomaterial in the
desired shape, which will be the base onto which the seeded cells will adhere
and consequently grow. In this study a number possibilities for scaffolds were
investigated, differing from each other by shape and/or material.
Scaffold materials cover a large range of materials and can be divided in
numerous categories, which can include, but are not limited to:
• Biodegradable (covering the range from highly (polylactic-co-glycolic
acid) to minimally biodegradable (segmented polyurethane)) or non-
biodegradable
• Made out of polymer or metal (although for the purpose of cell seeding,
plain or modified glass is also in use, it is not used as a scaffold
material)
• Smooth surface, knitted, woven, non-woven, moulded scaffold
• Made of one polymer or co-polymer mixture
• Purely synthetic, purely natural or a combination of the above
In recent years one research area, which has attracted attention is in vitro tissue
growth on scaffolds fabricated from different materials. These techniques are
used for the growth of ligament, bone, cartilage, soft tissue, blood vessels and
even whole organ growth, as is the case for the liver and skin. The scaffold
90
material must satisfy a number of requirements with respect to the application
of the “new body part” and cell type used to build the tissue structure.
While some of the materials used have been shown to be effective in certain
applications, others have been rejected, and even there is little or no knowledge
on the behaviour of materials for other applications.
Scaffold material can be classified based on mechanical properties:
• Stiff scaffolds, with minimally elastic or rigid properties (glass,
metal and some polymers). These materials are used in models where
the wall movements can be neglected. In this thesis a glass model is
used for LDA measurements because of its transparency although the
lack of flexibility of the wall of the vessel was deemed a limitation.
• Moderate compliant scaffolds, with some flexibility (most polymer
scaffolds). These are the most commonly used ones for tissue culture of
blood vessels as they, to some degree, mimic the movements of the
vessel wall.
• Highly compliant scaffolds (such as thin rubber). Although better
representing the elasticity of the blood vessel wall, these are difficult to
keep in shape and can deform more than the natural vessel they model.
Many studies over the past three decades have shown the importance of the
elasticity and compliance of vascular grafts being as close to those of the native
vessel as possible. Thus, mimicking the properties of the vessel where the graft
will be implanted is a pre-requisite for the success of the grafting. This has been
reviewed in more detail later in this chapter as well as in the following chapter
on Methodology.
Cell adhesion and proliferation on biomaterials is a key issue in the study of
cell-biomaterial interaction (Ellingsen and Lyngstadaas 2003; Van Kooten and
Von Recum 1999). With the development of new disciplines within
biomaterials research such as tissue engineering and cellular therapy,
information at molecular and structural levels is needed in order to envisage and
design biomaterials that bring out specific and functional cell responses.
91
The inner surface of a scaffold which is the point of contact with the blood can
be smooth, rough or porous (Ellingsen and Lyngstadaas 2003; Mori 1998). The
first ones are considered less thrombogenic, the second ones are also less
thrombogenic and preferred for vascular grafts, and the third one in still
controversial.
Usually, grafts fail due to one of the following four reasons (Mori 1998):
• Technical failure within days of the operation
• Failure within weeks due to inherited thrombogenicity of the graft
• Failure within months due to progressive occlusion of the graft
• Late stage failure (over a year) due to degradation and structural changes
of the graft
It is also important to find alternative techniques to suturing of the grafts, as
compliance of the vessel and the graft are mismatched at the point of the suture,
and thrombi formation is higher around the joint (Mitchell and Niklason 2003;
Mori 1989).
2.11.1. Comparison of materials
In recent years, apart from donor or animal blood vessel transplantation, a push
for artificially creating a vascular graft resulted in many new materials being
developed. Some natural polymers are used because they have better material-
blood cell interaction and they promote the maintenance of cell differentiation
(Langer and Vacanti 1993). On the other hand, they are difficult to control, thus
synthetic polymers, which are easy to control in terms of rate of degradation,
chain length and molecular weight are used as an alternative material. A
combination of both types of polymers would presumably be the ideal material
for vascular grafts. So far no ideal material has been found, although some
materials show promising results.
In 1999 researchers (Van Kooten and Von Recum) determined the formation of
focal adhesions and fibronectin fibrillar structures by human fibroblasts and
human umbilical vein endothelial cells adhered to fibronectin-precoated,
smooth, and textured silicones as a function of time. Textures consisted of
parallel ridges and 0.5 mm deep grooves with a width of 2, 5, and 10mm. Cells
did not proliferate on the silicone surfaces without fibronectin pre-adsorption.
92
Cells adhered to glass removed all the pre-absorbed fibronectin, whereas on
silicone, they did not.
Sefton et al. (2001) performed a series of assays for the evaluation of
hemocompatibility of cardiovascular devices in 2001. Leukocyte and platelet
activation was studied by them and the materials used were tubes (inside and
outside surfaces) 5-7mm in length. Heparinized whole blood (1 U/ml heparin)
was incubated inside the tubes. The summary of SEM results after 1h exposure
to heparinized whole blood (1 U/ml) showed the following (Table 2.10.1):
Pellethane® (PEU) A little fibrin, occasional platelet, a few rbcs; moderate
fibrin with more rbcs, rare plt; red cells, fibrin, increased
spread platelets
NH4 plasma treated PEU
(PEU-NH4)
Some to moderate activated platelets, plus leukocytes and
rbcs
H2O plasma treated PEU
(PEU-H2O)
Increased fibrin/platelet rbc aggregates; pseudopodial to
fully-spread platelets
Fluorinated PEU (PEU-
fluorine)
High number of activated platelets plus secondary
pseudopodial platelets
Polyethylene imine
treated PEU (PEU-PEI)
Activated leukocytes
Heparin treated PEU
(PEU-heparin)
Activated leukocytes
Polyethylene (PE) Non-reactive for the most part; a few areas of activated
platelets
H2O plasma treated PE
(PE-H2O)
A few activated leukocytes & platelets
CF4 plasma treated PE
(PE-CF4)
Fibrin masses, some with rbcs
Nylon Moderate amount of spread platelets
Latex Fibrin masses, thrombi
Table 2.11.1.1. Comparison of different materials (Sefton et al. 2001)
93
In 2001 researchers (Meng-Yen and Jui-Che) carried out a study on
biocompatibility and thrombogenesis of self-assembled monolayers (SAM)
containing alkanethiol with phosphonate ester and phosphonic acid
functionalities on gold model surface.
In contrast to a polymer material modified by plasma processing or a grafting
reaction, the SAM technique can provide a densely packed, well-defined, and
highly ordered surface.
Those examples are given to illustrate the diversity of materials used or studies
nowadays. Materials more appropriate for blood vessel scaffolding are
discussed further in this chapter.
When we were looking into the available materials from which to choose,
surgical suture and mesh material was considered the most suitable for building
the scaffold. More focused research pointed out that the absorbable sutures
might be suitable. These are the most popular type of sutures that lose their
tensile strength, to various degrees, after 60 days under the skin (i.e. implanted):
• Catgut Suture
• Treated Catgut Suture (Mild Chromic Gut)
• Polyglycolic Acid Suture (Dexon)
• Polylactic Acid Suture (Vicryl)
• Polydioxanone (PDS)
• Polyglyconate (Maxon)
For some of the materials on this list a brief comparison was made and is
presented below with more detailed comparison further in this chapter:
Vicryl is made out of polymer – Lactide and Glycolide; the coating is
Polyglactin 370 and Calcium stearate. The suture is completely absorbed
between days 60 and 90. The coating mixture forms an absorbable, adherent,
non-flaking lubricant. All these components are water repelling, which slows
tissue fluid penetration and absorption. This suture is commercially available
coated and uncoated (Johnson & Johnson Gateway).
94
Vicryl is hard to shape, thus is difficult to use to manufacture complicated
shapes and has proven unsuccessful for the creation of a tri-leaflet heart valve
(Shinoka 2002).
Polydioxanone (PDS): Complete Suture absorption day 180 after the operation,
which might not be suitable for vascular grafting.
Maxon has the advantage of being available as a monofilament. A possible
disadvantage is the long degradation time - complete absorption of suture is
from day 180 to 210 post operative.
Suture materials
Polyglycolic, polyglactic acid polymer-derived sutures (PGLA) such as Vicryl
and Dexon are absorbed via enzymatic degradation by hydrolysis (Aderriotis
and Sandor 1999).
Irradiated polyglactin 910 (IRPG) Vicryl Rapide (Ethicon, Somerville, N.J.)
was the suture we used in our tissue experiments and it is a braided co-polymer
of glycolic and lactic acid that is surface treated with polyglactin 370 and
calcium stearate and has received gamma radiation (Johnson & Johnson
Gateway).
Name Materials Absorption Support Novelty
Vicryl (Braided)
90%Glycolide 10%L-lactide Coating: polyglactin 370 and calcium stearate
65%strength at 14 days, 40% still at 21 days, Complete absorption by 70 Days
Still not established safety in cardiovascular and neural tissues
Panacryl (Braided)
95%lactide 5%glucolide coated with 90%caprolactode and 10%glucolide
Between absorbent and non-absorbent Strength after 3months 80% and 60% after 6 months
For long-term wound support- up to 6 mnths
Not for use in cardiovascular tissues
Coated Vicryl Rapide (polyglactin 910)
90%glucolide 10% L-lactide (C2H2O2)m(C3H4O2)n Coating polyglactin37& calcium stearate
All initial strength is lost by 10-14 days. Full absorption by day 42 via hydrolysis
For short support within 7-10 days
Only available undyed
Monocryl (poliglecaprone 25)
Copolymer of glycolide and epsilon-
Strength after day 7 60-70%, 30-40% after day 14. Total strength loss
Not for use in cardiovascular tissues
95
Monofilament
caprolactone at day 28. Total absorption via hydrolysis day 91-119.
PDS ІІ (polydioxanone) Dyed and Clear Monofilament
Prepared form the polyester, poly (p-dioxanone) (C4H6O3)x
Original strength after 14days 70%, 59% after 28days, 25% after 42day with total absorption after 6 months
Not indicated in adult cardio-vascular tissue, but used in pediatric cardio-vascular tissue Not to be used in conjunction with prosthetic devices
Table 2.11.1.2. Comparison of commercially available suture materials
In the tissue culturing tests done as part of developing the bioreactor, cell
adhesion on Vicryl (braided) was performed, chosen for its gradual degradation,
low cost and easy accessibility in any surgical clinic. Those experiments are
discussed in brief in the Methodology Chapter of this thesis.
Mesh used in surgery
Historically (Seiler and Mariani 2000), collagen-coated Vicryl mesh composed
of a watertight film of bovine collagen and polyglactin 910 (Vicryl, Ethicon)
has been used. Because of the theoretical danger of transmitting bovine
spongiform encephalitis with bovine collagen, this material was later changed to
a fleece composed of polyglactin 910 and poly-p-dioxanone (Ethisorb, Ethicon).
The patch is available in different sizes, softens when immersed for a few
seconds in liquid, and can be cut to any size. It is easily handled and relatively
inexpensive, and elicits a minimal inflammatory response.
Some researchers (Burg et al. 1999) created a scaffold using a Vicryl mesh
folded into the desired shape with its edges heat-sealed leaving the centre
empty.
Both Vicryl woven and Vicryl knitted mesh are prepared from uncoated,
undyed fibre identical in composition to that used in Vicryl synthetic absorbable
suture. They are available in single sheets sized 15x15cm and 30x30cm.
Minimal absorption of the mesh until 6 weeks and complete absorption between
60 and 90 days of implantation provides a good operation range for tissue
repair. The knitted one has less strength but maintains 80% of the original
96
strength after 14 days. The woven one has double initial strength but only about
23% remains after 14 days. The use of neither has been established for
cardiovascular tissue (Johnson & Johnson Gateway).
We recommend manufacturing the mesh in the desired scaffolding shape
instead of cutting sheets of mesh to appropriate size, and then folding it up to
achieve a 3-D scaffold. This would avoid the non uniform mesh of the scaffold
due to the joining of the ends of the sheet, which creates a region similar to the
one of sutured shunts to native vessels, thus being an area of changed
hemodynamics and increased thrombogenicity.
2.11.2. Techniques for manufacturing scaffolds
A number of different processing techniques have been developed to design and
fabricate three-dimensional (3D) scaffolds for tissue-engineering applications.
The imperfection of the current techniques has encouraged the use of a rapid
prototyping (RP) technology known as fused deposition modelling (FDM). The
FDM method is an RP technique that builds a physical model by depositing
layers of thermoplastic material one at a time.
Results from a study (Hutmacher et al. 2001) showed that FDM allowed the
design and fabrication of highly reproducible bioresorbable 3-D scaffolds with a
fully interconnected pore network. This study looked at the mechanical
properties and in vitro biocompatibility of polycaprolactone scaffolds with
honeycomb-like pores and a porosity of 61 ± 1% and two matrix architectures -
the first scaffolds had a 0/60/120o lay-down pattern and the second scaffolds
with a 0/72/144/36/108o lay-down pattern. The second pattern had a
compressive stiffness of ½ of the first pattern and 1% offset yield strength in air
nearly 25% lower, and in simulated physiological conditions over 25% lower
compressive stiffness and over 10% lower in the 1% The stress-stain curves
obtained for both scaffold architectures demonstrated the typical behaviour of a
honeycomb structure undergoing deformation. In vitro studies were conducted
by the authors (Hutmacher et al. 2001) with primary human fibroblasts and
periosteal cells showing that both cells could proliferate, differentiate, and
produce a cellular tissue in an entirely interconnected 3D Polycaprolactone
97
matrix. The above technique was given as an example and other techniques can
be used for FDM.
2.11.3. Dacron prostheses
Human umbilical vein endothelial cells (HUVEC) on knitted and woven Dacron
prostheses were compared with HUVEC on smooth surfaces (tissue culture
polystyrene, PET film, and Natrix) with regard to adherence, growth, and
susceptibility to injury by neutrophils (PMN) in 1995 (Tunstall et al.). For
prosthetic materials of given macroscopic dimensions, more endothelial cells
(ES) adhered to these materials than to smooth surfaces. However, the
prostheses had a greater effective surface area as determined by the number of
EC at confluency. When this parameter was taken into account, fewer EC were
found adhering to prosthetic materials per unit effective surface area than for
the smooth surface substrates. Growth on prostheses was clearly inferior to that
on smooth surfaces, and EC on prostheses were more prone to attack by
activated PMN than on smooth surfaces. These differences may reflect the
topographic differences in cells attached to fibres where they assume more
distorted shapes by stretching to span gaps in the fibres.
2.11.4. Non-woven scaffold
Highly porous grafts (loosely woven or knitted) can cause blood leakage
through the wall, thus a careful evaluation of the porosity and the method for
manufacturing of the scaffold needs to take place.
The results of a study performed by Pahernik et al. (2001) indicated that non-
woven polyurethane sheets supplied a biocompatible support structure for
functionally active high-density cultures. According to the researchers, the
optimal cell density in a three-dimensional culture configuration was 1x106
cells/cm2.
Polyester non-woven fabric (Naruse et al. 2001) has been successfully used to
culture porcine hepatocytes.
Textured surfaces stimulate adhesion and cell growth, as opposed to smooth
surfaces such as polyurethane (Belanger et al. 2000).
98
2.11.5. Modified ePTFE and PTFE
There is evidence to suggest thrombogenesis in polytetrafluoroethylene (PTFE)
grafts used in patients with renal failure (Anderson et al. 1980;Rapaport et al.
1981) is due to changes in graft diameter or unevenly distributed pressure and
shear stress.
McKeown et al. (1991) have modified expanded PTFE (ePTFE) with a simple
chemical modification which facilitates endothelialization without using
thrombogenic cell adhesives.
When comparing ePTFE and PTFE with or without coating, as in any other
comparison, the surface treatment and type of cells (including cell seeding
method) have to be the same. Such a study (Sipehia et al. 1996) using human
umbilical vein and human saphenous vein endothelial cells on ammonia plasma
treated ePTFE and PTFE has shown that both have significantly better cell
growth on the coated surface compared to uncoated material.
There are animal studies supporting the benefits of using PTFE in TIPS.
Following failure of a polyethylene terephthalate (PET) stent due to thrombosis,
PTFE TIPS implanted in the same animals showed patency and good function
(Haskal et al. 2002). Similar findings have been recently reported (April 2004)
by a group in UK (Barkell et al. 2004) as a retrospective study in 100 patients, 9
of which had PTFE covered TIPSS following stenosis of the primary uncovered
stent, and the rest had covered stent placement only. They have reported
improved patency rate, which might reduce the invasive portography follow
ups.
2.11.6. Biodegradable scaffold
A synthetic biodegradable scaffold consisting of polyglactin and polyglycolic
acid fibers has been seeded in vitro with mixed (endothelial and fibroblasts) in
(Shinoka et al. 2000). The key benefit of a biodegradable polymer scaffold is
that it will degrade in vivo as seeded cells proliferate, so the long-term presence
of foreign materials would be avoided. In that study the mesh used as scaffold
consisted of a polyglactin woven mesh sandwiched between either two
nonwoven PGA mesh sheets or a polyglactin woven mesh reinforced with
copolymer of caprolactone and lactide, in both the mesh matrix had more than
99
95% porosity before seeding. This provides a strong shape and also ensures the
degradation of the material. Some of the other requirements of course include
the ability to determine the degradation rate.
Synthetic biodegradable scaffold consisting of polyglycolic acid fibres, seeded
with fibroblasts and subsequently coated with endothelial cells could be used
and the rate of degradation can be monitored and modified (Shinoka 2002).
2.11.7. Other types of scaffolds
The large variety of scaffolding materials currently used for blood vessel and
hepatocyte cultivation makes it impossible to describe all materials and
methods. Apart from the ones mentioned previously, in this part of the chapter
some examples of other scaffold options are given.
Highly porous biodegradable poly(D,L-lactic-co-glycolic acid) with
immobilized galactose onto its internal surface has been successfully seeded
with rat hepatocytes (Park 2002).
Highly porous chitosan (partially deacetylated derivative of chitin) with fructose
onto the inner surface has shown improved cell density with rat hepatocytes (Li
et al. 2003). Endothelial cell adhesion to the fibronectin in artificial
extracellular matrix proteins shows good prospects for small-size vascular grafts
(Heilshorn et al. 2003) when seeded with human umbilical vein endothelial
cells.
For growing cardiomyocytes from rats, fibrin glue has been shown to be a
suitable ground matrix (Kofidis et al. 2003) with cells viable even in the
periphery of the tissue block.
Collagen components were successfully used for myocardial grafts (Kofidis et
al. 2003).
Tests carried out by Bèlanger et al. (2000) showed poor endothelial cells (from
human umbilical vein) and fibroblasts (from skin) adhesion on all polyurethane.
Polycaprolactone (PCL) is a semi-crystalline bioresorbable polymer with a low
glass-transition temperature of –60oC, a melting point of 60oC, and high
decomposition temperature of 350oC with a wide range of temperatures that
allows extrusion. A study (Hutmacher et al. 2001) using pellets of PCL
extruded and manufactured into 3-D scaffolds tested them in a phosphate-buffer
100
saline (PBS) solution preconditioned in PBS for 1 day, or at ambient conditions.
Polycaprolactone has been used in many studies due to its relatively simple
extrusion.
2.12. Coating of biomaterials The benefit of coating grafts prior to implantation can be seen from studies,
which compared coated and uncoated grafts. Most studies suggested that intimal
hyperplasia, which leads to graft failure, develops faster in uncoated grafts
(Debski et al. 1982; Mori 1989).
Although monoprotein coatings of biomaterials with either natural adhesion
molecules or genetically designed analogues have been used to assist
attachment and spreading of endothelial cells, such treatments were found
unsatisfactory in maintaining the integrity of the endothelial surface under
turbulent flow conditions (Nikolaychik et al. 1994).
There are different requirements and expectations from coating of devices for
implantation and external use (still in contact with the blood). For
extracorporeal circuits reduction of the inflammatory response can be achieved
using coating with poly(2-methoxyethylacrylate) (Saito et al. 2000), diamond-
like carbon film (Alanazi et al. 2000) and heparin coating.
Apart from improving endothelial cell adhesion, coating of biomaterials is
important for improving the thrombogenesis of shunts. As was discussed
earlier, the development of thrombi is one of the main disadvantages of TIPS.
This could be eliminated by coating the shunt with materials that discourage the
formation of thrombi (Collons and Sarfeh 1998).
Rough surface of the in vitro created blood vessel could lead to hemolysis of the
blood due to shear stress (Yousef 2001), hence smoothness of the graft needs to
be taken into consideration when choosing the scaffold material and cell
seeding methods.
2.12.1. Coating the material with a layer of endothelial cells
The endothelium has a number of vital roles in the functioning of the blood
vessel and allograft, some of which are releasing of pro-fibrotic cytokines,
101
taking on circulating leucocytes, proliferation of vascular smooth muscle cells
and deposition of extracellular matrix proteins (Waller et al. 2003).
For the improvement of vascular graft patency, an endothelial cell (EC) lining is
desirable as has been demonstrated by Seifalian et al. (2001), where they used
human umbilical vein endothelial cells (HUVECs) seeded onto graft materials
(CPU, ePTFE, and Dacron). EC seeding significantly improves the blood
compatibility of artificial surfaces (Bos et al. 1999). It is also essential that the
EC remains viable after being seeded onto the prosthetic graft. The polymers
currently used are Dacron and ePTFE, while a new compliant polyurethane
(CPU) is under clinical trial. Clinical studies have shown Dacron causes
thrombosis and neointimal thickening in low-flow states, so currently it is only
used in large-vessel implantation. The only alternative to autologous materials
in small-vessel reconstructions is ePTFE, however the long-term patency rate of
autologous implants is approximately 75% after 2 years, whereas the rate of
ePTFE is only about 30% (Esquivel and Blaisdell 1986). The principal reason
for late graft failure is neointimal hyperplasia and clinically it accounts for
about 80% of occluded vascular grafts (Chervu and Moore 1990). Such poor
long-term patency rates have driven the current search for new polymers and
novel biological vascular grafts with superior biocompatibility. EC seeding
improves patency rates and reduces early thrombus formation in some animal
models (Herring et al. 1994; Schneider et al. 1988; Stanley et al. 1982).
Endothelial cell attachment to a synthetic substrate, which surface was
chemically modified using either laminin or fibronectin, was studied (Scott and
Mann 1990) using an in vitro model system. That study confirmed that
biomolecules increase the attachment rate of endothelial cells to synthetic
substrate.
The adhesive interactions between blood cells and endothelial cells in regions of
low shear stress are assisted by the prolonged contact blood elements have with
the vessel wall (Henry and Chen 1993; Hinds et al. 2001; Lappin et al. 1998).
Small-diameter vascular grafts, for example, tend to have an early and high
occlusion rate (Hedeman et al. 1998). Cell seeding on the luminal surfaces of
small-diameter prostheses may provide an antithrombotic lining and improve
both the short-term and the long-term patency rates (Hedeman et al. 1998).
102
Kam and Boxer (2001) have described a method for endothelial cell seeding on
protein-micro patterned lipid bilayer surfaces, whereas the cells were cultured in
DMEM (Dulbecco’s modified Eagle’s medium) supplemented with 20% fetal
bovine serum under standard cell culture conditions (a humidified, 5% CO2/
95%air environment maintained at 37oC). Endothelial cells seeded onto glass-
supported bi-layers of egg PC had cell density reduced by 85% compared to
those on plain glass. This minimal adhesion of endothelial cells onto fluid lipid
bilayers supports the previous reports, which showed that lipid structures
inhibited fibroblast adhesion. Other conclusions that can be made are that the
bigger the gap in the surface, the better adhesion, and the thicker the strings of
the material, the lower cell spreading. The reduction of cell spreading has also
been associated with a decrease in cell survival. This example highlights the
importance of understanding the underlying principles of cell adhesion before
attempting to seed cells on grafts.
2.12.2. Coating with fibronectin and E-selectin
Conjugates of albumin and heparin provide non-thrombogenic coatings for
vascular grafts (Bos et al. 1999; Bos et al. 1998) which can be further enhanced
by adding fibronectin.
Heparin coating has shown to decrease the initial thrombus formation, but does
not favour endothelialization (Mori 1989, Chapter 22; Noishiki and Miyata
1986).
Heparin coating has to be done so to minimise the negative impact the coating
can have on the coated polymer (Tayama et al. 2000), and some studies with
Bioline coating system, although showing improvement in biocompatibility in
terms of leukocyte and complement activation do not improve platelet
activation and coagulation, thus have minimal clinical benefit.
In the study by Van Wachem et al. (1988) cellular fibronectin was deposited on
tissue culture polystyrene during the adhesion and spreading of cultured human
endothelial cells (HEC) indicating that the ability to deposit cellular fibronectin
onto a polymeric surface is a condition for the spreading and proliferation of
HEC.
103
The role that surface properties play in influencing the extent of
endothelialization of polymer surfaces was investigated by Absolom DR et al.
(1988). Their research suggested that for a wide range of polymer surfaces the
degree of endothelialisation (but not cell spreading) for both porcine and bovine
endothelial cells were directly related to polymer surface tension, i.e. the higher
the surface tension the higher the endothelialisation.
The benefit of 90o every 15 minutes for an hour graft rotation method is that the
presence of E-selectin allows for cell adhesion at higher wall shear stresses
under steady conditions compared to uncoated models (Hinds et al. 2001). It
has also been shown in that study that cell adhesion of coated models under
pulsatile flow is lower in magnitude and distribution compared to steady flow.
In addition, the adhesion rates before and after the stenosis region (if any) were
very different depending on the type of flow present (Hinds et al. 2001).
2.12.3. Carbon-deposited surface and Diamond-like Carbon coating
The adhesion and proliferation of endothelial cells can be drastically improved,
according to Kaibara et al. (1996), when cells are cultivated on a carbon-
deposited polymer surface pre-coated with either fibronectin or laminin (which
are most likely the reason for cell adhesion).
All mechanical devices in contact with blood have to be conditioned to prevent
thrombogenecity as much as possible. Not only clot formation is dangerous to
the patient, but it also disturbs the normal work of the device. Diamond-like
carbon films, having physical properties in-between diamond and graphite, are
hard, chemically inert and un-reactive. They have been used to coat rotary
blood pumps showing good biocompatibility (Alanazi et al. 2000). Although
Alanazi et al. (2000) showed better compatibility of diamond-like carbon coated
polymers compared to heparin and polycarbonate coated, they still do not
advice those devices to be used long-term.
104
2.12.4. Coated with grafted adhesion peptides
Decreased hepatocyte adhesion to polymeric constructs limits the function of
tissue engineered hepatic assist devices, but the study by Eric S. Carlisle et al.
(Charlisle et al. 2000) where they grafted adhesion peptides to polycaprolactone
(PCL) and poly-L-lactic acid (PLLA) using pulsed plasmadeposition in order to
mimic the in vivo extracellular matrix in an attempt to enhance hepatocyte
adhesion showed promising results..
2.12.5. Encapsulation of the graft
In order to avoid or minimise immunosuppression resulting from the need to
administer medication for the prevention of graft rejection, encapsulation
(immuno-isolation) has been developed as a novel technique. Some of the
limitations of this method include the type and size of the available
microcapsules and the optimal site for transplantation. The usual site is the
peritoneal cavity, but application through the portal vein has been tested in pigs
(Toso et al. 2003), while intra-portal administration is performed in most islet
allografts transplantations. Encapsulation of the graft has been shown to
increase the portal pressure initially and then decrease it to a normal level
(Toso et al. 2003) thus this method, if further studied and proven effective,
could complement the portal vein shunts currently in use.
2.13. Why pulsatile flow is important From the large number of reports on modelling of blood vessels, a very limited
number concentrate on, or take into account, the effects of pulsatility on the
model (Charara et al. 1999). There is also a lack of literature examining the
waveforms and pulsatility in patients with portal hypertension (Barakat 2002).
Under flow conditions vascular endothelial cells change their shape and
orientation depending on the nature of the flow (Nerem et al. 1998; Shiomi et
al. 2000; Verweyveld 1997). The flow can be laminar, pulsatile, turbulent, and
random or a combination of these. Increasing the level of shear stress benefits
the cell growth until a certain point, after which increased levels of shear stress
result in a decrease of cell replication. The aim is to find the right level of shear
stress so optimal conditions for cell growth can be achieved.
105
Cell adhesion rates before and after the stenosis region (if any) is very different
depending on whether the flow is steady or pulsatile (Hinds et al. 2001). Under
pulsatile flow, contrary to steady flow, adhesion in the stenosis is significantly
greater than in the recirculation.
The importance of pulsatile flow for arterial growth in vitro has been confirmed
by Niklason et al. (1999), where the strength of small diameter bovine vessels
cultured in vitro onto polyglycolic acid scaffold (with supplemented medium)
was even higher than the one of native human saphenous vein.
Portal vein flow has been usually described as non-pulsatile or continuous
(Keller et al. 1989; Taylor et al. 1985), although several authors have described
pulsatile flow (Barakat 2002; Gallix et al. 1997; Koslin et al. 1992; Partiquin et
al. 1987).
One common opinion is that the portal vein represents a low-pressure system
without significant pulsatility of flow (Hűbner 2000), thus the flow can be
modelled either as non-pulsatile or as continuum flow with minor pulsations.
Other studies examining the pulsatility of the flow in the portal vein (Barakat
2002) show a clear presence of pulse in the vein.
Both points of view have good arguments in their favour, and the limited
number of studies evaluating the presence of pulsatile flow in the portal vein
allow for either modelling technique. In this thesis pulsatile flow was simulated
in both computer and experimental simulations as this added complexity to the
study of flow through obstructed vessel.
2.13.1. Waveforms and pulsatility
A study carried out to examine the relationships between the hepatic vein
(HVW) and portal vein (PVW) waveforms in patients with cirrhosis and portal
hypertension (Barakat 2003; Barakat 2004) has shown the difference between
those waveforms and the ones in healthy subjects. In all healthy subjects the
PVW recorded by Doppler Ultrasound was pulsatile, however from the 148
patients with liver cirrhosis, 37.8% had flat PVW, and only about a quarter of
them also had a flat HVW. One of the most important conclusions of another
study by Barakat et al. (2002) was that the percentage of flat waveforms
increased with the progression of liver cirrhosis. Taking this into account, one
106
can assume flat PVW in advanced and end stage cirrhosis in most patiens. On
the other hand, the same study showed that there were differences between
individual patients and, as such, individual study of each patient is required to
evaluate the form of the PVW. This effect is taken into account from the author
of this thesis by suggesting the proposed model be individualised for each
patients’ portal blood flow velocity, pressure, pulsatility, stenosis, hematocrit,
cell adhesion etc.
The pulsatility index (PI) is inversely related to pulsatility ratio, where:
PI = [(peak maximum velocity) – (peak minimum velocity)] / peak maximum
velocity, or can be expressed as PI = [peak systolic velocity – end diastolic
velocity/ mean velocity (Tasu 2002).
In healthy subjects (Barakat 2002) the PI has been calculated as ranging
between 0.21-0.58 (with mean 0.39 ± 0.1), whereas in patients with cirrhosis it
was mean 0.23 ± 0.1
Interestingly, only 38% of cirrhotic patients had PI less than 0.2, i.e. almost
non-pulsatile and high pulsatility (0.5) was rare (in less than 2%). In all healthy
subjects however there was some pulsatility.
Portal vein mean flow velocity decreased in patients with chronic liver diseases
(Chawla et al. 1998). During pregnancy however the portal blood flow
increases in most women (Van Splunder et al. 1994).
Portal vein pulsatility and spectral width can be used as indicators for early
hemodynamic changes in patients with CLD (Barakat 2002).
Pulsatility is determined in this case as portal vein waveform fluctuations over
time (Gallix et al. 2002).
In healthy people the portal vein pulsatility ratio is around 0.66±0.08 and is
negatively correlated with right atrial pressure (Rengo et al. 1998).
Experimental studies have shown that arteries cultured in vitro under pulsed
flow more closely mimic the native artery than those cultivated in a non-
pulsatile environment (Bilodeau et al. 2005; Niklason et al. 1999) and also have
better and longer patency.
107
There are differences in the literature in describing the flow of blood to the liver
through the portal vein. In this work, two scenarios have been considered:
modelling using steady flow and then introducing pulsatility.
2.13.2. Endothelial cells – graft relationship
The establishment of an early blood-contacting endothelialised surface may
improve the graft-host relationship. In some studies endothelial cells seeded on
fibronectin-treated polyester elastomer (Greisler et al. 1989) show no
differences in EC adhesion between high- and low-shear conditions or proximal
vs. distal graft segments.
Endothelial cells in arteries are sensitive to pulsatile flow and flow-induced
shear stress as shown through numerous studies by L. Schalina over the past
two decades. If pulsatile flow is transmitted to veins a similar response of the
endothelial cells could be expected, and this could explain some of the changes
in the vessel shape (Schalina and Liepsch 2001).
2.13.3. Effect of hemodynamics on endothelial cells
Endothelial cells at the arterial wall subjected to various mechanical stresses
due to the flow of blood, the most important of which are the pressure force
acting normal to the cells and the shear stress acting tangentially and they both
vary with time (due to the pulsating flow) (Lin 2000; Nerem et al.1998; Shiomi
et al. 2000). Studies of the effect of flow on cell proliferation have been carried
out, and these have shown that the rate of cell replication decreased with
increased levels of shear stress.
Endothelial cells in vivo are highly adherent and can resist disruption by
hemodynamic shear stress at levels that far exceed physiological conditions.
Ballerman and Ott (1995) found that endothelial cells exposed to chronic shear
stress in vitro, applied in a stepwise fashion over several days, are provoked to
become “tightly adherent to the substratum and exhibit more differentiated
features”. Thus, pre-conditioning of endothelial cells seeded on vascular grafts
with stepwise shear stress in vitro could be used to improve endothelial cell
retention and differentiation for subsequent in vivo use.
108
A novel perfusion system for culturing human endothelial cells on small-
diameter PTFE grafts under defined pulsatile shear stress (Dunkern et al. 1999)
has been developed in 1999. This perfusion system enables culture of
endothelial cells on PTFE grafts to confluence under a wide range of shear
stress conditions in order to benefit form stronger adhesion of endothelial cells
to the substrate. The application of pulsatile flow with high shear stress (6.6
dyn/cm2, 5min) to a graft endothelialised under perfusion didn’t cause any
damage to the cells, whereas a shear stress of 3 dyn/cm2 applied for 5 min has
been shown to wash more than 50% of endothelial cells off PTFE graft when
cultured to confluence under static conditions (Dunkern et al. 1999).
Endothelialised vascular grafts can be pre-conditioned to defined shear stress
values.
With the use of ultrasound and pulsed Doppler Kiserud et al. (2003) have
described pulsation in the left portal branch in all studied subjects (10 fetuses
under 33 weeks with smaller diameter of the portal vein which might be the
reason for the pulsatility index being higher than that in the umbilical vein).
Another point of that study was that pulse wave and blood flow run in the same
direction in the left portal vein. To the best of our knowledge there are only few
studies mentioning or assuming pulsatility in the portal vein. Thus, the flow
model has been developed for both pulsatile and non-pulsatile flow in this
thesis.
2.13.4. Vessel compliance
Arterial tissue is continuously exposed to a dynamic mechanical environment
induced by pulsatile blood flow that exerts shear stress (tangential force),
pressure (normal force), and cyclic stretching. Recent biomechanical studies
have strongly implied that a combination of all these factors contributes to the
maintenance or regeneration of vascular tissue architecture (Hiromichi et al.
2001). Compliance mismatch between native artery and artificial graft has been
long discussed as a cause of graft failure during a prolonged period of
implantation of an artificial graft with a small diameter. In their study Sonoda
Hiromichi et al. (2001) have realised the artificial graft using segmented
109
polyurethane (SPU), which has been proven to be a highly durable, minimally
biodegradable synthetic elastomer in order to address this issue.
2.14. Other methods and approaches for addressing the
problems of cirrhosis of the Liver and vessel transplant in
general This brief discussion on liver assist devices is in addition to that described
previously in relation to methods for overcoming portal hypertension.
Hepatocyte systems for extracorporeal devices as well as implants are under
rapid development with somewhat promising results. Implants could provide a
permanent solution to end-stage liver disease, if successful. On the other hand,
extracorporeal devices have two advantages: control over the medium
surrounding the cells and decrease in the chance of immune rejection. The main
two disadvantages of the extracorporeal assist devices are the access of blood
between the body and the apparatus and the need of hospitalisation for the
duration of the procedure. In general, while extracorporeal devices are for
temporary use to assist a recovering liver or to serve as bridge until
transplantation, the implantable hepatocyte-based methods are for long-term
treatment with the intention of improving liver function. In both methods, the
ability to isolate and culture liver cells without losing their differentiated
functions on a large scale is still a challenge. Isolation of liver cells includes
mechanical dissociation or enzymatic digestion (Puviani et al. 1998). The later
one has the benefit of more viable cells as a percentage of liver volume in the
suspension after isolation. Establishment of a bioartificial liver support system
using genetically modified hepatocytes is a potential approach to improve the
treatment of severe liver failure. Kawashita et al. (2000) designed a method for
medicated gene transfer porcine hepatocytes growth for the creation of
bioartificial liver support system.
Another, even more rapidly progressing area is development of artificial blood
vessels, predominantly arteries. Many materials have been used for large
diameter grafts for arteries with some success, and some of them were discussed
earlier in this chapter. For the small diameter grafts, more inert materials
110
(heparin coated or endothelial lined) could be more appropriate. A challenge is
how to coat the material to increase biocompatibility, and in the same time
decrease adhesion of blood cells and substances to the graft. Heparin coating of
coronary artery bypass graft, for example, has shown little clinical or biological
benefit, but might be useful for at risk patients or some more complex
procedures (Collart et al. 2000).
Even though most of this research can be utilised for designing artificial veins,
due to histological and functional differences, this area needs more research and
special consideration.
Most treatments incorporate some of the detoxifying functions of the liver,
using dialysis, charcoal hemoperfusion, immobilized enzymes and exchange
transfusion (Langer and Vacanti 1993).
Recently, interest in hepatocyte transplantation has increased and the clinical
experimentation of hepatocyte-based liver support has attracted many
researchers. Promising reports of clinical usage of isolated allogenic
hepatocytes in hepatocellular transplantation and of xenogenic liver cells in
constructing bio-artificial liver support systems come from various groups
worldwide. From a clinical perspective the advantages and use of isolated
hepatocytes for supporting an acutely devastated liver or a chronically diseased
liver, and for correcting genetic disorders resulting in metabolically deficient
stages, are major reasons for the interest in this approach (Puviani et al. 1998)
Gene transfer and epithelial cell transplantation technologies play important
roles in the development of new therapeutic concepts for liver diseases (Ott et
al. 2000).
Animal experiments have been carried out on orthotopic liver
autotransplantation and even though reports on such procedures have been
promising (Gruttadauria et al. 2001; Urban et al. 2002), the impact this
technique has on the long-term function of the liver has not been studied. Many
recent animal experiments have shown promising results, and some have shown
liver regeneration after partial hepatectomy. In rats, for example, the original
liver mass could be restored after few days (Nadal 2000). This thesis does not
111
look into this or similar techniques, as the possible shunt model is for
maintaining the blood flow, not redirecting it away from the liver.
In vitro growth of thick tissues, like the liver, as part of culturing the organ in
laboratory conditions so far have not been successful (Kaihara et al. 2000).
Healthy liver cells were shown to grow well on sponge-like silicon chips during
a two-week trial carried out by a group from the University of California, San
Diego (Gornam 2001).
2.15. Future work The fused deposition modelling (FDM) method is a rapid prototype (RP)
technique that builds a physical model by depositing layers of thermoplastic
material one at a time. It was deemed that this is the most precise technique as
FDM reproduces an exact copy from a computer file (usually a CAD file) of the
scaffold. It lays the filament working on one plane at a time and can achieve a
smooth surface if desired for seeding purpose. The technical difficulty we had
using this “ideal” technique was the diameter and stiffness of the filament to be
fed into the FDM machine. Our machine could not be manipulated to accept
various thickness filaments, as the lead rolls were factory-fixed at a specific
diameter. This prevented us from using any of the biodegradable materials in
our simulations in the form of 3-D scaffold, thus leaving this goal for future
work in collaboration with investigators having access to more advanced FDM
machines. We have carried out some experiments with biodegradable filaments
and now the next step in building 3-D scaffolds could prove to be an optimal
combination of desired properties of scaffold in terms of material,
manufacturing, biocompatibility and cell proliferation.
FDM manufacturing of the scaffold was not carried out as part of this thesis,
because the feeding mechanism of our FDM machine was not suitable for soft
material, nor was the filament diameter adapted. These issues are now
addressed as part of another Ph.D. project in our department.
The reason for choosing a suture material was the high rate of cell adhesion in
vivo on the suture surface, which provides wound healing at the site of suturing
after surgical procedures. In vitro, the suture has also demonstrated a good
112
degree of cell adhesion, thus promising a good rate of scaffold proliferation
prior to implantation of the graft.
2.16. Conclusions There are many ways to combat portal hypertension and its complications. All
have their advantages and disadvantages, and there are examples in which they
are more or less effective than other methods. Thus, all treatment methods need
to be available in any medical centre undertaking hepatic surgery, and in each
individual case the appropriate approach needs to be chosen. In some cases a
combination of different treatments might be the correct approach, and the
search for novel methods needs to continue.
This project started as an investigation of the development of a micro-axial
pump for liver perfusion and was changed due to the lack of materials to
develop the device, the difficulty in driving and monitoring the pump over a
long period of time and the damage it can cause to the blood flow.
Thrombosis and other occlusions in the portal vein change the flow dynamics
and the prognosis of the disease. The modelling done in this thesis to compare
flow behaviour in a simplified portal vein model with and without blockages
can be used to further understand the impact this complication has on the
outcome for the patient.
In this chapter a review of different scaffold materials and methods for cell
seeding have been discussed. Pulsatile flow was discussed in general and in
relation to its application to blood flows in the portal vein. The importance of
pulsatility flow for culturing the new vessel and other current methods for liver
disease treatment were briefly presented. Experimental work indicates that more
studies are needed to investigate the relationship between scaffold porosity and
cell adhesion in both steady and pulsatile flow conditions. The review carried
out in this chapter presents the basis and ideas for future work.
113
CHAPTER 3
Commonly used methods and parameters in blood flow
modelling and thesis specific used theoretical and experimental
methods
3.1 Introduction In this chapter the different methods for blood flow measurements and Laser
Doppler Anemometry (LDA) are presented. Flow properties and behaviour of
the blood as well as common assumptions and simplifications used to describe
its behaviour, both theoretically and experimentally, are given and the
principles of the Computational Flow Dynamics (CFD) model used in this
thesis are described. Recommendations for future research are given at the end
of this chapter based on observation made during the LDA measurements or the
computer model results.
3.1.1. Methods for measurement of portal blood flow
Portal blood flow is not easy to measure and in most cases requires an invasive
procedure (either adding substances to the blood or biopsy). However, non-
invasive methods are being developed and tested.
Doppler ultrasound (DUS) is used to confirm the normal hepatic structure,
presence of transformations in the portal vein and the patency of the left
intrahepatic portal system in patients with extra hepatic portal hypertension due
to idiopathic portal vein thrombosis (St Moravec 1987).
To determine the patency of the portal vein a color-flow DUS can be used
before a shunt is performed. When there are vague findings on color-flow DUS,
venous-phase visceral angiographies should be undertaken (In Seok Kang
2002).
Nowadays, there are several methodologies for diagnosing natural shunts
and liver blood flow (Durst et al. 1976), although many methods, such as Au
colloid and 32P-chromophosphate, are no longer in use. Some examples are
given below:
114
• Clearance methods measure the rate of disappearance of radioactive
indicators from the system. Some of them are briefly outlined below.
• Intake methods are based on the ability of the liver to remove the
indicator (radioactive elements, bromsulphthalein etc.) Some of the
most common ones are described below.
• Portal or transcolonic scintigraphy – intake of radiochemicals from
the intestinal tract (usually of technetium 99m pertechnetate). The
method involves monitoring whether the chemicals bypass the liver
and are detected in the heart first, which shows that the blood has
been shunted away from the liver. It does not provide information on
the location of the shunt, but the degree of shunting can be
determined with good accuracy.
• Contrast radiography using a marker dye injected into a vein
draining the intestine is accompanied by radiographs and allows
good visualization of the portal vein and shunts, but is invasive and
is only to be used to assist surgery.
• Contrast-enhanced agent detection imaging (Youk et al. 2003) has
been shown as useful and as effective as helical computed
tomography for evaluating the therapeutic effects of interventional
therapeutic procedures for malignant hepatic masses. It is usually
used to supplement ultrasound or Doppler measurements.
• Continuous indocyanine green infusion method is one of the
invasive methods (Miyamoto et al. 2003) evaluating the metabolic
activity of the hepatocytes.
• Ultrasonography is a non-invasive approach used for detection of
intrahepatic shunts. Doppler ultrasonography is used intraoperative
for shunts and flow visualization (Sekido et al. 2002). Using this
method, the cross-sectional area of the vessel provides information
on the size and quantity of intimal thickness next to the wall. It can
be used for the visualization of vessels, shunts or even spontaneous
collaterals.
• Doppler ultrasound is another non-invasive method and is used
primarily for detection of extrahepatic shunts. It has also been used
115
to investigate the turbulence intensity in the carotid bifurcation in
vitro, showing promising results (Glanemann et al. 2001; Kok et al.
1999; Levick 1995; Poepping et al. 2004). Pulsed Doppler
ultrasonography can be used to measure velocity waveforms in the
portal vein branches (Kiserud et al. 2003).
• Colour Doppler ultrasonography has been used for non-invasive
measurements of renal resistance and pulsatility (Koda et al. 2000).
• Contrast harmonic ultrasound has been successfully used in dogs for
determination of macrovascular and perfusion patterns (Salwei et al.
2003).
• Contrast-enhanced computer tomography (Miyamoto et al. 2003).
• Liver biopsy is an invasive method used when shunts cannot be
detected or there are multiple extrahepatic shunts.
• Ultrasound guided biopsy is used for visualization in patients with
portal vein thrombosis (Spircher et al. 2003) for detection of
hepatocellular carcinoma. Other applications of this method are still
to be tested.
• Doppler sonography is non-invasive and can be transcutaneous or
intravascular (Glanemann et al. 2001; Haag et a;. 1999; Marsutani et
al. 2003) and specially used for pulsatile flow (Hűbner et al. 2000).
As a cross-sectional imaging technique, it is very useful for
demonstrating aneurysms of the portal venous system and bland or
neoplastic portal vein thrombosis (Gallego et al. 2002). This method
is also effective for long-term TIPS follow-ups ( Žižka et al. 2000).
• Doppler sonography aided by Levovist for improvement of the
diagnostic efficiency (Drelich-Zbroja et al. 2003; Fischer et al.
1998).
• Colour velocity imaging quantification takes into account the blood
flow profile and might have advantages over conventional Doppler
flow measurements (Kawasaki et al. 1999).
• Helical computed tomography has shown to be a useful, less
invasive method for 3-D anatomic analysis of large intrahepatic
portosystemic venous shunts (Nagafuchi et al. 19996).
116
• Laser Doppler light scattering instruments for measurement of flow
within blood vessels in vivo have been under extensive development
since the early 1970’s (Bonner and Nossal 1981).
• Measurements of blood flow using pressure, or flow parameters of
other nearby vessels (Bolondi et al. 1990; Burns and Jaffe 1985;
Deplano et al. 1999; Gibson et al. 1993). Some of those methods are
presented later in this chapter in 3.1.3. Flow measurements are based
on pressure gradients, flow in other blood vessels, or numerical
estimation.
3.1.2. Generic flow measurements
3.1.2.1. Electromagnetic Flowmeters
Magnetic flowmeters invasively measure the blood velocity and are known to
have good linearity, direction sensitivity, capability of monitoring pulsatile flow
and capability of monitoring flow in intact blood vessels. The basic operational
principles of this method involve an electromagnetic field around a vessel and
the use of the blood’s properties as a conductor of electricity. This technique
was used by Schenk et al. (1962) to measure the hepatic flow, which showed
that on average 26% of the total hepatic blood is being supplied by the hepatic
artery (with upper limit of values 50%). Another use of this method has been
the measurements of shunt operations. The flow in this case can indicate the
success of the anastomosis. One of the negatives of this method is the invasive
approach as the probe needs to be around the vessel and an in vivo calibration is
required; this causes added risk and discomfort to the patients.
3.1.2.2. Ultrasonic Methods
The Ultrasonic Doppler Technique was first implemented for blood flowmeters
by Satomura and Kaneko (1961) so that an ultrasound beam was directed onto
the blood flow and the frequency changes produced in the backscattered
radiation were monitored. The Doppler shift is a term used to describe the
frequency shift when the sound source moves relative to the observer, i.e. to
higher frequency when approaching the observer and to lower frequency as it
117
moves away. The absorption coefficient α of most biological tissues is
approximately proportional to frequency f.
In blood, the ratio α / f increases with frequency up to about 10MHz (Rowan
1981). The error in the measurement can occur due to unknown dimension and
orientation of the blood vessel and the shape of the velocity profile. In practice
the angle between the blood vessel and the beam used is 30-45O although the
ideal angle is 0O, i.e. the ultrasound beam is parallel to the direction of the flow.
This is not possible for a non-invasive technique and the compromise increased
angle is used. One of the main practical limitations of this method is not being
able to identify the direction of the flow, thus, a Directional Doppler instrument
has been designed. For clinical use the pulsatile index is defined as PI = peak-
to-peak velocity /mean velocity; as this formula is independent of probe-to-
vessel angle. This index (PI) increases from the aorta to the small vessels and
can be used with accuracy even when there are vessel occlusions up to 50%
(and some authors claim more than 50%). The two most frequently used
instruments are the continuous wave Doppler and the pulsed Doppler. The first
records not only the target but all moving structures in the way of the beam and
is used to create an outline of the blood vessels, whereas the pulsed Doppler
(Kiserud et al. 2003) is used to create a two-dimensional section of the vessel,
i.e. vessel diameter as well as the blood flow can be measured.
Some researchers have used measurements of the hepatic arterial acceleration
index (the dependency between the total cross-sectional area and the early
systolic acceleration) for non-invasive evaluation of portal hypertension (Tasu
et al. 2002).
In flowing blood, when measurements are made with a Doppler ultrasound, the
Doppler-shift echo is given by the red blood cells, and its size is much smaller
than the specular reflection given by solid tissue interfaces (Burns and Jaffe
1985).
Other applications of Doppler methods are presented later in this chapter.
118
3.1.2.3. Electrical Impedance Techniques
These plethysmography methods use the change in the impedance in the blood
vessel due to the systolic and diastolic cardiac cycle. Venous occlusion via cuff
is the principle for blood flow measurements; hence these are non-invasive
measurement techniques. The main applications of this technique are deep vein
thrombosis monitoring, monitoring stroke volume (pulse) and blood flow
through the brain.
3.1.2.4. Tracer Techniques
Metabolically inert tracers, which have the ability to diffuse rapidly between
blood and tissue, are used in these methods. They are based on the principle of
conservation of mass, so that the quantity of the substance taken up by tissue for
one unit time is equal to the quantity of arterial blood brought to the tissue
minus the quantity of venous blood carried away from the tissue. One of the
most popular for liver blood flow measurement is the 133Xe injection technique.
It consists of external monitoring of the clearance of gamma ray activity after
the injection of 133Xe. This method does provide information on different
segments of the liver and the mean blood flow within those segments, but
provides no information on the overall variability of blood flow in the whole
organ.
3.1.3. Flow measurements based on pressure gradients, flow in other
blood vessels, or numerical estimation.
In this part of the thesis measurement techniques are discussed with the
recognition that they have been used in studies of the portal blood flow and
most likely will be used in the future for the establishment of a unified
mathematical formula for portal flow calculation. The different methods for
measurement and visualisation are discussed further in this chapter.
One important factor affecting the accuracy of any blood flow measurement
would be the impact of gravity on flow, and in relation to this, the posture of the
patient during the measurements.
119
Due to the collapsible nature of veins, even small changes in venous pressure
will have large effects on the venous blood flow (Levick 1995).
Portal hypertension combined with ascites, variceal bleeding, esophageal
varices or splenomegaly is making it more difficult for diagnostic methods to
measure portal vein flow.
Portal circulation is regulated by three factors - blood pressure, blood flow and
vascular resistance - as are other splanchnic circulations. To evaluate the degree
of abnormality in the portal circulation, the portal venous pressure and the
portal blood flow have to be measured (Moriyasu et al. 1986; Richardson and
Withrington 1981, 1978; Sherlock 1974; Ueda et al. 1971). The values of both
volume and pressure are regulated and affected by several other factors. Such
factors are, for instance, intrahepatic changes like fibrosis that can elevate the
portal venous pressure and decrease the portal blood flow (Sherlock 1974).
Splenic enlargement can increase the portal blood flow by increasing the splenic
blood flow and, in turn, elevate the portal venous pressure (Ueda et al. 1971).
The formation of extrahepatic porto-systemic collateral pathways can lower the
portal venous pressure and decrease the inflow volume of portal blood.
Intrahepatic shunts between the portal and hepatic veins can lower the
intrahepatic resistance of the portal blood flow and decrease the portal venous
pressure (Moriyasu et al.1986). So the factors that affect the portal venous
pressure and the volume of the portal blood flow are complicated. Therefore the
data for both pressure and flow volume should be treated with care, as it is
difficult to completely comprehend the hemodynamics of portal hypertension
simply by measuring the portal venous pressure or the portal blood flow.
However, by focusing on the perfusion of the portal blood flow through the
liver, that is, the inflow to the liver through the hepatic portal vein and the
outflow from the liver through the hepatic vein, there are only three factors
influencing the hemodynamics that need to be considered: volume of portal
blood flow, portal perfusion pressure, and portal vascular resistance (Moriyasu
et al. 1986; Richardson and Withrington 1981, 1978).
120
3.1.3.1. Portal vein blood flow measurement based on pressure gradient
between portal and hepatic veins
One of the most commonly seen problems with those approaches is the error
from the zero point, which is usually external to the vein. There are suggestions
that using an internal zero point will eliminate most of the source of error in
percutaneous transhepatic measurement of portal and hepatic veins pressure
gradient (Gibson et al. 1993). As use of wedged pressure is a well-established
technique, it needs to be noted that it does not always reflect the portal pressure
and depends on the underlying disease for each patient (Gibson et al. 1993).
In patients with thrombosis without cirrhosis the spleen-to-wedged hepatic
venous pressure (WHVP) gradient has shown to be more than double, and the
splenic-to-free liver vein pressure gradient more than 50% increased when
compared to patients with cirrhosis but without thrombosis (Keiding et al.
2004).
3.1.3.2. Measurements based on pressure drop within the blood vessel
In an idealized vessel, flowing blood would speed up in an area of narrowing.
Then, the velocity change could be used to determine the magnitude of the
pressure drop. In a blood vessel, as in any tube, there are cohesive forces
between the nearby laminae and the moving blood, so there is some resistance
to acceleration of a single stream within the whole lumen vessel, i.e. the so
called drag of viscous friction (Burns and Jaffe 1985). The drag is dependent on
the viscosity of the blood and the size of the narrowing of the blood vessel. The
computer model presented in this thesis has tried to take this into account and
simulations of the same flow with variations in the viscosity of the fluid only
were carried out and are presented in a later chapter of this work.
The pressure drop in this method can be described by the following equation:
Pressure drop = kinetic energy gain (due to acceleration) + viscous loss
(due to friction) + inertial energy gain (due to changing flow rate)
(3.1.3.2)
This is the Bernoulli equation in its physical representation as given by Burns et
al. (1985).
121
3.1.3.3. Volume flow measurements
There are several types of measurements in this category. To use the velocity
profile method, the velocity profile has to be known and needs to be constant
during a cardiac cycle. For the cross sectional method, the area needs to be
assumed circular and has to be known.
Portal blood flow mean velocity was calculated according to Moriyasu et al.
(1986) as maximum velocity times 0.57 and expressed as centimetres per
second. Portal blood flow volume was calculated as portal blood flow mean
velocity times cross-sectional area and expressed as millilitres per minute. This
estimate is known to be quite inaccurate, but it is the most widely used method
to measure portal blood flow non-invasively. Resistance indexes from the left
(L) and right (R) branches of the hepatic artery were similar in both controls
and cirrhotics (cirrhotics: L-PI, 1.35 ± 0.37; L-RI, 0.72 ± 0.08; R-PI, 1.27 ±
0.42; R-RI, 0.69 ± 0.09). Portal blood flow was not significantly different in
cirrhotic patients and controls (866 ± 363 vs. 948 ± 303 ml min-1; P = 0.27)
(Shiomi et al. 2000).
This brief review of methods is given to help better understand the complexity
of blood flow measurements and not to critically examine the existing methods.
3.1.3.4. Measurements of Portal Vascular Resistance
According to a study (Sacerdoti et al. 1995) involving 31 controls and 171
cirrhotic patients with (n=13) and without (n=158) portal vein thrombosis, who
were measured using duplex Doppler ultrasonography (DDU), hepatic arterial
resistance indexes increased in cirrhosis, particularly with portal vein
thrombosis. To the authors, the pathophysiology of the increase in hepatic
arterial resistance seemed to be parallel to that of portal resistance. Peak
systolic, end diastolic, and temporal mean velocity were determined and the
pulsatility index (PI) and the resistive index (RI) were calculated according to
the following formulas:
PI= (Peak Systolic – End Diastolic Velocity) / Mean Velocity, and
RI = (Peak Systolic – End Diastolic Velocity) / Peak Systolic Velocity
122
Measuring the portal vascular resistance in patients with portal hypertension
(Moriyasu et al. 1986) has shown that in patients with cirrhosis, both the portal
venous pressure and the portal perfusion pressure were elevated, the wedged
hepatic venous pressure was also higher than normal, and there was only a
slight presinusoidal pressure difference. The reduction of portal blood flow was
insignificant, but the portal vascular resistance was five times as great as that in
the control group (Moriyasu et al. 1986).
Richardson and Withrington (1981) have also estimated the vascular resistance
of the normal human liver assuming that the portal venous pressure is between 5
and 10 mmHg, and the hepatic venous pressure is between 1 and 2 mmHg.
Increased resistance to portal blood flow is the main indicator of portal
hypertension and is mainly determined by the morphological changes occurring
in chronic liver disease (Bosh and Garsia-Pagan 2000).
3.1.3.5. Measurements of the Hepatic and Portal Venous Pressure
Under normal conditions the portal vein pressure is around 9mmHg, while the
pressure in the hepatic vein is close to 0mmHg, thus the pressure difference is
around 9mmHg. This pressure difference increases with the increase of the
severity of portal hypertension.
One of the fundamental studies that compared free portal venous pressure and
wedged hepatic venous pressure was carried out by Viallet et al. (1970), which
recognised the importance of evaluating the porto-hepatic gradient. Good
correlation between Wedged Hepatic Venous Pressure (WHVP) values and free
portal vein pressure (FPVP) values was seen in all 43 patients in that study and
the maximum difference between these two parameters was 4mm Hg (WHVP
ranged from 9.5 to 40 mm Hg and FPVP ranged from 9 to 38mm Hg). The
difference between FPVP and free hepatic venous pressure (FHVP) (porto-
hepatic gradient) was calculated (range between 3 and 24mmHg) and used as an
index of portal hypertension (Viallet et al. 1970).
Hepatic venous pressure gradient (HVPG), the difference between the WHVP
and FHVP, can be used to describe the severity of the portal hypertension –
123
mild when HPVG=12mmHg and severe when HPVG>12mmHg (Tasu et al.
2002).
Another study of cirrhosis due to hepatitis C infection (Deplano et al. 1999)
revealed that the difference between WHVP and portal vein pressure (PVP) was
inversely related to the portal flow velocity and directly related to the portal
vascular resistance. Predominantly left portosystemic collateral blood flow can
be observed when WHVP is higher than PVP Deplano et al. 1999).
3.1.3.6. Relationship between vessel diameter and velocity
In chronic liver diseases, especially in liver diseases accompanying portal
hypertension, the diameter of the vessel increases and the flow velocity
decreases. From the experimental study in (Moriyasu et al. 1986), in which the
authors had varied the diameter of the vessel and the viscosity of the blood, it
was concluded that the ratio between the mean and the maximum velocities was
constant as long as the flow was not turbulent.
3.1.3.7. Measurement of Portal blood flow
In the early 1970’s, a study for assessing the portal blood flow (PVF) carried
out by Sovak et al. (1999) showed that as the disease progresses, PVF may
become hepatofugal due to the high hepatic artery flow and hepatoportal
shunting.
Strandell et al., (1973) undertaking measurements in 8 conscious patients, noted
that portal vein blood flow varied from 0.82 to 2 litres/min-1 and hepatic artery
flow from 15 to 56% of total hepatic flow. It is important to note that two
patients with identical total hepatic blood flow may have a markedly different
distribution of flow between the portal vein and hepatic artery.
In 2001 a study by Bolognesi et al. (2001) created a formula for prediction of
the grade of portal hypertension as follows:
[splenic pulsatility index * 0.066 - 0.044] * portal blood flow (3.1.3.7)
This group showed good accuracy in the study of 19 initial patients and 21
further patients.
124
A study performed by Richter et al. (2000) in 36 cirrhotic rats showed that flow
reduction of the hepatic artery did not influence portal venous blood flow.
Normally, the portal vein provides the major blood supply of oxygen to the liver
(Mathie andAlexander 1990). In cirrhosis, the change of the ratio of portal
venous to hepatic arterial blood flow in favour of the hepatic artery may sustain
oxygen delivery and exert a protective effect on organ function and integrity
(Mathie and Blumgart 1983).
3.1.4. Doppler flowmetry
Doppler flowmetry is useful in assessing the direction, presence and
characteristics of blood flow in hepatic vessels (Bolondi et al. 1990).
Determining the presence of blood flow is the easiest Doppler finding, and the
absence of Doppler signal from the portal vein confirms the presence of
thrombosis.
Because the portal vein is large in size and has around 3-4 centimetres of
straight course (Bolondi et al. 1990), measurements using Doppler are quite
accurate, provided they are done in small time intervals (4-6 seconds) and are
repeated to minimize error.
3.2. Laser Doppler Anemometer
3.2.1. Principle of Laser Doppler Anemometry
In this work Laser Doppler Anemometer (LDA) will be used as a tool for fluid
dynamic investigations in the liquid (blood simulation). Because of its non-
intrusive principle and directional sensitivity, LDA is very suitable for flow
measurements in applications with reversing flow, or in biological systems
where physical sensors are difficult or impossible to use. LDA requires tracer
particles in the flow and transparency of the shunt. The particular advantages of
this non-invasive method are:
• velocity range from 0 to supersonic
• high spatial and temporal resolution
• no need for calibration
125
• the ability to measure in reversing flows (Coherent Scientific and
Dantec Dynamics).
LDA is becoming the most used method for measurements of blood flow and
optimisation of artificial hearts and heart valve function due to its functionality.
Two-dimensional LDA is the most common method used even for complicated
structures and flow profiles like the one across heart valves. Many studies by
Grigioni, Barbaro, Daniele and D’Avenio published between 1997 and 2000
have used two-dimensional LDA and have shown the advantages of the method.
The basic configuration of any LDA consists of a continuous wave laser,
transmitting optics (including a beam splitter and a focusing lens), receiving
optics (comprising of a focusing lens and a photo-detector), a signal conditioner
and a signal processor (Figures 3.2.1.1 and 3.2.1.2). Most advanced systems
may also include traverse systems and angular encoders.
The principle of LDA involves division of the laser beam and then intersecting
the two beams using a focusing lens. Tracer particles in the flow scatter light
which gets picked up by a receiver lens and then focused onto a photo-detector.
The noise from ambient light and from other wavelengths is removed by an
interference filter mounted before the photo-detector, which passes only the
required wavelength to the photo-detector.
“The scattered light contains a Doppler shift (Doppler frequency fD), which is
proportional to the velocity component perpendicular to the bisector of the two
laser beams” (corresponds to the x axis shown in the probe volume in Figure
3.2.1.2) (Coherent Scientific and Dantec Dynamics).
The other very popular technique for flow measurements is Particle Image
Velocimetry (PIV) which allows all three velocity components to be recorded
and thus provides 3D velocity vectors for the whole area instantaneously.
Although this technique has similar advantages as LDA it was not used in this
study due to unavailability of the equipment.
126
Figure 3.2.1.1. Experimental setting of LDA – basic operational principles (with permission from Coherent Scientific and Dantec Dynamics, reference (Coherent Scientific and Dantec Dynamics))
Figure 3.2.1.2. The probe and the probe volume (with permission from Coherent Scientific and Dantec Dynamics, reference (Coherent Scientific and Dantec Dynamics))
The probe volume is typically a few millimetres long. The light intensity is
altered due to interference between the laser beams, which produces fringes
(parallel planes of high light intensity) (Coherent Scientific and Dantec
127
Dynamics). The fringe distance (df) is defined by the wavelength of the laser
light and the angle between the beams by the following formula:
df = λ ⁄ 2 sin ( θ/ 2) (3.2.1.)
Depending on the local light intensity, different particle passages scatter light
differently and thus provide information on the flow velocity when passing
through the probe volume.
The edge (fringe) spacing df provides information about the distance travelled
by the particle and the Doppler frequency fD provides information about the
time: t = 1/fD and velocity equals distance divided by time (V = df /fD)
Figure 3.2.1.3. Doppler frequency to velocity transfer function for a
frequency shifted LDA system (Coherent Scientific and Dantec Dynamics)
The frequency shift obtained by the Bragg cell (glass crystal with a vibrating
piezo-crystal attached, used as the beam splitter) makes the fringe pattern move
at a constant velocity (Coherent Scientific and Dantec Dynamics). The particles
which are not moving will generate a signal of the shift frequency fshift, and the
velocities Vpos and Vneg will generate signal frequencies fpos and fneg,
respectively. LDA systems which do not possess frequency shift cannot
differentiate between positive and negative flow direction and cannot measure
zero velocity (Coherent Scientific and Dantec Dynamics).
128
LDA studies of the flow behaviour and velocity in varicose vein models
(Schalina and Liepsch 2001; St Moravec 1987) and flow measurements (Durst
et al. 1976; Liepsch 1978; Mori 1989, Chapter 19) have been very popular in
the last three decades.
The methodology for flow visualization used by L. Schalina and D. Liepsch
(2001) is simple and effective for the purpose of studying flow behaviour in
model veins. Part of the future studies proposed in this thesis is to use this
technique to model the portal vein flow and adjust the flow parameters in the
computer model as appropriate. The method described involves four steps:
• Obtain the vein from the patient (in the case of the portal vein, this
would most likely be during liver transplantation),
• Filling the vein from both ends by injecting silicone rubber mixed with
a hardener
• Corrode the vessel wall by dissolving the tissue in 30% potassium
hydroxide leaving a silicon cast
• Use this cast for the preparation of translucent rigid vein models with
polyester resin and for the preparation of translucent elastic silicone
models
This model (see Figure 3.2.2.1. below) can then be used for LDA measurements
of flow patterns (velocity fluctuations, shear stress on the interior wall of the
vein, non-Newtonian blood flow) using coloured dyes.
Another possible model constructed from Sylgard 184 silicone
elastomer (Hinds et al. 2001) could also be used for LDA measurements as it
has optical clarity.
Properties/ requirements of the fluid:
• The fluid used for the LDA measurements needs to have similar flow
properties to that of blood (refraction index must be same as the one of
the model n = 1.41)
• A suitable fluid could be any of the following solutions: 52% (v/v)
glycerine with a density of 1.150 kg/m3 and titanium dioxide particles
for Newtonian flow (Shalina and Liepschb 2001); 58% glycerol with a
density of 1.14gr/cm3 (Hinds et al. 2001); or 32.7% (v/v) glycerol with
129
5.2% (w/v) CaCl2 in distilled water with a density of 1.1gr/cm3 and
3.625 centipoise dynamic viscosity at room temperature (Mori 1989,
Chapter 19)
• Must have added particles in the flow to allow measurements (particles
refract the laser beam)
• Must be transparent to allow the laser beam to pass through it
uninterrupted.
The usual seeding particles used for LDA measurements are polyamide seeding
particles (round but not exactly spherical) or hollow glass spheres and silver-
coated glass spheres (the silver coating increases the reflection of the laser
beam). There are many other types of particles (fluorescent polymer and
others) that can be used to perform the measurements.
In a study of blood flow in stenosis, for example the liquid used was 58%
glycerol in water (refractive index 1.412) with suspension of 10-20μm
polymethylmethacrylate particles (Hinds et al. 2001).
In the LDA measurements done as part of the research of this thesis, Meta DC
Coated Particles (Model 10037) were suspended in a glycerol solution to enable
refraction of the laser beam and allow flow measurements.
3.2.2. Models used for LDA
LDA or other visualization methods are usually used to represent flow
behaviour in arteries (Strackee and Westerhof 1993). Normally, the models
have to be enlarged to enable the measurements, and thus the Navier-Stokes
equation has to be transformed to take this into account. In the experiments
performed as part of this thesis no scaling was done and no adjustment to the
equation was needed.
The following three figures depict a three-stage model creation technique used
to create a varicose vein model for LDA measurements. It could be applied to
the portal vein in the future, the subject of this thesis.
130
Figure 3.2.2.1. Silicone cast of varicose vein (stage 1) (Schalina and Liepsch
2001)
Figure 3.2.2.2. Rigid polyester resin model of the varicose vein (stage 2)
(Schalina and Liepsch 2001)
Figure 3.2.2.3. Elastic silicone rubber model based on the previous two
models (stage 3) (Schalina and Liepsch 2001)
131
The final model was installed in a circulatory system using pressure valves,
piston pump and dyes (Schalina and Liepsch 2001). This model can be used for
3D visualizing of flow via LDA. A similar approach can be used to visualize the
portal vein, but due to equipment and material limitations, the model used in
this thesis was made out of glass, and hence was constructed of rigid, not elastic
walls. This limitation is acknowledged and in future work the model needs to be
compared to one made either using the above method or another method that
allows elasticity of the model vessel wall.
For example, a bifurcation as shown below can be used for 3D flow
visualization:
Figure 3.2.2.4. Representation of bi (tri) furcation (Dinnar and Raton 1981)
The model used as part of the research in this thesis (Figure 3.2.2.5 below), as
mentioned above, was glass-blown as two different variations – one without any
blockages, and the other one with three blockages, representing obstructions
respectively in the trunk and one in each or the main branches (left and right).
132
Figure 3.2.2.5. Comparison between the two glass models and scale
The results from the LDA measurements using the above models of the portal
vein are presented in the Results Chapter of this thesis. A 2-dimentional Laser
Doppler Anemometer was used to measure the velocity of the flow at a number
of points within the vessel model. Once the Z plane was known from the focal
point, having the X and Y velocities were sufficient to measure the flow in the
simplified model.
3.3. Computational Fluid Dynamics (CFD) Modelling Computational fluid dynamics (CFD) has been widely used for the
characterisation and visualization of flow field as well as obtaining data on wall
shear stresses (Hinds et al. 2001; Niu et al. 2002; Xu et al. 1999; Zhao et al.
2000). CFD can use in vivo measured parameters, such as velocity and flow, as
boundary conditions (Song et al. 2000). Most beneficial for the development of
a model to be used in clinical practice is validation of the CFD results with
experimental particle tracking techniques. It is one of the aims of this thesis to
compare experimental values to CFD modelled values. In this part of the thesis
133
discussion on and description of the CFD methods and proposed model are
explained. In this research the physical models were created based on the
dimensions of the CFD models for validation purposes.
The FLUENT software package has been used for a wide range of applications
to model the physiology of the human body (Fluent Europe Ltd 2002). CFD
uses numerical methods to solve the equation governing the fluid flow by
splitting the analyzed domain into small volumes and elements and a set of
partial differential equations is solved for each one of them. FLUENT allows
simple adjustment of the flow parameters and can deal with complicated models
involving a two-phase flow and provides three-dimensional visualization of the
model.
FLUENT was used for simulating flow through a pulmonary artery to predict
and visualize areas where clotting and aneurysms were most likely to occur,
based on non-invasive MRI carried out by a group at Sheffield University,
England (Thilmany 2003) and other groups (Moore et al. 1998). This group has
also shown the possibility of studying stent size and orientation by modelling
the stent location and design using this computer method. FLUENT has been
used for turbulent blood flow simulations (Varghese and Frankel 2003) and for
many non-biological simulations (like turbo machinery, wind and aircraft
models).
Some studies using a 3-D model of the aorta (Yamaguchi et al. 2002) have used
another commercially available software package (Software Cradle, Japan)
called SCRYU version 1.4, which allows for the solving of unsteady Navier-
Stokes equations for incompressible flow (high Reynolds number, no-slip at the
wall, velocity perpendicular to the aortic cross section, and zero velocity and
pressure at the outlets). An example of this work is shown below.
134
Figure 3.3. Wall shear stress and flow streamline patterns at different cross
sections of the aorta using CFD analysis, where the model is based on
magnetic resonance imaging (MRI) data (Yamaguchi et al. 2002).
Similar studies were carried out by Xu et al. (1999) in the human carotid and
aortic bifurcations.
FIDAP (Fluid Dynamics International, IL USA), a general-purpose code, has
been used for computing velocity, pressure and wall shear stress in coronary
artery bypass grafts (Song et al. 2000). The assumptions made were common to
other studies, and include steady, symmetric, incompressible, homogeneous
Newtonian flow through round diameter vessels (and all vessels with the same
diameter) with rigid walls.
ABAQUS and CFX4 (Long et al. 2000; Morsi et al. 2001; Starmans-Kool et al.
2002; Thilmany 2003; Xu et al. 1999; Zhao et al. 2000) were also successfully
used in modelling and simulation of arterial bifurcated blood flow. CFX4
135
provides the opportunity to simulate a moving grid of the finite element solid
mechanics multi-block structured grid created with ABAQUS. Recently,
FLUENT has been upgraded to provide similar moving grid simulations on
models created with ABAQUS or Gambit.
Materialise Mimics (Belgium) software, used for converting MRI slices into 3-
D solid models suitable for export into some of the most popular solid
mechanics modeling software (Thilmany 2003) is a good example of linkage
between non-invasive in vivo measurements and computer simulations on a
practical level.
Comparisons between 3-D ultrasound measurements in vivo have been
successfully used to generate realistic geometry, suitable for CFD simulations
(Augst et al. 2003), and multiple measurements have shown reasonable
reliability of these methods.
CFD has been used successfully in modelling post-stenotic flow in the aorta
(Niu et al. 2002), demonstrating its capabilities to simulate the wall shear stress
and model the unstable flow in that region of the blood vessel.
Even though most studies of blood flow assume Newtonian behavior (Bonert et
al. 2003; Finol and Amon 2002; Gurlek et al. 2002; Hinds et al. 2001; Marques
et al. 2003; Moore et al. 1998; Niu et al. 2002; Siro et al. 2002; Starmans-Kool
et al. 2002; Xu et al. 1999; Zhao et al. 2000), in this thesis non-Newtonian flow
has been modeled. Newtonian behavior of blood flow can be assumed for
simplicity in large vessels (aorta and large arteries) with high shear rate, which
is not the case in the portal vein. A comparison of results obtained with the
same model using parameters for Newtonian and non-Newtonian flow showed
some differences in the flow behaviour. These are presented in the Results
chapter of this thesis.
Fluid-solid mechanical interactions between the blood flow and the vessel wall
are difficult to study in vivo and computational, non-invasive methods may
provide the way to evaluate, predict and prevent fatal consequences of vascular
disease (Long et al. 2000; Yamaguchi et al. 2002).
136
3.3.1. Non-Newtonian flow
A large number of blood flow studies treat blood as non-Newtonian fluid (Fung
1993; Jalan et al. 2004; Petkova et al. 2003; Sugiura 1988; Zhang and Kuang
2000). Blood constitutive equations (BCEs) provide important information
about hemorheology and hemodynamics. There have been many theoretical
(Brunn and Vorwerk 1993; Krieger and Elrod 1953; Krieger and Maron 1954;
Mooney 1931) and practical ways to tackle the problem (model the blood flow).
BCEs are related to the mechanical characteristics of blood. Different equations
have been used to solve the problem - Casson equation (Aroesty and Gross
1972; Bate 1977), Walburn equation (Easthrope 1980; Walburn and Schneck
1976), Weaver equation (Suguira 1988), K-L equation (Wang and Stolz 1994),
Bi-exponent equation and Quemada equation (Zhang and Kuang 2000). They
can be divided in two different types – Casson and Power-law. In this thesis the
non-Newtonian Power-law equations have been used for simplicity, although
the Casson model can also be applied.
A new technology/program has been developed at the Department of Aerospace
Engineering, University of Bristol, under the leadership of Dr. Chris Allen, to
predict unsteady flows. This CFD method is based on a moving mesh for
solving the flow past deforming shapes. So far it has only been tested for
Newtonian flow, but is under development for non-Newtonian flow (such as
blood). This technique, when available, could prove to be more effective and
time saving than the current CFD techniques, but until this has been tested the
well-known FLUENT operations will remain most commonly used. FLUENT is
used to model the blood flow in this thesis.
3.3.2. Numerical simulations and modelling
Sometimes, physical experiments are difficult to perform, or too expensive and
time-consuming. After the initial stage of parameter and grid formulation,
computer modelling is a reasonably rapid and inexpensive way to visualise flow
patterns.
137
For simplicity, the blood vessel wall is treated as a rigid structure (Bonert et al.
2003; Gurlek et al. 2002; Marques et al. 2003; Petkova et al. 2003; Siro et al.
2002; Starmans-Kool et al. 2002; Xu et al. 1999).
In some cases, based on numerical simulations, the shape of graft vessels can be
altered to improve hemodynamics. An example is the following structure (Siro
et al. 2002):
Figure 3.3.2.1. Normal and Protrusion model vessel of prosthetics graft
connection to a blood vessel (Siro et al. 2002).
The protrusion model shows lower variation in shear stress on the wall opposite
the bifurcation compared to the normal one; hence the intimal thickening is
suppressed in the protrusion type model (Siro et al. 2002).
Due to the wide variation in the structure, function and pathology of patients,
portal vein and blood flow, each computer model needs to take into account the
clinically measured parameters (derived from imaging), the patient disease
history, and other conditions the individual might suffer from (i.e. kidney
failure, elevated blood pressure, etc.). Computer modelling therefore has to be
carried out on a case-by-case basis with regards to each patient (Long et al.
2000; Starmans-Kool et al. 2002; Thilmany 2003; Xu et al. 1999; Yamaguchi et
al. 2002; Yedavalli et al. 2001). Until a system is developed that allows rapid
modelling (within minutes) and still be based on the individual patient,
computer models cannot be used for emergency procedures. With the
development of new software, the possibility of changing the model by simply
changing the input parameters may arise. So far, and from our experience, the
simulation requires re-drawing of the grid and mesh for each case, which is in
essence the most time-consuming part of the modelling process.
138
In order to minimise the cost of modelling, the use of ultrasound geometry
measurements and pulsed Doppler velocity measurement in vivo can be
successfully combined with CFD simulations (Starmans-Kool et al. 2002). This
equipment is more widely available than MRI and can be used by most medical
professionals. As with all the other methods for combining real-life
measurements and computer simulations, many assumptions need to be made to
allow for modelling.
The simulations and modelling can be validated either against clinical data or
against in vitro experiments. In this thesis, LDA was used to validate the
principle of the model created with GAMBIT and simulated using FLUENT.
3.3.3. Limitations of CFD and future work
Ideally, computer models would use in vivo measurements and geometries to
build the simulation. There is no disagreement in the literature on the benefit of
real-life geometry and hemodynamics parameters in correctly predicting the
flow behaviour in individual patients. Obtaining the data is not problematic, as
non-invasive techniques are advancing very rapidly and there is a variety
available. The limitation is in what to do with the data once available, i.e. how
to convert the data into a “readable” form for the computer software used for
the simulations. Some researchers have developed their own custom-written
computer package to transfer MRI, or MRA (magnetic resonance angiography)
data to CFD code (CFX4) (Xu et al. 1999); others have used specialised
commercially available software like Materialise Mimics (Belgium) for
converting MRI slices into a 3-D solid model suitable for export into some of
the most popular solid mechanics modelling software (Thilmany 2003). Most
modelling programs require code writing or drawing and data input directly into
the system, or imported data from limited sources (usually CAD (Augst et al.
2003)) to create the model (including grid and mesh generation, flow pattern
and shear stress distribution), or a combination of Matlab and FLUENT (Moore
et al. 1998). In this thesis the geometry was created using GAMBIT and was
139
imported into FLUENT for flow simulations. Details are presented in the results
chapter of this thesis.
As previously mentioned, computer simulations need to be carried out for
individual patients to both predict their flow pattern and aid understanding of
the relationships between vessel geometry and angles of bifurcation, velocity
profiles and wall shear stress in general (Xu et al. 1999). More studies are
needed and many research groups worldwide are currently working to provide
such information.
CFD provides the benefit of easy visualization of flow path and the possibility
to predict flow, pressure and velocity fields as a function of time and position
within the geometry (Song et al. 2000).
3.3.4. FLUENT model used in this thesis
In this thesis, a simple 3-D geometry was used, as shown below (Figs. 3.3.4 a)
and b), which has four branches with different flow rates. Table 3.3.4.1 shows
the dimensions of the geometry used in this model. At inlet (bottom) the
velocity is considered to be 0.07 m/s and operating pressure is 3922.66 Pascal
according to Tasu (2002). The velocity magnitude, pressure, and dimensions of
the geometry are an approximation from the majority of published values. This
model geometry was constructed by using GAMBIT 2.0.4 (FLUENT 6.0) and
the simulations were carried out using a super computer at VPAC (Victorian
Partnership of Advanced Computing), which takes up to 60 minutes to converge
using FLUENT 6.0. The flow properties of blood used in this study are given in
Table 3.3.4.2 (Oka 1980). The convergence criterion of reduction of residuals
by five orders of magnitude for continuity and three orders of magnitude for
other transport equations was used.
140
(a) (b) Figure 3.3.4. Grid for (a) normal model; (b) blocked model (Petkova et al.
2003)
Dimensions mm
Inlet diameter 10
First branching diameter 8.5
Outlet diameter 6.375
Total hight of the vessel 91
Table 3.3.4.1. Dimensions of the geometry used in this model (Petkova et al.
2003)
Power law index (n) 0.4851
Consistency index k (kg-s^n-2/m) 0.2073
Reference temperature (0K) 310
Minimum viscosity limit ηmax (kg/m-s) 0.00125
Maximum viscosity limit ηmin (kg/m-s) 0.003
Table 3.3.4.2. Non-Newtonian power law parameters used in this study
(Petkova et al. 2003)
3.4. Blood flow properties
3.4.1. Rheological properties of human blood
Under low shear stress the blood experiences shear thinning and aggregation of
red blood cells, while in high shear the deformation and separation of red blood
cells is present. The flow properties of blood are also dependent on the cell
141
concentration, coagulation, adhesion and oxygen concentration. In this respect,
Kang (2002) based his study on the view that blood is a suspension of red blood
cells.
There are many theories dealing with the changing behaviour of blood. In this
thesis those changes have not been taken into account, but these are to be
considered for future research. Some of these are Batchelor’s theory (helping
predict the effective viscosity of blood), Hinch and Leal’s theory (continuation
from previous theory dealing with spherical particles) and Keller and Skalak’s
theory (the axisymmetric flow and tank treading motion with or without
flipping of the red blood cells).
Human red blood cells have a sphericity index about 44% larger than the
minimum area required for the spherical shape (Galbraith et al. 1998). The
sphericity index is represented by:
S= (A/4π)1/2/ (3V/4π)1/3 (3.4.1.)
where A is the total surface and V is the volume of the cells, and for spheroidal
particles S=a/b=r , where a is the elongated diameter and b is the smaller
diameter of the red blood cell
S= 1.2 and
r= 0.25 for oblate spheroid and r= 6 for an prolate spheroid (Galbraith et al.
1998).
In uniaxial straining flow, each blood cell is deformed into a prolate shape,
while in biaxial straining flow, the cells have oblate shape. In axisymmetric
straining flow the membrane tension is assumed to be isotropic and the shape of
the cells is independent of the strain rate if the bending resistance of the cells
are neglected (In Seok Kang 2002). The importance of the deformation
characteristics of blood cells on the flow has been demonstrated in many studies
in the last over three decades (Nakamura and Sawada 1988; Pozrikidis 1990;
Richardson 1974) and will be the focus for future research.
The tables below (3.4.1.1. and 3.4.1.2.) show some of the difficulties
researchers face when trying to model blood flow, namely the discrepancies of
the blood parameters. They might be due to the conditions of the blood donor,
his/her diet, the equipment used, or even the skills of the laboratory technician.
142
Regardless of the reasons, the differences are quite significant, induce
discrepancies in modelling, and need to be specified by the researcher in detail
(i.e. it can not be sufficient to state that the blood used in the model is human
whole blood, but rather the cell count, cell size, density and viscosity for that
specific sample need to be specified).
Reference Leukocytes Erythrocytes Platelets Plasma
Diameter/
number per
mm3
Diameter/ number
per mm3
Diameter/
number per
mm3
% of
total
blood
(Arey 1957) 5000-
9000
8.5μm 5.5 million1
5 million2
2-3μ 200-
350,000
55
(Miller and
Leavell 1972)
7-
20μ
5000-
9000
7.7μm 5.5-7
million1 4.5-
6 million2
2-4μ 400,000 ~50
(Petrov 1994) 6000 4.5-5 million 3-
500,000
(Strackee and
Westerhof 1993)
6-15μ 8μm 3μ
(Jensen 1996) 9-25μ 8000 7μm 5 million 2-4μ 250-
500,000
(Smith and
Kampine 1984)
5000-
7000
8 μm 4.5-5.5
million
2-3μ 150-
300,000
55
(Jalan et al.
2004)
7-
22μ
4000-
11000
8 μm 4-6 million 2-4μ 250-
500,000
~55
(Encyclopaedia
Britannica 2005)
4500-
11000
7.8μm 5.2 million 150-
400,000
Table 3.4.1.1. Composition of human blood; 1 Man, 2 Woman
The viscosity of blood varies with the hematocrit, which is the percentage of the
total blood volume occupied by blood cells.
143
Reference Density Viscosity at 37oC Kinematic viscosity
(Pedley 1980) 1.05x103 kgm-3 0.004 kgm-1s-1 4x10-6 m2s-1
(Strackee and
Westerhof 1993)
1.05x103 kgm-3 3-4mNsm-2
(Jensen 1996) 1.06x103kg/m3 0.004 kgm-1s-1
(Jalan et al. 2004) 103kg/m3 3-4mNsm-2
Table 3.4.1.2. Blood properties according to the literature
The relationship between hematocrit and viscosity is non linear. At hematocrit
of 40% the relative viscosity is 4, and at hematocrit of 60% is 8 (Klabunde
2003). An increase of 10 in hematocrit above the level of 40 results in around
25% increase in relative viscosity, and an increase of 20 (hence hematocrit
reaching 60) would result in around 60% viscosity increase (Smith and
Kampine 1984). Another factor that has a major impact on the viscosity is
temperature – for every 1oC decrease in temperature the viscosity increases by
approximately 2% (Klabunde 2003; Smith and Kampine 1984). In very low
flow, the cell-to-cell and protein-to-cell adhesion increases and thus the blood
viscosity increases too.
Below (Table 3.4.1.3.) are some examples of the blood flow parameters in the
portal vein, and it can be noted that they are varying largely between different
studies, thus again highlighting the importance of quoting the patient specific
conditions and the measurement methods used when presenting the results.
Those three examples (Tables 3.4.1.1. to 3.4.1.3.) highlight the benefit of
individualising the model to fit each patient’s individual conditions, in terms of
accuracy and realistic representation.
144
Reference Portal
vein
pressure
PV
pressure
after
procedur
e
Pressure
gradient
Pressure
gradient
after
procedure
Portal
vein flow
velocity
Total PV
blood flow
(Oikawa
et al.
1998)
26-42
(mean
36.44)
cmH2O
- - - - -
(Ochs et
al. 1995)
40 ± 10
cmH2O
28 ± 7
cmH2O
30 ± 10
cmH2O
11 ± 6
cmH2O
- -
(Rőssle et
al. 1994)
- - 21.5 ± 5
mmHg
9.2 ± 4.1
mmHg
cm/sec
7.7 ± 4.8
before
shunting
19.7 ± 5.2
after
shunting
ml/min
800 ± 500
before
shunting
1900 ± 800
after
shunting
(Okazaki
et al.
1986)
- - - - 10.2 ± 3.5
cm/sec
579 ± 262
ml/min
(Zardi et
al. 2003)
- - - - 23.9cm/se
c
[29cm/sec
after
iloprost
treatment]
1824.6ml/m
in
[2294.4ml/
min after
iloprost
treatment]
Table 3.4.1.3. Portal vein flow
In healthy people the portal vein pressure is around 10-14cmH2O (8-10mmHg),
which increases in portal hypertension to 25-35cmH2O (20-25mmHg) (Smith
and Kampine 1984).
145
The perfusion pressure, vascular resistance and blood viscosity determine organ
blood flow. In the case of a rigid tube the changes in apparent viscosity would
lead to parallel changes in resistance and a reciprocal change in flow (Chen et
al. 1989). In vivo changes in viscosity and flow may be compensated by
autoregulatory change in vascular geometry (Dalinghaus et al. 1994). While this
study has not aimed to measure blood properties, it is the intentions of future
research to have those measured and the model adjusted for the specific
conditions. Blood viscoelasticity, for example, is mainly represented by the
behaviour of red blood cells, thus further studies are needed taking this effect in
consideration.
Studies have shown that blood viscosity and elasticity change during and after
coronary artery bypass grafting (Undar and Vaughn 2002). This is an area
where research is starting to become more intense and the results of those
scientific efforts would greatly benefit the model proposed in this thesis.
One of the new methods for blood plasma viscosity measurements is fluorescent
molecular rotors (Haidekker et al. 2002). While this method is accurate, it
cannot be used for whole blood measurements. For these measurements, a novel
method is the use of a capillary viscometer with a mass-detecting sensor (Shin
and Keum 2002) over a range of shear rates without using anticoagulants
(measurements need to be completed within three minutes). The major
advantage of this method is the accuracy that is provided by using unaltered
whole blood.
The viscosity of plasma is 1.2mPa at 37oC and its behaviour is Newtonian
(Fung 1993; Jalan et al. 2004; Levick 1995; Petrov 1994; Strackee and
Westerhof 1993).
The number of blood cells determines the viscous behaviour of the blood.
Normally, there are 4.5-5 x 106 erythrocytes, 6 x 103 leukocytes, and 3-5 x 105
platelets per mm3 of blood (Petrov 1994), although there is some discrepancy in
these in the literature. For example, in Y.C. Fung’s book Biomechanics (1993)
the number of cells is given as 5-8 x 103 per mm3 for leukocytes and 2.5-3 x 105
per mm3 for platelets.
The size of the different blood cells under normal, non-deformed conditions is
approximately the following (Strackee and Westerhof 1993):
146
Erythrocytes are about 8μm length and 2μm and 1μm thickness at the thickest
and thinnest points respectively; leukocytes are about 6-7.5μm in diameter
(when flattened, the diameter increases to about 15μm); and platelets are
approximately 3μm long and 0.6-1μm thick.
The model developed in this thesis is applicable to different cell volume and can
be adjusted to both Newtonian and non-Newtonian flow.
The specific gravity of blood plasma is 1.03 (Levick 1995; Miller and Leavell
1972), compared to 1.10 for erythrocytes. The specific gravity of whole blood is
between 1.041 and 1.067 and is taken as an average of 1.058 (Miller and
Leavell 1972).
The relative viscosity of blood is dependent on the hematocrit and for human
blood would be approximately 4 for hematocrit of 47% (Levick 1995).
3.4.1.1. Properties of blood in patients with chronic liver disease
In patients with chronic liver diseases alterations in the blood level of proteins,
lipids and fibrinogen, and changes in blood viscosity lead to structural and
metabolic abnormalities in the membrane of the erythrocytes (Sule et al. 2002;
Takashinizu et al. 2000). Sule Tamer et al. (2002) have shown a decrease in
plasma and blood viscosity, and a significant reduction in hematocrit in patients
with chronic liver disease (29.83±5.9%) compared to healthy subjects
(45.61±2.15%). Suggestions that erythrocytes become more rigid and have
decreased deformation ability in non-alcoholic liver disease show gaps in the
current knowledge of blood properties. Advances in this area can benefit the
model proposed in this thesis by altering the fluid parameters to represent blood
properties in diseased patients. The hematocrit of the blood, which determines
the blood viscosity, varies with changes of temperature and with the stage of the
disease.
3.4.1.2. Non-Newtonian properties of blood
It is well known from viscometry that blood plasma behaves like a Newtonian
fluid. The composition of the blood plasma is approximately 90% water and
10% of a combination of proteins (predominantly), organic and non-organic
147
substances. This is not the case with whole blood, revealing the role blood
elements play in the properties of the flow.
Another very important factor in describing the behaviour of whole blood is the
shear rate. Data suggest that for low shear rates of less than 10s-1 and hematocrit
less than 40%, the Casson’s equation can be used (Fung 1993; Strackee and
Westerhof 1993):
γηττ &+= y (3.4.1.2.1.)
where τ is shear stress, γ& is the shear strain rate, η is a constant, and τy is yield
stress in shear constant. At high shear rate the viscosity μ can be considered constant and the behaviour
of blood Newtonian:
γμτ &= (3.4.1.2.2.)
Investigations and analysis of flow through stenosis or at the point of branching
of an artery or vein (point of bifurcation) cannot usually be carried via the
previous equations as the stress and strain rate distributions are unknown (Fung
1993)..
Blood flow through the portal vein occurs at low Reynolds numbers, and the
Navier-Stokes equations can be used for its mathematical modelling (Yousef et
al. 2001).
Reynolds numbers (Re) are calculated using the following formula:
μρLU
=Re (3.4.1.2.3.),
where U is the mean velocity at the inlet, L is the diameter, μ is the viscosity
and ρ is the density (Strackee and Westerhof 1993).
Reynolds number indicates whether a flow is laminar or turbulent (Re above
2500 usually indicates turbulent flow in a long, smooth tube) (Jensen 1996).
Reynolds numbers are equal to the inertial forces divided by the viscous forces
(Jalan et al. 2004).
3.4.1.3. Blood viscosity
As mentioned earlier blood viscosity in most cases is not a constant. Here, some
of the possible reasons for this, in addition to the ones featured above, are
discussed.
148
Red blood cells in humans have the ability to form aggregates, called rouleaux
(Jalan et al. 2004). These are chain-like structures, which can branch and take
any form. The individual cells are attached to each other via their larger faces.
The smaller the shear rate, the greater or larger aggregates that are formed
(Fung 1993; Levick 1995; Petrov 1994; Strackee and Westerhof 1993). Large
rouleaux cause a greater disturbance in flow than that which occurs from the
individual cells. The ability of erythrocytes to deform and elongate also plays a
role in determining viscosity – increasing the shear rate leads to reduction in
viscosity. Although shear rate plays an important role in the deformation of the
red blood cells it is not the only factor for this phenomenon.
Tubular flow of blood shows a zone “free” of cells next to the wall. The width
of this wall “plasma-only zone” increases with the increase in the shear rate
(Fung 1993; Strackee and Westerhof 1993). Platelets, on the contrary, tend to
move towards the wall of the vessel. When measuring blood viscosity it is
important to recognise the existence of this zone and to take it into account
when representing the data. Thus, the term apparent viscosity is usually used to
describe the blood viscosity. The apparent viscosity and the hematocrit are
lower in smaller vessels and where branching from larger vessels into smaller
vessels occurs (Fung 1993).
Viscosity can be defined as the ratio of shear stress to shear rate. Shear stress
(N/m2) is the shearing between the layers of laminal flow and depends on the
axial pressure gradient. Shear rate [(m/s)/m or s-1] is the change in velocity per
unit radial distance (Levick 1995, Chapter 8).
3.4.1.4. Blood cell behaviour as suspended particles in the blood flow
Blood cells, flowing in an axial direction, would tumble, thus disturbing the
flow pattern. Rouleaux of cells would have a greater impact on the flow
behaviour than the disturbance of a single cell (Fung 1993). This also means
that at any given time at any point in the cross sectional area of the blood vessel
there will be different flow disturbances based on the path of the blood cells and
rouleaux. If the rouleaux are broken into individual cells, due to increase in the
shear rate, the viscosity of the fluid will decrease (Fung 1993). This is one of
149
the explanations of why the profile of whole blood flow is not parabolic as in
Poiseuille flow. Researchers (Gauthier et al. 1972; Goldsmith 1971) have
described the erratic sidewise movement of the blood cells due to random and
frequent encounters with other cells (Fung 1993, Chapter 3).
3.4.2. Newtonian flow
The Navier-Stokes equation (Smart Measurement 2004) can be described as
follows:
where p is pressure, v is velocity, μ is viscosity, ρ is density and b is body force
(Smart Measurement 2004).
An incompressible fluid is one whose density is constant throughout.
The Navier-Stokes equation, in combination with the equation of continuity,
describes completely incompressible Newtonian flow with constant viscosity
(Dinnar 1981; Jensen 1996). Whole blood behaves as a Newtonian fluid at high
shear rate (Fung 1993; Jensen 1996). In the portal vein, shear rates are lower,
and the flow has to be seen as non-Newtonian.
By definition, the Newtonian flow can be described as
drduμτ −= (3.4.2.),
Where τ is the shear stress, μ is coefficient of viscosity, and drdu is the velocity
gradient perpendicular to the direction of shear [s−1].
3.4.3. Factors governing portal vein hemodynamics
Patients with advanced chronic liver diseases often have abnormal
hemodynamic parameters showing hyperdynamic circulation in the splanchnic
and systemic vessels (Koda et al. 2000). There are many factors determining the
local hemodynamics of the blood flow through the liver. These include the size
of the shunt, its diameter and position, the type of bifurcation and the curvature
150
of the vein and shunt, and the size and position of blockages within the portal
vein or liver.
In portal hypertension, the portal vein diameter enlarges (>13mm) and the
platelet count is low (<140,000 per mm3) (Grace et al. 1996).
3.4.4. Specific factors impacting on branched vessels
Branching of blood vessels is frequent in the human circulatory system. These
changes in the vessel diameter, and consequently the velocity of the flow are
difficult to model and predict. Commonly used assumptions (as outlined further
in this chapter) are that the walls are rigid, the branches have equal dimensions
and the flow is steady and incompressible.
Figure 3.4.4. Rigid, simple model of equal dimension branching of a vessel
(Jensen 1996)
When the total cross sectional area of all vessels (post-branching) becomes
larger, the velocity decreases and the velocity profile changes (Jensen 1996).
The pressure difference needs to overcome the viscous resistance. If all
branches have the same dimensions, the radius of each branch Rn must be equal
to the initial vessel radius R0 divided by 4 n , where n is the number of branches
(Jensen 1996) if resistance is to remain constant. As such, it would be very
difficult to estimate the radius if the branches have different dimensions. Other
factors, such as the maintenance of constant shear stress may play part in
keeping the resistance constant.
151
Experimental studies have shown that continuous loads of hemodynamic
stresses influence the tissue architecture of a branched vessel (Kobashi and
Takehisa 2000). Under in vitro continuous flow conditions, cultured arterial
tissue from bovine smooth muscle cells, for example showed endothelial cell
alignment in the direction of the flow in the branched region after 24 hours,
with the exception of the region of predicted flow separation where they
retained polygonal form (Kobashi and Takehisa 2000).
Flow disturbance is usually seen at bifurcations, branches or stenotic areas in
blood vessels, which are also the regions responsible for damage to blood cells
and vessel wall and where platelet thrombi are usually observed (Mori 1989,
Chapter 19).
It was mentioned above that erythrocytes tend to concentrate along the axis of
the blood flow while platelets tend to move closer to the wall. This might be an
explanation for the frequent findings of platelet thrombi downstream of
bifurcations and stenosis. Another explanation is given by the convection-
diffusion theory, which looks at thrombus formation as a result of endothelial
layer damage, and the adhesion of platelets to the subendothelium (Strackee &
Westerhof 1993, Chapter 16).
Most studies involving branched vessels are related to the arterial tree, but
might be used in studies on the venous system. In studying the uniform shear
stress hypothesis for a mean value of the branching exponent = 3 (where the
parent diameter equals the sum of the two daughter diameters), Karau et al
(Karau et al. 2001) demonstrated that there was little correlation between the
above, for a heterogeneous distribution of the exponent values in the vascular
tree.
At the point of branching the flow profile will not be parabolic, and can be
assumed to be blunt in the beginning, and gradually reaching a parabolic profile
with increasing distance travelled (Jensen 1996).
The simplified portal vein models used in the experimental and computer
simulations in this thesis have two main branches, which each branch into two
sub-branches, as described above in section 3.4 of this chapter.
152
3.4.5. Hemodynamics of vascular grafts
It is necessary to investigate the hemodynamics that affects the performance,
durability and effectiveness of vascular grafts. Different graft models need to be
compared to find ways to improve them so they can be as close to their natural
values as possible. Wall shear stress, for example is deemed to be responsible
for development of intimal hyperplasia, which ultimately leads to stenosis and
thrombosis (Bonert et al. 2003). When a graft is implanted in the human body,
the native vessels which are affected experience change in their hemodynamics
due to, for example, increased blood flow and perfusion pressure. The angle at
which the graft joins the vessel is also very important, as the wall shear stress
changes in the native vessel in the region of impact with the graft flow. The
vessel wall can relax and thus increase the vessel diameter. Some studies have
even shown that such changes can lead to structural changes in the endothelial
cells (Levesque et al. 1986; Reidy and Schwartz 1981) and their orientation.
Many studies over the past three decades have shown the importance of the
elasticity and compliance of vascular grafts being as close to those of the native
vessel as possible. The compliance hypothesis is described as “the patency of a
vascular prosthesis will be optimal if its mechanical properties match those of
the anastomosed natural vessel” (Mori 1989, Chapter 21). Thus, mimicking the
properties of the vessel where the graft will be implanted is a pre-requisite for
the success of the grafting. But even if this was so at the time of first
implantation, changes to hemodynamics over time may affect the structure and
behaviour of the graft.
3.4.6. Impact of portal hypertension on vascular hemodynamics
Shunting after portal hypertension has an effect on the whole blood circulation.
For example, after portocaval end-to-side shunt, an increase of 27±19% in the
cardiac output and a decrease in the peripheral vascular resistance of -23±18%
can be observed, while those parameters are 18±28% and -11±27%,
respectively, after distal splenorenal shunts (Luca et al. 1999). The fact that
peripheral vasodilatation deteriorates after shunt procedures suggest that portal
systemic shunting is more important than increased portal pressure in
153
determining peripheral and pulmonary vasodilatation in patients with cirrhosis
(Luca et al. 1999). Both shunts also decrease hepatic blood flow.
In a study to evaluate portal hypertension and cirrhosis using hepatic venous
pressure gradients measured by duplex ultrasound method in 72 patients (Tasu
et al. 2002), the following comparison between the patients and a control group
was made in terms of portal vein parameters:
Group / parameter Control Child-Pugh classified (A, B, C)
Portal vein velocity
(cm/s)
20.5 ± 4.71 Between 9.02 and 15.42
Portal vein diameter
(cm)
1.1 ± 0.26 1.22±0.31 and 1.35±0.35
Portal flow rate (cm/s) 21.68 ± 7.68 Between 14.6 and 21.38 max
variation ± 16.6
Table 3.4.6.1. Comparison between control and Child-Pugh classified
patients (Tasu et al. 2002)
Child-Pugh classification is used to determine the severity of the liver disease
based on scores given according to the degree of ascites, the plasma
concentrations of bilirubin and albumin, the prothrombin time, and the degree
of encephalopathy, and are grouped in three classes:
Points Class Life expectancy Perioperative mortality
5-6 A 15-20 10%
7-9 B Candidate for transplant 30%
10-15 C 1-3 months 82%
The measured values of vessel diameter and velocity are also dependent on
administration of medications by the studied subjects (Gibson et al. 1996), and
Table 3.4.6.2 below represents the baseline values only (before administrating
ketanserin) which differ from the post administrative ones. This table shows the
usefulness of Doppler flowmetry and the type of data obtainable with this
154
method. These values can be used in the creating of the computer model
described in this thesis.
Vessel / Parameter Vessel diameter (cm) Peak velocity (cm/s)
Hepatic artery 0.43 43.2
Main portal vein 1.33 19.6
Right portal vein 0.95 14.7
Para-umbilical vein 0.51 18.9
Table 3.4.6.2. Duplex Doppler Ultrasound measurements of vessel diameter
and average velocity in 14 patients with alcoholic cirrhosis and portal
hypertension (Gibson et al. 1996)
To illustrate the diversity of data obtained from different research groups, the
following table shows the same parameters measured by Duplex
Ultrasonography in control and portal hypertensive patients (Ozaki et al. 1988).
In this study there are no differences in the vessel diameter in the two groups,
yet the average velocity is nearly three times, and the blood volume more than
60% lower, in portal hypertensive patients.
Group / parameter Control Portal hypertension
Portal vein velocity (cm/s) 18.99 ± 0.86 6.21 ± 1.6
Portal vein diameter (cm) 0.95 ± 0.033 1.008 ± 0.049
Portal flow volume
(ml/min)
874 ± 44 450 ± 86
Table 3.4.6.3. Duplex Ultrasonographic measurements in 22 control and 29
PH patients (Ozaki et al. 1988)
Maximal blood flow through the portal vein in cirrhotic patients did not exceed
15.5cm/sec, whereas the minimal flow velocity in the control group was
15cm/sec. The congestion index in cirrhosis is shown to be greater than 18%
(Maisaia et al. 2001).
155
Group/Parameter Control Cirrhotic portal hypertension
Portal vein diameter Average 0.95cm Average 1.278cm
Maximal flow velocity 13.87cm/sec 21.79cm/sec
Mean velocity 11.13cm/sec 17.96cm/sec
Minimal flow velocity 9.53cm/sec 15.01cm/sec
Congestion index 6.547% 47.616%
Table 3.4.6.4. Comparison between control and cirrhotic patients (Maisaia
et al. 2001) using Doppler Ultrasonography data
Comparison between the vessel diameter and peak velocity (Table 3.4.6.2) in
hepatic blood vessels shows the usefulness of measuring more than one vessel
to determine and grade portal hypertension (Gibson et al. 1996).
Contrary to the above, Duplex sonography is shown to diagnose portal
hypertension but is unable to assist in its grading. Haag et al. (1999) made the
following observation in 375 patients with portal hypertension and a control
group of 100 patients, where either velocity under 21cm/sec or a portal vein
diameter larger than 1.25cm was indicated the presence of portal hypertension.
Portal vein diameter Portal vein flow velocity
Portal hypertension
compared to healthy
individuals
+ 30% - 44%
Table 3.4.6.5. Difference between healthy individuals and patients with
portal hypertension (Haag et al. 1999)
Although scarce in the literature, follow-up studies of patients with cirrhosis can
be very useful to determine the outcome of treatment and to investigate possible
alternative methods to improve hepatic hemodynamics. In 2002 Bolognesi et al.
(2002) reported on 4 years of follow-up in 41 patients with cirrhosis and 35
controls. In cirrhotic patients, portal blood velocity increased post-
transplantation from 9.1±3.7cm/sec to 38.3±14.6cm/sec, and the blood flow
volumes increased from 808±479 mL/min to 2817±1153mL/min (Bolognesi et
156
al. 2002). Pulsatility index increase corresponded with the increase of portal
blood flow and velocity from 1.36±0.32 to 2.34±1.29.
3.5. Theoretical reasoning for the proposed model During the past 80 years, modelling of different parts of the cardiovascular
system has played a significant role in our understanding of its function in
health and disease. Most models have been developed to deal with the workings
of the heart or parts of the heart (valves, ventricles and carotic artery), or to
describe the aorta. Some studies have investigated and modelled the whole
cardiovascular system (Žáček and Krause 1996) using a system of pipes and
reservoirs with good agreement to experimental measurements. All of those
models have some degree of simplification, and the one in this thesis is no
different.
In many of the models, blood is assumed to be Newtonian in large vessels and
sometimes considered non-Newtonian only in the capillary system (Žáček and
Krause 1996), but still making provisions in the mathematics to account for the
non-Newtonian flow in small vessels (introducing drag coefficient for each
component of the system, for example).
The model of the shunt and the portal vein presented here uses both Newtonian
and non-Newtonian behaviour for comparison and to enable the use of either of
them in future modelling of components of the vascular system.
3.5.1. Laminar flow of blood
Even when the flow is laminar, the profile of blood flow in a cylindrical tube is
more blunted than parabolic (Levick 1995). At high shear rates red blood cells
tend to orient parallel to the direction of the flow due to the shearing of lamina
against lamina. As the cells are displaced closer to the centre of the tube, a small
layer next to the wall becomes nearly “free” of cells, thus playing role in the
blood viscosity measurements (mentioned in the previous chapter). Shear rate
could be defined as the change in blood velocity per unit distance across the
tube. In this thesis turbulent flow has not been modelled as it occurs when the
Reynolds number is very high (somewhere around 2000), as this is not the case
157
in the portal vein. The Reynolds numbers in the model have not exceeded 600
in any part of the vessel.
If the flow is laminar, the resistance is due to internal friction of the lamina
layers and not friction between the blood and vessel wall (based on the zero-slip
condition). Shear rates depend on the diameter of the vessel, and increase with
the decrease of diameter provided the flow rate is the same.
In case of laminar Newtonian flow through a tube, Poiseuille’s law can be
applied:
lQ
ηπα8
4
ΔΡ=& (3.5.1.)
Where Q is the total flow per unit time; & ΔΡ is the pressure difference (pressure
drop) between the two ends of the tube; η is the viscosity coefficient, α and l
are radius and length.
With increase in viscosity there is an increase in the pressure gradient to
maintain the given flow rate (Jalan 2004). To maintain laminar flow after
shunting, the ratio between the native vessel and the graft also plays an
important role. Szilagy et al. (1960) and Kinley et al. (1974) have modelled
empirically and mathematically, respectively, the optimal graft to native vessel
diameter ratio for end-to-end anastomoses to be 1.5.
Sometimes there are disturbances in the flow due to wall-fluid interaction,
branching or stenosis, but this does not mean there is turbulence. In most cases
it is a temporary condition and depends on the oscillatory cycle (Dinnar and
Raton 1981).
3.5.2. Fluid mechanics definitions
A number of terms relevant to subsequent sections are defined below:
Viscosity is the resistance to flow due to friction of molecules in a moving
liquid, and is dependent on the concentration of suspended cells, the radius of
the vessel and velocity of the flow (Smith, J J & Kampine, J P 1984). It is
constant in homogeneous fluids, but changes for non-homogeneous liquids like
blood.
Stress can be defined as force per unit of perpendicular area, thus being
associated with a direction and a plane, hence, a tensor (Dinnar and Raton1981).
158
Strain is a non-dimensional ratio between deformation of a parameter and its
pre-load value.
The relationship between stress and strain will be linear for perfectly elastic
material under low acting forces (Hooke’s law). Hooke’s law for a linear elastic
material means that the material would deform proportionally to the force
applied and with the cessation of its application will return to its original
dimensions (Strackee and Westerhof 1993, Chapter 13). In a fluid such as
blood, on the other hand, the magnitude of the applied force will determine the
movement of the particles, hence viscosity. Blood vessels have some of both
relationships and are considered viscoelastic (Dinnar 1981; Strackee and
Westerhof 1993). Blood vessels, such as veins, which deform axisymmetrically
because they have thin walls, do not obey Hooke’s law (Dinnar and Raton 1981,
Chapter 8).
When tensile stress is applied to a material, the resulting strain is determined by
Young's modulus constant (force/length2) defined as the ratio of the stress to the
corresponding strain. It is a way to measure the stiffness of a material, and may
be used as a measure of the stiffness of the vessel wall material (Greenwald
2002).
When the two tensors, stress and strain, act in the same direction (the ratio
between them is in that direction) the material is called isotropic (Dinnar and
Raton1981). The material has equal properties in all directions.
A homogenous material is one which has the same structure and properties at
all points, and is also isotropic.
Blood vessels are neither homogeneous, nor isotropic (Jalan et al. 2004;
Strackee and Westerhof 1993, Chapter 13).
Shear rate at any point of the vessel wall is the normal derivative of the fluid
velocity of the blood. Shear stress is the product of shear rate and kinematic
viscosity of blood at that shear rate (Mori 1989). Both parameters vary in the
vascular system with time.
Streamline is an empirically drawn imaginary curve in space at a given time
which is parallel to the direction of the motion of the fluid at that point, and
does not change with time in steady flows (Jalan et al. 2004).
159
The flow remote from the entrance of the vessel is called Fully Developed
(assuming the entrance is far away and does not have an impact on the flow, i.e.
the flow has reached its profile and will not change further if the vessel
diameter is constant) (Moore et al. 1998), and is characterised by a linear fall in
pressure with distance and a linear increase in pressure gradient with flow-rate
(Jalan et al. 2004).
3.5.3. Basic laws governing the cardiovascular system
If we simplify the human circulation system, from Newton’s law for the
conservation of mass, the relationship between the changes in venous versus the
change in arterial blood volume is adverse. If that is so, any decrease in the
venous volume will lead to a decrease in venous pressure determined by the
magnitude of venous compliance (Strackee and Westerhof 1993). As veins
carry around 75% of the blood in the body their high sensitivity to mechanical
stimuli (due to the smooth muscle cells in their walls) and high compliance (low
elastic modulus), as well as their lower Young’s modulus compared to arteries
has to be taken into account when modelling blood flow (Fung 1993). Young’s
modulus (modulus of elasticity) is the ratio of stress and strain, and is a straight
line only for purely elastic materials (Strackee and Westerhof 1993). Venous
pressure is determined as the pressure which blood exerts within the vein and is
usually between 60 and 120mmH2O in a lying position (Miller and Leavell
1972). The pressure in the portal vein model in this thesis is 40mmH2O, which
is the average in portal hypertension.
3.5.4. Commonly used assumptions
Homogeneity. The assumption used for simplifying the mathematical model of
blood flow is that blood is homogenous. This however cannot be a valid
assumption when considering blood cells and particle deposition on the blood
vessel wall (Dinnar and Raton 1981; Finol and Amon 2002; Fung 1993; Song et
al. 2000; Strackee and Westerhof 1993).
160
Axisymmetric flow. Most models of blood assume the existence of
axisymmetric flow. For example, such flow was used to model stenotic arterial
flow in a vascular tube (Gurlek et al. 2002; Marques et al. 2003) or to model
pulsatile flow (Marques et al. 2003).
Example:
Let us assume that the vessel is perfectly elastic and the flow is incompressible
(Poisson’s ration =0.5, as for rubber). εσ E=
Poisson’s ratio is the ratio of transverse contraction strain to longitudinal
extension strain in the direction of stretching force, and is positive for all
common materials as they become narrower in cross section when they are
stretched.
If we want to solve the elasticity equation for a tube wall motion under varying
internal pressure (i.e. pulse), and the wall radius expands from a to a + η(x,t),
then the strain is: aa
aa ηπ
πηπε =−+
=2
2)(2
The forces relationship tension = pressure + inertia must be observed.
The circumferential tension T divided by the tube thickness φ can give the
value of stress per unit length: ( )a
txET ,ηφ=
Endothelialization after fibrin deposition on the vessel wall can be viewed as
incorporation of the deposition into the vessel wall. Thus, when modelling the
region of stenosis, it can be considered as a region of different wall thickness,
elasticity and radius (Dinnar and Raton 1981).
Steady flow. This is a flow in which the velocity at any one point of the flow
never changes (and the viscosity is neglected) (Jensen 1996): 0=∂∂
tυ
Incompressible flow. This is a flow which neither enters nor exits the vessel
(Song et al. 2000) and is used in CFD modelling of arteries (Moore et al. 1998;
(Morsi et al. 2001). If an incompressible fluid flows through an infinitely long
vessel, the cross sectional area at one point multiplied by the velocity at that
point will equal the cross sectional area at another, different point multiplied by
the velocity at that point.
161
3.5.5. Additional effects impacting the circulation
Shear Stress: The vessel wall shear stress is one of the factors considered
important for vessel growth. In a theoretical study, Hacking et al. (Hacking et
al. 1996) evaluated the effect of shear stress on vascular diameter, and
concluded that there must be other factors leading to steady network structures,
as shear stress alone does not cause such stability. The calculations carried out
showed a development of an optimal constant diameter when the vessel or
vessel tree was perfused with constant flow or a constant pressure source with
internal resistance. This study also showed a regression diameter (increasing to
infinity) of the vessel when constant pressure perfusion was applied. The non-
homogeneous distribution of stress is responsible for non-homogeneous
remodelling of tissue due to stress changes in the organ or vessel (Fung 1993).
There are numerous theoretical and experimental studies of the effects of shear
and circumferential stress on vessel growth ranging in subjects from rats to
human, and investigating Newtonian and non-Newtonian flow conditions.
Shunting and portal vein ligation: Cirrhotic portal hypertension may alter the
relationship between portal pressure and capacity to develop shunts in rats
(Geraghty et al. 1989).
Partial ligation of the portal vein in rats leads to increased portal venous inflow,
which helped in maintaining portal hypertension (Sikuler et al. 1985). That
study, on 45 portal hypertensive rats and 29 control rats, showed increase in
portal pressure and resistance (around 80%) and decrease in portal venous
inflow (around 80%).
Shunting and fibrosis: Comparing dogs with artificially induced portal fibrosis
and healthy dogs (Sugita et al. 1987), there were no portal systemic shunts,
although the intrahepatic presinusoidal portal hypertension developed by the
treated dogs could be seen as the portal vein pressure increased by 50% in those
animals compared to controls. Increase in both the portal venous flow (around
30%) and intrahepatic portal vascular resistance (over 50%) has also been
recorded.
Clamping of the vessel: A problem when using blood flow models is the
mismatch impedance caused by the clamps on the vessel segment studied
(Charara et al. 1999).
162
Reynolds number (Re): Reynolds numbers in the human circulation are
usually below 2000 (or even assumed to be 125 in arteries (Moore et al. 1998)),
but stenosis or clots on the vessel wall can increase velocity and lead to a
disturbed or even turbulent flow close to the stenosis (Jensen 1996) with
Reynolds numbers above 2500.
3.5.6. Wave propagation in the cardiovascular system
Waves generated from the heart travel in the direction of the blood flow, and
due to their reflection at sites of discontinuities, waves in the opposite direction
are observed. Bifurcations are a good example of regions responsible for wave
reflection.
In non-viscous models the governing equations for wave propagation are linear
(Fig 3.5.6).
Figure 3.5.6. Relationship between pressure, area of vessels and the speed
blood moves with through them in the circulation (Miller and Leavell 1972)
From Figure 3.5.6, it is evident that pressure is lower in the veins compared to
other parts of the circulation. In the portal vein the blood velocity is increasing,
as shown above for veins in general, with the high increase of flow and smaller
increase of vessel area. The blood would be expected to have higher velocity in
163
the portal vein after it is joined by the splenic vein (which itself usually already
incorporates the flow from the inferior mesenteric vein), as the area will
increase less than the flow. In portal hypertension, this relationship is still valid,
but the values are reversed. The pressure is higher and the velocity lower, with
the area being larger or smaller depending on vessel patency.
The pressure distribution without the area and velocity are shown in the
following figure 3.5.6.1.1.
3.5.6.1. Pulsatile flow
The pumping of the heart and the action of the aortic valve generate highly
pulsatile flow. This has an impact on the analytical modelling of the
cardiovascular system, and the flow can be considered stable only in part of the
microcirculation. Laminar flow can be assumed in most parts of the circulatory
system when turbulence is not observed. The flow in the microcirculation is
closest to steady laminar flow than any other region in the human body.
Peripheral veins have a similar flow pattern, but the closer the vein is to the
heart’s right atrium, the greater the suction effect creating a pulsatile pressure
gradient. The valves of the small veins contribute to the pulsatility of the venous
flow by preventing back flow, thus creating a positive pressure gradient (Dinnar
and Raton 1981, Chapter 6).
Figure 3.5.6.1.1. Pulsation of pressure in the circulation (Dinnar and Raton
1981)
164
The vessel impedance and the pressure gradient determine flow. Because the
pressure gradient depends on the initial and reflected wave and the phase
between them, it is difficult to measure or estimate its value. The wave
propagation depends on the branching and thickening of the wall of the blood
vessel. The last two factors, together with the cross-sectional area of the vessel,
have a linear relationship with the average velocity of the blood. The small
blood cell “free” zone near the vessel wall is responsible not only for the non-
parabolic profile of the flow, but also for the initiation of reversal flow (the
reversal spreads from that zone to the axial layer due to viscosity). If this is the
case, both viscous and inertial terms must be considered (Dinnar and Raton
1981).
Figure 3.5.6.1.2. Mean velocity and velocity fluctuations in the cardiovascular system (Dinnar and Raton 1981) As can be seen from figure 3.5.6.1.2. velocity fluctuations can be expected in
veins as large as the portal vein.
Blood vessel walls are commonly modelled as rigid (Bonert et al. 2003; Gurlek
et al. 2002; Marques et al. 2003; Petkova et al. 2003; Siro et al. 2002;
Starmans-Kool et al 2002; Xu et al. 1999), but in nature their elastic properties
introduce a propagation speed for the blood pressure wave (Jensen 1996).
Researchers model compliant vessels (Hiromichi et al. 2001), but for simplicity
165
in manufacturing vessels for laser Doppler anemometry (LDA) rigid walls have
been assumed in this thesis. Incompressible flow through a rigid vessel would
enable immediate propagation of a pressure wave through the tube with any
pressure change. Human blood vessels accommodate steady, gradual wave
propagation (Jensen 1996) due to their mechanical properties.
3.5.6.2. Importance of hemodynamics on modelling of blood flow
Graft patency and vascular pathophysiology are impacted by mechanical forces,
generated by the blood flow (Charara et al. 1999; Nerem 1991; Stanley et al.
1982) and the endothelial layer is most affected by those forces. A common
place for the appearance of atherosclerosis is in areas of bifurcation in arteries,
which is included in the model used in this thesis. As mentioned previously,
thrombosis and intimal hyperplasia are other complications in areas of
bifurcation (Charara et al. 1999). Flow separation is also possible if an
incompliant graft is connected. Flow separation is identical to incidence and
occurrence of negative axial velocity (Strackee and Westerhof 1993).
Limitations to organ culture models designed to examine the vascular
remodelling are the absence of the effect of blood forming elements and
pulsatile flow on the hemodynamics (Del Rizzo et al. 2001), even though such
models are useful for analysis of early vascular occlusive disease.
3.5.7. Tissue culturing studies
Tissue culturing experiments were conducted in order to establish what
improvements are required to the device in which scaffolds would be seeded to
produce novel vessels. These experiments aided in understanding of the process
of growing in vitro blood vessels on biodegradable polymers and were
undertaken simultaneously with the computer modelling of the portal vein flow.
3.5.7.1. Tissue culture methods
Tissue culture consumables (Life Technologies Pty Limited, Mulgrave,
Victoria, Australia) used in these experiments included 5mL, 10mL and 25mL
pipettes, 50mL centrifuge tubes, 100mm petri dishes, 25cm2 and 75cm2 vented
flasks, DMEM (Dulbecco’s Modified Eagle’s Medium) high glucose, Foetal
166
Bovine Serum, Trypsin-EDTA (ethylenediaminetetraacetic acid), and
Penicillin/Streptomycin Solution.
Phosphate-buffered saline (PBS) was prepared as follows: 8g NaCl, 0.2g KCl,
1.44g Na2HPO4 and 0.24g KH2PO4 were dissolved into 800mL distilled water.
The pH of the solution was adjusted to 7.4 with HCl, following which water
was added to 1L. The PBS solution was sterilized by autoclaving at 121°C and
then stored at room temperature.
For long-term storage, cells were frozen using the following methodology: cell
monolayers in flasks were washed with PBS, trypsin-EDTA was added and the
flask incubating for approximately 5-10 minutes at 37°C. Growth medium was
added and the cells collected by centrifugation (1500rpm at 4°C for 5 minutes).
The cells were cooled sequentially at -20°C and -80°C, and finally placed in
liquid nitrogen.
As sterility of equipment and solutions was required at all times, the novel
design of the Bioreactor necessitated its manufacture solely from materials
suitable for autoclaving (glass and medical grade stainless steel) which could be
easily and completely disassembled.
3.5.7.2. Preparation and test studies
Dulbecco’s Growth Media (Dulbecco’s phosphate-buffered saline (PBS)
solution) was supplemented with 1% (v/v) Penicillin/Streptomycin Solution and
10% (v/v) Foetal Bovine Serum (Life Technologies Pty Limited, Mulgrave,
Victoria, Australia). As a sterility control, 20 mL of the media was placed into
75mL flasks and incubated for 48 hours at 37°C in a humidified atmosphere
with 5% CO2. After exchanging the media and a further incubation for 72 hours,
samples of the media and placed onto nutrient agar plates and examined after 24
hour for growth of bacteria and fungi.
The next step involved the addition of polymer squares (1x1 cm) to the flasks
with fresh media. If no contamination was observed after 72 hours of
incubation, the above procedures were repeated using 3T3 rat fibroblast cells.
This cell line was chosen as it was easy to maintain and to passage, and allowed
167
for optimisation of the protocols and methods. The cells were grown until they
reached confluence (usually after 3-4 days) and passaged.
Experimental work was carried out with 3T3 cells and primary endothelial cells
on Vicryl and several polymers (Polycaprolactone 3D scaffolds with different
porosity, Acrylonitrile-Butadiene-Styrene).
Parallel experiments were carried out with and without polymers added to the
flasks, using cells from the same passage incubated for the same length of time
and under the same conditions.
From the experimental work, it was observed that when growing cells on
polymers (in this case, Acrylonitrile-Butadiene-Styrene at first), there is not
only a risk of stripping the tissue from the scaffold, but after passaging, the cells
which grew on the polymer showed different behaviour. Some of the cells
started growing in strips and did not spread as much as the ones which did not
have polymer in the flask. In experiments with polycarbonate, sandblasted
blocks of approximately 10mm2 were placed in the flasks, and cells attached to
the polymer within 2 minutes of incubation. Some experiments were carried out
with aortic leaflet, aortic and coronary artery, thoracic and radial artery cultures,
either via extracting the cells or culturing off-cuts of those tissues. Aorta cells
were also cultured with Vicryl coated braided (polyglactin 910) suture
(undyed), where the coating was stripped with acetone, and also some sutures
into which knots were introduced. Tissue culturing took place over 4 months,
with confluency taking longer after the third passage. Cell proliferation was not
complete, and did not have a clear pattern or preference in some areas on the
polymers Experiments were also carried out using Polycaprolactone three-
dimensional tubing with a range of porosities for improving cell-adhesion. More
studies on the relationship between scaffold wall porosity and cell adhesion are
needed, with advances in this area occurring rapidly. As developing a
biodegradable scaffold with seeded cells was not the aim of the research, but
rather the understanding of the requirements from a bioreactor for such tissue
culturing studies, the experiments were not continued. However, preliminary
observations showed that the type of material used as scaffold and the cell
adhesion method used are important factors.
168
3.6. LDA experimental set-up The LDA experimental set-up is described below. The CFD simulations and the
LDA visualizations are presented in the Results chapter. Details on the design
of the bioreactor are presented in the Bioreactor chapter of this thesis.
3.6.1. Bioreactor and LDA
The current design of the bioreactor is described in detail in the next chapter.
Here, information on the experimental set-up is presented.
3
Fig 3.6.1. Schematic representation of the experimental set-up
Photographic representations of the experimental set-up are presented in the
Results chapter of this thesis.
The bioreactor, made entirely from glass for better transparency, was connected
to the pump via silicone rubber tubes, and was emersed in a fish tank filled with
water to improve the reflection surface of the laser beam. The water in the fish
1. Personal Computer 2. Laser Doppler Anemometer 3. LDA Cooling System 4. Fish Tank with Bioreactor inside 5. Pump 6. Reservoir
4 2
1
5
6
169
tank was changed daily to minimise the noise in the measurements from dust
particles.
3.6.2. Fluid and model vessel
The Bioreactor was filled with a solution of approximately 32.7% Glycerol,
5.2% CaCl2 (by weight) in distilled water to mimic blood viscosity (Mori 1989).
While limitations of this approach are recognised, some suggestions for future
work to overcome these are proposed. In brief, the viscosity of the fluid could
not be kept constant, as the temperature of the fluid dropped (Klabunde et al.
2003; Smith and Kampine 1984) during measurements. Reasonable attempts to
keep the experiment at room temperature were made by adding warm water to
the fish tank in which the bioreactor was submerged. However this did not
include heating the tank, tubes and reservoir of fluid in which the bioreactor
was emerged. Ways for heating need to take into account the sensitivity of the
laser beam to disturbances in the fluid (which can reflect the beam before
reaching the flow inside the scaffold, leading to unrealistic results, or an
inability to measure any flow). To partially overcome those limitations, future
work may involve the incorporation of a hot plate beneath the tank, as well as
hot plates on the sides which do not obstruct the laser beam. Heating the fluid
inside of the bioreactor could be achieved with a separate heated reservoir
through which all of the suspension has to pass before entering the bioreactor.
This heated vessel should be attached to the reservoir (grey filled cylinder in
Figure 3.6.1, above) and should take the solution from and return it to the
reservoir.
Meta DC Coated Particles (Model 10037) were suspended in the Glycerol
solution to enable refraction of the laser beam and allow flow measurements.
The vessel model, which is a glass representation on the desired scaffold, was
used to allow for LDA measurements. Because the scaffold surface is round,
which causes reflection of the beam and makes measurements very difficult,
and the bioreactor has a rounding wall surface, the experiment was submerged
in the water-filled tank with straight angular (90o) walls.
170
The glass vessels were manufactured by Bartelt Instruments (Heidelberg West,
Melbourne, Australia) based on the two designs (with and without blockages)
supplied by the researcher and used in the computer modelling.
3.6.3. Pump and reservoirs
Different types of pumps have been used in flow measurements, and all of them
have certain advantages and disadvantages. Roller pumps, although useful for
studying the behaviour of seeded grafts, produce uncontrolled sine wave flows
(Charara et al. 1999). High-pressure systems, for example, generate pulse, but
do not represent physiological conditions. On the other hand, computer-
controlled gear pumps generate a more natural pulse, but are difficult to keep
sterile (while the fluid passes through).
We have used a peristaltic pump (Easy-Load™ Master Flex, Millipore), which
operates by squeezing the outside of the silicone tubes, carrying the blood-like
fluid. Thus, there is no contact of the pump to the surfaces which are
biologically sensitive and require sterility. The operational range of the pump is
6-600 rpm, with a frequency between 50 and 60Hz.
3.7. Ideas for future work
3.7.1. Heating of fluid
For the purpose of stably maintaining the viscosity of the fluid in the
bioreactor during LDA flow measurements, heating options for both the
glycerol (or other) solution and the water in the tank are needed to maintain
a constant room temperature (or body temperature for other solutions).
Heating plates around the tank, a separate heating chamber for the blood-
like solution or alternative methods need to be explored to resolve this
problem.
For the purpose of tissue engineering, the growth medium can be heated by
covering the silicone tubes with a warming jacket, or simply placing the
whole system into an incubator (if there is no need of constant monitoring).
The bioreactor developed in this work can be safely placed inside
commercially available incubators.
171
3.7.2. Developing the model vessel from different material
In the present study the model blood vessel was made of glass to allow for a
high degree of transparency, but the limitations of this material are its
stiffness and lack of flexibility, both of which do not correctly represent the
native vessel or the ideal scaffold. Alternative materials, with similar
transparency but also with a reasonable degree of flexibility, need to be
developed for LDA measurements. It is worth remembering that
transparency is not an issue for biodegradable scaffolds used for tissue
engineering, as the cells growing on the material will not permit LDA
measurements.
3.7.3. Blood flow modelling
There are many theories dealing with the changing behaviour of blood. In
this thesis those changes have not been taken into account, but these are to
be considered for future research. It is important to individualise each model
to represent the specific changes occurring in patients’ blood flow so
realistic modelling can be done.
3.8. Conclusions of the Chapter Recent advances in ultrasound have made it possible for non-invasive and
accurate measurements of portal vein blood flow. Computational Fluid
Dynamics (CFD) applied to non-Newtonian blood properties in a branched
vessel by means of specialised software package (FLUENT) assist in
developing a better understanding of hemodynamics in the portal vein. The
results obtained from the computer model can be compared to the flow
measurements in a glass model vessel using Laser Doppler Anemometry
(LDA) The physical experiment is easy to perform for non-experts and has
the advantage of low cost, but some of its limitations include the changing
viscosity of the fluid (due to temperature changes), the stiffness of the
model vessel wall, the errors due to the dependency of the laser beam focus
point on the operator (human error) and the noise due to impurity of the
liquids.
172
As part of this research the glass vessels were made as simplified models of
the portal vein based on the dimensions and shape created for the computer
simulations.
A better understanding of the blood flow pattern using modelling and
visualization may help to minimize the thrombogenesis of artificial blood
vessels and organs (Marques et al. 2003).
Tissue culturing experiments have been carried to gain a better
understanding of the required qualities of the bioreactor which was
developed as part of this research and is presented in the next chapter.
173
CHAPTER 4
Bioreactor
4.1. Introduction The purpose and benefits of developing bioreactor systems was briefly discussed in
previous chapters. In the current chapter an overview of existing designs, ideas and
principles are presented, aiming to cover the designs most widely used, which have
also contributed to the design of the device used in the research carried out in this
thesis. The design of the device is described and its operational performance is
discussed. Advantages and disadvantages are presented and recommendations for
optimisation are made at the end of the chapter.
4.1.1. The use and historical development of Bioreactors Why do we use bioreactors for tissue culturing? As stated by Martin et al. (2004)
‘By enabling reproducible and controlled changes of specific environmental
factors, bioreactor systems provide both the technological means to reveal
fundamental mechanisms of cell function in a 3D environment, and the potential to
improve the quality of engineered tissues’.
One of the limitations of unperfused in vitro tissue engineering systems for growth
of thick, fully grown three-dimensional grafts is that of cell viability (Kofidis et al.
2003; Kofodis et al. 2003). This, in combination with the desire to have stronger
tissues, has lead many researchers to develop different types of pulsatile
bioreactors. All such systems have shown to be beneficial to the cultured tissue, be
it blood vessel or heart valve, in comparison with tissues grown under steady
conditions.
An experimental bioreactor with multiple chambers was used successfully for
myocardial grafts on collagen components under dynamic conditions (Kofidis et al.
2003). Transparent chambers for tissue growth, which can be assembled in parallel
174
so that more than one vessel can be grown under pulsatile conditions
simultaneously and is small enough to fit in an incubator, is presented below
(fig.4.1.1) (figure from original article). The first figure (a) represents the multi
chambers, while (b) shows the experimental set-up as presented by the authors. As
can be seen, such a system is not suitable for growing of 3-D blood vessels, but the
results for smaller grafts are satisfactory.
Figure 4.1.1. Multichamber pulsatile bioreactor (a) and experimental set-up (b) (Kofidis et al. 2003). A bioartificial liver bioreactor, filled with porcine hepatocytes immobilized on
polyester nonwoven fabric (Naruse et al. 2001) operating on a plasma-whole blood
separation principle, is given here as an example of a possible application of these
techniques. Naruse et al. (2001) developed the bioreactor system consisting of two
circuits. One, which was used to separate the plasma and whole blood, and was
thus connected to the model’s circulation, incorporates plasma separation, a plasma
175
reservoir and a roller pump. The other comprises nonwoven fabric bioreactors, an
immunoglobulin adsorbent column, an oxygenator and warmer, a dissolved oxygen
meter, a cell filter and a roller pump. Although the results appeared promising, one
needs to be mindful of the fact that, as with any other blood plasma separation
method, negative impacts could be expected. Another feature of the above-
mentioned experiment was that the xenogenetic perfusion obstacles were bypassed
due to the immunoglobulin adsorbent column.
Some researchers (Shinoka 2002) are of the view that once cells are attached to the
three-dimensional scaffold the construction can be implanted in vivo, and that this
will provide the environment for growth and development of the tissue. Ideally, the
scaffold will degrade totally after implantation, but not too fast to jeopardize the
full growth of the new tissue. As an example, Shinoka (2002) constructed scaffold
that would degrade within 6-8 weeks and had 95% porosity.
4.2. Types of Bioreactors
4.2.1. Tissue culture static bioreactors Static type bioreactors provide good culturing conditions for tissue growth. There
are several designs commonly used.
The two most commonly are those for in vitro growing of tissues and those used as
a host for extracorporeal support. Examples of the second type are presented below
in brief, as the next sections deal with the tissue growth devices only.
From the flow point of view these systems either have a continuum of flow passing
through the scaffold, or there is no flow as such, but rather the graft is submerged
in growth medium in a similar manner to tissues grown in petri dishes. Of those
two options, continuous flow provides better mixing of the medium and flow
conditions, which better resemble those that occur in vivo. Discussions on the
benefits of pulsatile flow were presented in the previous chapter and will not be
outlined here (see Methodology chapter for details).
176
4.2.2. Extracorporeal bioreactor systems – some examples
Some researchers have used a bioartificial liver (BAL) for plasma perfusion
consisting of a column with activated charcoal and porcine hepatocytes (Khalili et
al. 2001). This system is based on the principles used in apparatus for kidney
dialysis.
Modular extracorporeal liver support (MELS) is a large, very complex system,
suitable for hospital use only and requiring the patient to be connected to the
machine for extensive periods of time. This is the first complete system,
incorporating the “CellModule” which harvests human liver cells, the
“DetoxModule” which replaces the bile excretion function of hepatocytes and
removes albumin-bound toxins, and the optional “DialysisModule” for continuous
veno-venous hemofiltration (Sauer and Gerlach 2002).
3-Dimensional pulsatile horizontal bioreactor, for example, was developed by
Bilodeau et al. (2005) providing perfusion to both the inside and outside of the
seeded graft.
4.3. Development of the Bioreactor used in this study 4.3.1. The initial idea and design
Initially, the bioreactor needed to be able to perform a variety of functions,
including providing an appropriate environment for 3-dimensional blood vessel and
heart valve or flat 2-dimensional tissue film development. Although that would be
possible to achieve by exchanging the stainless steel fixating rings, it was not
needed for the purpose of the research conducted for this thesis and did not
progress any further. The diversity of devices used nowadays meant they are highly
specialised to grow one type of tissue only. One of the initial ideas was to allow for
easy adjustment and re-design of the device by non-engineers so the person
responsible for tissue culturing could perform these tasks. The device needs to be
easily sterilised with widely available sterilisation methods, such as autoclaving. In
the proposed design as developed in this thesis, glass, medical grade stainless steel
177
and silicone tubing would be the only materials in contact with the tissue growth
environment, all of which can be autoclaved.
Another requirement was the need for incubation of the whole system to allow
extended periods for tissue growth. The tube, which connects the two inlets and is
used by the pump to provide pulsatility, can be safely driven outside the incubator.
The same applies to the tube providing the outflow from the bioreactor or the
reservoir (i.e. the one which returns the medium to the pump, thus closing the
circuit).
All of those tasks were accomplished and the outcome of the development of such
a bioreactor is presented below in Section 4.3.3. In addition to solving the
requirements outlined above, the design is also inexpensive and can be produced in
large quantities allowing it to be constructed using materials that can be purchased
‘off the shelf’.
4.3.1.1. Parallel plate bioreactor
The flow system developed by Brian Lin (2000), described in this section (Figure
4.3.1.1), consists of an upper reservoir to store the culture medium (Dullbecco’s
Modified Eagle’s medium (DMEM) containing 25 mM HEPES buffer, 10-20%
Foetal bovine serum, and antibiotics). The tubing exiting from the upper reservoir
is wrapped with 3ft. of heating cord using a variable DC power supply to heat the
media to 37oC, which then flows through a flow meter before entering the parallel
plate flow chamber. Such wrapping or alternative heating methods are
recommended as part of future work as a result of the experiments carried out in
this research.
In their design pressure drop due to a height difference between the upper and the
lower reservoirs is responsible for the media flow, which is diffused with a 95% air
with 5% carbon dioxide mixture (in order to control pH levels of the media).
The flow deck is made from cast acrylic, and can be adapted with a variety of
inlet/outlet designs, and silicone rubber with different sizes – for Reynolds numbers
in the order of 100 or less. The top of Lin’s flow deck has three threaded holes: two
are for connectors leading to the inlet and outlet of the flow deck; and the third is
178
for the attachment of the vacuum pump to ensure a tight seal between the flow deck
and the bottom of the 35mm culture dish. Silicone sealant is used to prevent leaks.
Figure 4.3.1.1.Parallel Plate Bioreactor (Lin 2000)
The oxygen in the air allows the cells to live under aerobic conditions. The CO2
regulates the pH by providing a buffering system regulated by the chemical
equation
CO2 (g) + 2H2O ↔ HCO3- + H3O+
The reversibility of the carbonic acid-bicarbonate conversion buffers the media by
releasing or removing H3O+ ions from the solution depending on the pH.
Although different types of endothelial cells (from bovine to human umbilical cord)
are used in flow studies the culturing techniques are essentially the same. Usually
the specific cells are acquired and then grown in tissue culture dishes. They are fed
with minimum essential medium along with antibiotics to prevent contamination.
After 1-5 days, confluent cells are detached by exposure to a solution of trypsin-
EDTA. The detached cells in suspension are then centrifuged and discarded. The
remaining attached cells are then remixed with medium at a 1:3 dilution. Then, the
cells are plated on tissue culture dishes, which are sterilized with UV light.
Experiments are then performed when the cells are nearly confluent but not inert.
179
For the purpose of experiments carried out in this study, and due to the three
dimensional shape of the scaffold, the cells need to be seeded directly on the
polymer and placed in the bioreactor for incubation. If a simple sheet of cells is
required, then the method outlined above can be applied.
4.3.1.2. Pulsatile Bioreactor – where the idea came from
The idea of simulating the physiological pressure and flow of the growth medium
in bioreactors for culturing tissues in order to achieve higher strength of the vessel
or valve is becoming an area of increased scientific interest.
In early 2000 a technical report by Hoerstrup et al. (2000) on the functioning and
development of pulsatile bioreactor for heart valve culturing provided a foundation
idea of what was required in the bioreactor developed in this thesis.
The design was initially reviewed in my research group, resulting in the bioreactor
presented in Figures 4.3.2.1 and 4.3.2.2 of this thesis. The author’s participation in
the re-design included the following areas:
• Removing the angles in the pressure chamber to avoid ‘dead corners’ where
detached cells might adhere,
• Adding another inlet and creating an angle between inlets and chamber to
allow for better mixing of culture medium,
• Increasing the volume in the perfusion chamber and creating two outlets
positioned central to the perfusion chamber,
Later, as part of the research done in this thesis, further revisions were made and
the bioreactor was substantially modified by:
• Eliminating the screws. In the initial design these were present (see below).
However, rust developed after a few weeks, even though they were made of
stainless steel, which is unacceptable in tissue engineering and can cause
contamination of the culture,
• Designing the body of the bioreactor from glass instead of Plexiglass,
• In the final design, the diaphragm was removed, thus removing one of the
areas connecting the chambers where leakage can occur,
180
• The design allowed fixation of the scaffold at both ends (bottom and top),
• Introducing stainless steel as support for the scaffold, thus allowing for non-
stiff, more flexible material to be used for creating the graft.
The bioreactor created by Hoerstrup et al. is presented below, and the picture is
taken from original article (2000).
Figure 4.3.1.2.1 Pulsatile Bioreactor (Hoerstrup et al. 2000)
The body of the bioreactor is made of
Plexiglass; the two chambers are
divided via a 0.5 mm thick silicone
diaphragm. The lower part (the air
chamber) is connected to a respirator
pump. The second chamber
comprises two parts – below and
above where the scaffold is fixed.
There is a provision for changing the
diameter of the scaffold.
1. Air chamber 2a and b are the bottom and top of the perfusion chamber 3. Silicone diaphragm 4. Tube 5. Removable silicone tube 6. Valves inlet 7. Outlet 8. Stainless steel screws
Figure 4.3.1.2.2 Schematic representation of pulsatile Bioreactor (Hoerstrup et al. 2000)
181
This work confirmed the importance of sterile conditions, long-term workability
and reliability as some of the most important aspects of developing a bioreactor for
in vitro growth of tissues.
4.3.1.3. Another example of pulsatile bioreactor – aortic heart valve growth
The bioreactor developed by Dumont et al. (2002) for heart valve tissue culturing is
sufficiently compact to fit in an incubator and consists of a left ventricle where the
valve is placed and an afterload part incorporating compliance (to mimic the elastic
functions of the large arteries) and resistances (representing the arterioles and
capillaries). The left ventricle is made out of silicone rubber and its compression
and decompression are achieved by the movement of a piston powered externally
(and its stroke volume can be adjusted to represent different pulsatility flow
conditions). The idea is to simulate physiological conditions ensuring the tissue has
mechanical and hemodynamic properties similar to those of the natural vessel.
The circuit is presented below (Figure 4.3.1.3), with the figure taken from the
original article.
Abbreviations: LV is the Left Ventricle E is the external circuit V is the tissue-engineered valve C is the compliance P is the pressure transducer R is the resistance O is the reservoir A is the aerator M is the mechanical valve
Figure 4.3.1.3. Pulsatile bioreactor for tissue engineered aortic heart valve
(Dumont et al. 2002)
182
4.3.2. First prototype of the Bioreactor Following from the designs of Lin (2000) and Hoerstrup et al. (2000), optimization
of certain parts of a prototype Bioreactor, and associated working principles, were
carried out in our research group (Damen 2003; Morsi et al. 2001).
This bioreactor design (Figures 4.3.2.1. and 4.3.2.2.) had to fulfil the following
functions – to grow veins, arteries and valves. Thus, the apparatus was
manufactured from stainless steel and transparent plastic, and had the following
parts:
• The air chamber is powered by an air pump and follows the desired
pulsatility depending on the tissue that is grown (frequency of the heart
cycle if heart valves are grown, arterial pulsatility or smaller vein
pulsatility). It is large enough to provide space for the membrane to move
with the pulsing air.
• The different pressure fibrillates the silicone diaphragm membrane at the
desired pulsatility rate. When the membrane is in the neutral position, fluid
will be sucked out of the perfusion chamber. When it is above the neutral
point, a pulse will be introduced through the perfusion chamber. The
membrane is made of 0.8-mm thick silicone rubber, which can be readily
sterilized.
• The pressure-chamber is filled with blood or a substitute and flows to the
tissue culturing chamber due to the fibrillating membrane.
In the final design presented in this thesis (from Figure 4.3.3.1 onwards) and
developed for the purpose of this research, the diaphragm and air chamber were
removed. This achieved several outcomes. Firstly, the device was easier to sterilize.
Secondly, the screws connecting the air and pressure chambers and holding the
diaphragm in-between these were no longer needed. In the prototype bioreactor,
problems were experienced with the screws, as mentioned above. This resulted in a
smaller gap between the chambers, thus relaxing the membrane and creating a
different, non-uniform push of the flow towards the perfusion chamber.
183
The two inlets are made
tangentially to create a
vortex of fluid to ensure
good fluid mixing before
it leaves the chamber
Figure 4.3.2.1. Diaphragm at neutral and above neutral position and the
pressure chamber (Damen 2003; Morsi et al. 2001)
• The tissue is grown inside the perfusion-chamber. This chamber is
customized for the different applications, depending on the intended use of
the bioreactor. The design shown below is for culturing of an artery or vein
without any branches, i.e. a straight tube. A security bridge made out of
three stainless steel bars is used to fix the upper end of the scaffold to the
bioreactor, so they can provide support to the growing tissue during the
cycles of pulsatile flow.
In the final design developed in this thesis, the bridge was removed for simplicity
and a novel way of securing the scaffold was developed. The newly developed
rings make it easier to adjust for a variety of branching options and are much
simpler to assemble and disassemble. Below (in the next section of this chapter) all
parts of the bioreactor are presented, and the assembled device can be seen as well
as the disassembled parts of the system.
184
Secure Bridge Upper scaffold securing Outlet Lower scaffold securing Pressure chamber Diaphragm Air chamber
Figure 4.3.2.2. Schematic diagram of the prototype bioreactor (Damen 2003;
Morsi et al. 2001)
4.3.3. New, simplified Bioreactor The body of the bioreactor is made out of one cylindrical glass vessel, with two
inlets and two outlets. As these are part of the body, this minimizes the number of
attached parts. The top part of the body is a semi-sphere with an outlet on the top to
allow for gas mixture to be controlled and to provide opportunity to vacuum the
bioreactor if needed to ensure no air is present in the system. The dimensions of the
Bioreactor are given below.
185
Gas mixture and vacuum inlet /outlet Glass lid Connection between lid and body of the bioreactor Metal upper ring For scaffold attachment Outlets Silicone tube showing scaffold attachment to the bottom ring Separation between perfusion and inflow chambers Tangential inlets
Figure 4.3.3.1. Simplified Bioreactor: front view with silicone tubing attached
Two rings inside the body, made of medical grade stainless steel, are used to fix the
scaffold. The materials used were chosen to allow easy autoclaving and provide a
reliable environment for tissue growth on the scaffold. The Bioreactor can be
assembled and disassembled in less than 2 minutes, making it easy for non-experts
to operate, and minimizing the risk of contamination related to handling the parts
after autoclaving.
186
The height of the bioreactor without the lid is 180mm, the lower perfusion chamber
is 60mm, and the distance from the bottom of the bioreactor to the outlets is
120mm, the diameter of the body is 100mm and the inlets and outlets have 13mm
diameters.
Both stainless steel rings comprise of an outer ring, which fixes the ring to the
body. These are positioned on specially designed small bumps on the inside wall of
the glass body. The outer ring is connected to the inner ring, to which the scaffold
is fixed with three spikes. They are screwed to the inner rings and only slotted in
the outer rings, thus making it easy to replace the inner ring to fit a different size or
configuration ring for other applications. A detailed view of the rings is shown
below (Figure 4.3.3.2.). A close-up view of the inlets is presented below in Figure
4.3.3.3 to show the principle of media mixing and vortex creation. Photographic
images of the bioreactor lid and the top view (without the lid) of the assembled
bioreactor and of the two rings are shown in Figures 4.3.2.4 and 4.3.2.5,
respectively.
Figure 4.3.3.2. Upper (left) and lower (right) rings with connected spikes
The left spike of the lower ring is unscrewed to show disassembling of the rings.
187
Figure 4.3.3.3. Inlets in front and outlets in the background
Figure 4.3.3.4. Glass lid of bioreactor with the gas inlet/outlet
188
Figure 4.3.3.5. Assembled Bioreactor with both rings without the lid
The Bioreactor is connected to two glass reservoirs via silicone tubing to ensure
circulation and possibility for exchange of the media solution. One of the reservoirs
has a special outlet-inlet on the top for the purpose of media changing. The two
reservoirs and the silicone tubing are easily autoclavable, thus ensuring the
189
biosafety of the tissue culture system. The pulsatility of flow is provided with a
peristaltic pump (Easy-Load™ Master Flex, Millipore Corporation Belford MA
USA, frequency 50/60 Hz, capacity 6-600rpm). This pump operates outside of the
closed sterile system, as the pump pulses on the outer side of the silicone tube
connected to the inlet.
Test runs were carried out to ensure the correct flow of 0.7m/sec was achieved, and
the pump was set to 4.3rpm. A representation of the experimental arrangement is
shown in Figure 3.6.1 in the Methodology chapter.
A technical representation of the lower ring of the bioreactor is given as an
illustration only below (Figure 4.3.3.6). The design can be seen in greater detail
(technical) in Appendix 2 of this thesis.
Figure 4.3.3.6. Technical drawing of the lower ring
190
4.4. Requirements of a bioreactor for tissue culture of blood vessels There are generic requirements (Dumont et al. 2002) which bioreactors used for
tissue engineering need to meet irrespective of the culture application. Below is a
list of the requirements most commonly referred to by researchers in the field of
tissue engineering:
• Sterility
• Inclusion of a scaffold or cell matrix (attachment, compatibility)
• Composition and easy exchange of the growth medium
• Control of gas phase and temperature in the incubator
• Ability for gas exchange between the incubator and bioreactor
• Mechanical stimuli (including pressure, shear stress and pulsatility)
• Diversity of applications
• Low price and easy accessibility
• Simple design and small size (needs to be able to fit in an incubator)
• Easy access to the graft
• Transparency (to allow for monitoring of the graft development)
• Ability to work under a variety of hemodynamic conditions (flow, pressure,
pulse and temperature)
Apart from these general requirements the bioreactor has to comply with the
specific need of the tissue depending on its application. Some of those are
discussed below and have been addressed for the purpose of in vitro growth of
portal veins and portal vein shunts. One of the requirements is to be able to grow
portal veins with different diameters and geometry to account of the variety of
physiology of the portal vein in patients. The bioreactor was designed so that it
could be easily adapted to the required size and was easily accessible for exchange
of the parts, which affix the scaffold. The scaffold, to ensure correct flow
conditions, needs to be fixed at both ends, including any branches.
An important factor for growing veins and other in vitro cultured vessels is
mimicking the environment of the human body while maintaining total sterility.
191
In this project both continuous and pulsatile flows have been simulated, as the
bioreactor needed to be able to work under both conditions.
The system needs to support a constant temperature of 37°C to ensure optimum
culturing conditions. Another requirement is for constant CO2 level of 5%, which
could be achieved either by addition directly to the media, or to the air above the
media in the bioreactor or one of the reservoirs.
Transparent scaffolding material is used to facilitate flow measurements, although
once seeded with tissue culture, those measurements would be difficult to carry out
and cannot be done using LDA.
The fluid chamber and its inlets and outlets have to provide proper mixture of the
growth media so that it flows in a natural way, thus preventing cluttering of growth
media and loose tissue cells in dead corners. To achieve that and the other aims, the
following design adaptations were made:
1) The inlets were made tangentially to one another and are 2 mm above the
bottom of the bioreactor. The lower chamber, where the mixing of the
medium occurs more rapidly due to the preserved entrance velocity, has a
small angle and narrows in diameter prior to opening again to the perfusion
chamber. In future a steeper angle should be studied, possibly the one used
in the prototype bioreactor.
Figure 4.4.1. Joining of lower and perfusion chambers
192
2) Joining of the lower and perfusion chambers (Figure 4.4.1) was made
smooth but also rapid to assist fluid acceleration. Different angles and their
effect on the flow need to be studied to determine the most suitable one.
3) The perfusion chamber is double the size of the lower chamber and had two
outlets situated in the middle of the chamber and on opposite sides.
4) The glass at the point of connection between the body and the lid was
sanded to allow for a closer seal without the need for silicone sealing (if
such a sealant is needed, the design allows this possibility as the two sliding
planes of the body and lid are wide enough to ensure no contact between
sealant and growth medium can occur). It did not leak during the
experiments carried in this research as there was an air cushion above the
liquid medium and the flow did not reach the connection point.
5) The extra outlet/inlet on the top of the lid can be used for vacuuming the
bioreactor if needed or for gas exchange between the bioreactor and the host
incubator, as long as sterility can be maintained, e.g. through the use of an
appropriate filter unit attached to the opening.
6) The inner cylinder (Fig. 4.4.2) of the ring has holes drilled through it to
allow growth media perfusion. The part where the scaffold attaches to the
cylinder is thinner, whereas the area where the spikes attach is thicker for
extra support and strength.
7) The spikes, which connect the outer rings to the inner cylinders, are round
bars, which screw to the inner cylinder only and slide freely in the other
ring (Fig 4.4.3).
193
Figure 4.4.2. Close up of central area where the scaffold is attached
Figure 4.4.3. Inner cylinder and spikes
194
Figure 4.4.4. Bioreactor during LDA measurements
For the purpose of LDA measurements the glass model of the portal vein was fixed
at the bottom ring only, which did not have implications for the stability of the
vessel. If softer scaffold is used, then upper and lower fixation will be needed.
Figure 4.4.4. shows the laser beam passing through the middle of the glass vessel at
the focal point of the laser during LDA measurements.
4.5. Advantages and disadvantages of the bioreactor
4.5.1. Advantages of the new bioreactor design 1. Different bioreactor systems use a variety of methods to prevent leakage in
areas of connection between parts of the bioreactor. Some use adhesive
cement and/or silicone sealant, while others use bolts and metal braces, or a
glue and rubber isolator. The bioreactor developed in this thesis can, if
required (e.g. if the pulsatile flow produces considerable pressure or simply
as an extra precaution against leakage), be sealed from the outside with
silicone or any other material. The sealant will not be in contact with the
media or tissue as the only place where those two parts of the bioreactor
(the body and lid) meet is above the level of the media. The surface where
195
the two parts of the bioreactor join has been mechanically treated and will
adhere well without the need for extra sealing. Those surfaces are wide
enough to allow outside sealing without any risks of the sealant coming in
contact with the sterile environment inside the bioreactor. Silicone grease
can be used to improve the sealing and a gasket or similar device will need
to be used if better leakage protection is needed or the medium reaches the
connection between the body and the lid. This is one of the advantages of
having the main body of the bioreactor as a single part.
2. The body of the bioreactor is made out of a single glass vessel providing
transparency to monitor the tissue growth. It is easy to sterilise and is low
cost to manufacture.
3. On the top of the lid there is an extra inlet/outlet for gas exchange between
the incubator and bioreactor. This inlet/outlet can be sealed when the
bioreactor is taken from a sterile system (e.g. a biosafety cabinet) or if not
needed. This outlet can also be used for vacuuming the system.
4. The two media inlets, which are part of the body of the bioreactor, are made
tangentially to one another and are few mm above the bottom of the
bioreactor. This allows better mixing of the growth media and creates a
small vortex. In this way, the initial velocity of the flow entering the
bioreactor can be preserved, if not accelerated, until the flow reaches the
perfusion chamber.
5. Medical grade stainless steel is used to manufacture the rings, spikes and
cylinders, which secure the scaffold and hold it in place. The inner cylinder
(Fig. 4.4.3) where the scaffold is attached has holes drilled through it to
allow growth media perfusion. In this way nutrients will be able to reach the
ends of the graft enabling complete tissue growth. The spikes, which
connect the outer rings to the inner cylinders, are round bars, which screw
to the inner cylinder only and slide freely in the other ring (Fig 4.4.2). As a
result, the only part that would need to be changed to facilitate a different
diameter, shape or even branching of the scaffold would be the inner
cylinder. The bottom one does not require changing, apart from varying the
196
diameter, but the principle of attachment is same for the bottom and top
cylinders.
6. The bioreactor has only five parts – body, lid, two rings and tubes. The first
two are made of glass, the rings are made of stainless steel, and the tubing is
silicone. All those parts are easy to sterilise, and the system is very easy to
assemble and disassemble.
7. The bioreactor together with the reservoirs can easily fit in a standard
incubator.
8. The system can operate under continuous or pulsatile flow as required for
the specific application.
4.5.2. Disadvantages of the new design From experience, there are only few disadvantages of this device, but extensive
testing in the future might uncover more areas for improvement.
1. Glass is very fragile and extra care is needed when operating and storing the
device.
2. The stainless steel parts are heavy and difficult to fabricate.
3. The inlet/outlet on the lid is difficult to seal as it is smooth (possible
improvement could be a screw top or an attachment onto which a standard
sterile filter unit can be added).
4. There is no provision for heating the growth media and hence the apparatus
must be placed in an incubator.
Some suggestions for future work and optimisation of the bioreactor are outlined
below.
4.6. Future work and optimization of the device Some work and testing needs to be carried out in the following areas:
1. Heating the growth medium either within the bioreactor or while in the
reservoirs or tubes outside the device is needed to ensure optimal tissue
culturing conditions. While this is currently achieved by placing the
197
bioreactor in an incubator, an alternative could be to use a similar system to
that described in the parallel plate bioreactor (Section 4.3.1.1). This utilises
heating of the tubing exiting one of the reservoirs by wrapping it in a 3ft
heating cord.
2. Tests with different angles of the inlets need to be carried out to determine
the best one from a hemodynamic point of view.
3. Different joining angles between the lower and perfusion chambers and their
effect on the flow need to be studied to determine the most suitable one.
4. Creating a screw lid for the inlet/outlet of the bioreactor’s lid to allow for
better sealing.
5. Designing and testing of other types and different shapes of scaffold.
4.7. Conclusion For the purpose of the research carried out in this thesis, a novel bioreactor with a
simplified design had been developed. The small number of parts allows easy
assembling by non-experts. The materials used (glass, medical grade stainless steel
and silicone tubing) can be sterilised by commonly used methods, including
autoclaving.
Silicone sealant can be used on the outer surface of the connection between the
body of the bioreactor and the lid to prevent leaks after the system has been
sterilized and assembled.
The device can be easily modified to accommodate vessels of different size,
diameter and branching geometry, by simply exchanging the top cylinder to which
the scaffold is attached.
The maximum cost of manufacture of the bioreactor is AUD 500. This cost will be
lower if a larger quantity is manufactured. The bioreactor can be easily
decontaminated, washed and re-used as required.
The design allows for control of the flow conditions and media mixture (including
media exchange). There are no moving parts within the bioreactor.
198
The durability of the medical grade stainless steel parts have been tested by being
left in a glycerol solution for over two weeks without any indication of rusting. The
first prototype bioreactor (Section 4.3.2.) showed rusting within the first week of
exposure to glycerol solution and this test made sure the new device would not face
similar problems.
Nevertheless, there are several areas, as outlined above, which need to be
investigated further and could lead to improved performance of the apparatus.
199
200
CHAPTER 5
Measurements, Simulations and Results
In this chapter the results from Laser Doppler Anemometry measurements of
simulated flow and the corresponding FLUENT computer simulations of the
same vessels are presented and described. The physical experiment has been
designed based on the computer model to allow for comparison of the results
using both methods. The CFD computer model has been simulated using a
variety of parameters, including Newtonian and non-Newtonian flow, low and
high velocity, and adding the effects of gravity, pressure and predetermined
outflow from each outlet (branch). The model used for the LDA measurements
had non-Newtonian flow simulated, and was compared to the non-Newtonian
results from the computer simulation. FLUENT provides information on many
parameters and combinations between relationships of parameters, and some of
the opportunities for obtaining a variety of information will be presented later in
this chapter. Some suggestions for optimization and future work are also given at
the end of the chapter.
5.1. Geometry and Grid generation
5.1.1. Grid generation
Research groups in the last decade have used a variety of computational cell
numbers in blood vessel modelling. That number depends on the shape of the
vessel, the complexity of the flow and the need for accuracy. For modelling of
real-life carotid bifurcation, the mesh was generated using CFX4 code, after the
refinement was set to 17,920 eight-noded cells, as a further decrease in the mesh
size did not show any significant difference in velocity prediction (Starmans-
Kool et al. 2002). For modelling of a vascular tube with two stenotic areas,
Gurlek et al., used between 17,000 and 26,000 triangular computational cells in
FLUENT (2002). A four-noded cell volume grid containing 1,600 for normal and
1,800 for stenotic vessel cells, respectively, has been used in simulations of
201
pulsatile arterial flow (Marques et al. 2003). Other authors (Moore et al. 1998)
have used 18,000 to 51,000 cells depending on the resolution of MRI to achieve
mesh independence.
The grid used in this study contained 27,457 computational tetrahedral and
hexahedral grid cells for the normal non-obstructed model, and 60,731 cells for
the obstructed model (Petkova et al. 2003). Although that number might seem
large, this is the optimal number of computational cells for solving the flow
problem, above which no improvement in accuracy can be achieved.
5.1.2. Scaling
Many studies use scaling of the model blood vessel to allow for easier
measurements (Bonert et al. 2003; Yedavalli et al. 2001). Although scaling can
be necessary to visualize very small vessels (Jalan et al. 2004), it was not found
to be imperative in measurements of portal vein model vessel. When a model has
been scaled up, the flow pattern will be different because the shear rate will
change with the change in the diameter. Hence, the ‘cell-free’ wall zone will take
a smaller percentage of the cross-section area of the flow and the viscosity will
change. Some of these discrepancies can be overcome mathematically, but the
intention here was to develop a simple, easy to use model and, hence, no scaling
was applied in the experimental or computer simulation work. The experimental
vessel was manufactured according to one that was computer generated and is
identical (as far as possible) to the vessel used in solving the computer flow
problem.
5.2. Model assumptions
5.2.1. Geometry
5.2.1.1. 3-D geometry
In the model developed here, planar geometry has been assumed for simplicity
and to generalise the findings. In recent years the importance of real-life, non-
planar geometry has attracted much attention and the work presented below in
this thesis needs to be repeated with a replica of the natural portal vein without
202
the idealisation. The only problem with this approach is that the model has to be
individually tailored to each patient to be realistic. This will increase the cost and
time of modelling significantly, but is needed to further our understanding of
hemodynamics within the portal circulation. Planar bifurcation is an acceptable
assumption (Starmans-Kool et al. 2002), and rigid (non-compliant) walls are
widely used even in artery models (Siro et al. 2002; Starmans-Kool et al. 2002;
Xu et al. 1999) although this assumption affects the accuracy of shear stress
values.
5.2.1.2. Size, diameter and branching
The current model has been created as a symmetric branched vessel, where the
branches on the left side are identical to (or very much alike) the branches on the
right.
A circular diameter has been assumed in most parts of the vessel, with the
exception of the areas where the diameter size changes or branching occurs, and
in the areas of obstruction. The reason behind creating a simplified model of the
portal vein is to allow for 2-D LDA measurement using available equipment
(which can be found in most fluid research laboratories) and to produce a base
model, upon which future patient-specific models can be created. The two
vessels are identical apart from the areas of blockages to research into the impact
those areas have on the flow behavior.
The dimensions of the models are given in Table 3.3.4.1 in the Methodology
chapter, and are as follows: the inlet diameter is 10mm, primary branches have
8.5mm diameters, and the diameters of the secondary branches are 6.375mm.
The total height of the vessel is 91mm.
5.2.2. Flow
5.2.2.1. Common flow assumptions
Laminar flow is assumed in this thesis, which is a commonly used assumption in
CFD modelling (Gurlek et al. 2002; Starmans-Kool et al. 2002; Xu et al. 1999).
203
Fully developed axial flow is a commonly used inlet assumption (Gurlek et al.
2002; Jensen 1996; Moore et al. 1998), and constant pressure as the outlet
boundary condition can be used (Gurlek et al. 2002).
Some authors assume zero velocity, since the initial conditions are unknown
(Marques et al. 2003).
Newtonian flow is an accepted assumption in large vessels, as discussed in the
Methodology chapter earlier, and this work has been carried out to simulate both
Newtonian and non-Newtonian flow behaviour. A brief presentation on
Newtonian flow is given later in this chapter, but more emphasis is given to non-
Newtonian simulations, as the current model is capable of solving those
problems.
5.2.2.2. Observations
When negative values of shear stress are recorded at a point, this is an indication
that the shear stress there is in the direction of the flow (Starmans-Kool et al.
2002).
During Laser Doppler Anemometry (LDA) measurements, it was observed that
neighboring points in the vessel showed very different flow behavior. Some
areas, where the flow was quickly accelerating, were next to areas of slower,
even reversed flow in some instances. On some occasions, flow occurred in a
different direction. The selection of points for the LDA measurements has been
made in a way to ensure more points are selected in and around areas of
blockages and branching. All points are measured on the Z=0 plane and each row
of points has constant Y value. The details of the points (including the measured
values) are shown in appendix 5 of this thesis alongside comparisons between the
measured and computer generated flow at those points. The LDA measurements
were carried first out, and then the same points were selected in the computer
model to cross-check their behavior.
In the computer model, the definition of the different phases in the stream is very
important, as the size and quantity of each phase affects the flow significantly.
Hence, another factor, which appears to affect the blood flow within the portal
204
vein, is the number and size of blood cells in the patient. It would be worth
examining the effects of the blood composition on the flow in other vessels, as
this can prove important not simply for predicting areas more likely to attract cell
deposition and obstructions, but for the overall flow properties. Some of those
visualisations are presented later in this chapter in the section on particle
tracking.
5.3. Benefits and limitations
CFD simulations offer superior accuracy (Starmans-Kool et al. 2002), but are
difficult to be validated as there is no set standard, and the models are
individualised to correlate with the in vivo data of only one patient.
Studies concentrating on the errors of modelling are not comprehensive (Moore
et al. 1998) and are an area requiring extensive research.
Physical modelling, based on data obtained from in vivo measurements, can also
provide valuable information to physicians about blood flow pattern, wall shear
stress and areas most likely to be affected by thrombosis, (Yedavalli et al. 2001)
but creating such models is more time-consuming and usually requires scaling of
the vessel.
Comparison between physical experiments and computer modelling are difficult
to make, but are beneficial and necessary for validation of the simulations.
Identical conditions in the two models (Henry et al. 1997) are needed and vessel
geometry has to be much the same in both. The particle paths in the numerical
model can be compared to the visualised flow pattern in the physical model
(Henry et al. 1997). In this study, comparisons of the two simulations have been
made and the results are presented below. As outlined above, the shape and size
of the model used in the LDA measurements was determined using the
parameters from the computer-generated model, and are nearly identical.
5.4. Visualisation
In this thesis, FLUENT visualisation of velocity contours and wall shear stress
distribution in both normal and stenotic vessels was modelled, in accordance to
Gurlek et al. (2002) and Petkova et al. (2003). Details on the model and fluid
parameters are given in the previous chapter.
Many factors affect the flow through a stenotic area according to CFD
simulations, including Reynolds number, structure geometry and size of
obstruction, viscosity, flow velocity and pressure.
Comparisons between stenotic and normal flow using the same flow conditions
(Marques et al. 2003; Petkova et al. 2003) are presented in Section 6 of this
chapter.
Figure 5.4.1. LDA visualisation experiments of obstructed vessel in bioreactor For LDA experiments, the laser beam was focused inside the inner left sub-
branch of the model while flow measurements are made, as seen in Figure 5.4.1.
One of the difficulties in those measurements was the round surface of the vessel,
which reflected the light and did not allow for measurements immediately next to
the wall. In future, the glass model will need to be modified to aid LDA as
follows:
205
* the shape needs to have 90° angles, similar to those of the tank in which the
bioreactor was immersed, to allow easy laser penetration and to minimise light
reflection
* the inside shape of the vessel needs to be preserved to represent realistic flow.
The cross section of future designs of the glass LDA model may need to be
something like the one presented below in Fig. 5.4.2.
Figure 5.4.2. Cross-section of the inner diameter of an ideal glass vessel and
the outside square shape
Figure 5.4.3. LDA visualization experiments with normal (non-obstructed)
model in bioreactor
206
Figure 5.4.4. Laser beam through the glass model – side view
The bubbles that form in the water solution (water plus additives to simulate
blood viscosity plus reflector particles) ‘stick’ to the wall and their impact on the
measurements is minimal (Figs. 5.4.3 and 5.4.4). However, regular removal of
the bubbles was undertaken to improve the accuracy of the measurements.
Figure 5.4.5. LDA experimental setup
207
The bioreactor was placed inside a large tank filled with water to minimise the
reflection of the laser beam (Figs. 5.4.5 and 5.4.6). The bioreactor itself was
filled with a clear solution mixed so as to mimic the viscosity of blood while
being transparent to aid measurements. Refracting particles (Meta DC Coated
Particles (Model 10037)) were added to the solution. The vessel glass model was
mounted inside the bioreactor, and the flow velocities within it were measured. A
black background was placed opposite the laser to minimise the effects of beam
reflection from the far wall of the tank. The pump and two reservoirs were
situated outside the tank. Due to the unavailability of a movable platform onto
which the laser head could be mounted and then moved to measure different
points in the model, manual elevation and movement in the x-direction had to be
carried out, thus increasing the percentage of error. With more complete or
advanced machinery, the measurements need to be repeated for accuracy.
Figure 5.4.6. Water tank with bioreactor submerged in it, the reservoir
behind (top of photo behind the tank) and LDA on the right of the tank
208
5.5. Mathematics and parameters
The values of k and n used in this study were obtained from data of shear rate
versus shear stress by Syoten Oka (Syoten 1981, page 35, taken from Huang et al
1975) and are graphically presented below in Figure 5.5.
This relationship determines the changes in the viscosity of whole human blood,
and in this case has been obtained using a Weissenberg Rheogoniometer.
Those values are obtained by mathematical conversion using log functions,
where Series 1 is based on the parameters as per reference (Syoten 1981) and
Series 2 are the values after conversion. One of the main reasons for using the
log function was to obtain the values of k and n for which the angle between the
line of points (Series 2) and the x-axis had to be known. Those values were used
in setting the CFD model parameters using the non-Newtonian Power Law.
y = 2 .0727x 0.48 51
y = 0 .2073x 0.4 85 1
0
1
2
3
4
5
6
7
0 2 4 6 8 10 12
S ta in R a te
h e a r S tre ss
8
S
S eries1S e ries2P ow er (S e ries1 )P ow er (S e ries2 )
Figure 5.5. Calculating k-n parameters for use in the Power Law equation in FLUENT Figure 5.5. represents the relationship between Shear Stress (y axis) and Stain
Rate (x axis) for whole blood.
209
The points were calculated based on the experimental values as outlined in
(Syoten 1981) and are presented in the table below (after mathematical
conversion).
X axis 0.7; 1.45; 2.1; 2.95; 3.6; 4.25; 5.6; 7; 8.3; 9.6
Y axis 0.18; 0.245; 0.31; 0.35; 0.37; 0.398; 0.445; 0.52; 0.6; 0.68
5.5.1. Continuity and Momentum Equations
For all flows, FLUENT solves conservation equations for mass and momentum.
For flows involving heat transfer or compressibility, an additional equation for
energy conservation is solved. For flows involving species mixing or reactions,
species conservation equations are solved or, if the non-premixed combustion
model is used, conservation equations for the mixture fraction and its variance
are solved. Additional transport equations are also solved when the flow is
turbulent (Petkova et al. 2003). In this section, the conservation equations for
laminar flow (in an inertial, non-accelerating, reference frame) are presented.
The Mass Conservation Equation
The equation for conservation of mass, or continuity equation, can be written as
follows:
210
( ) mSvt
=⋅∇+∂∂ rρρ
(5.5.1.1.)
Equation (5.5.1.1.) is the general form of the mass conservation equation and is
valid for incompressible as well as compressible flows. The source Sm is the mass
added to the continuous phase from the dispersed second phase (e.g., due to
vaporization of liquid droplets) and any user-defined sources. In this case it is 0.
Momentum Conservation Equations
Conservation of momentum in an inertial (non-accelerating) reference frame is
described by
211
( ) ( ) ( ) Fgpvvvt
rrrrr++⋅∇+−∇=⋅∇+
∂∂ ρτρρ
(5.5.1.2.)
Where p is the static pressure, τ is the stress tensor (described below), and
and
gr
ρ
Fr
are the gravitational body force and external body forces respectively.
is given by The stress tensor τ
( ) ⎥⎦
⎤⎢⎣
⎡⋅∇−+∇= Ivvrr
32μτ ∇v T
r
(5.5.1.3.)
Where μ is the molecular viscosity, I is the unit tensor, and the second term on
the right hand side is the effect of volume dilation.
5.5.2. Viscosity equations
Viscosity for Non-Newtonian Fluids
For incompressible Newtonian fluids, the shear stress is proportional to the rate-
of-deformation tensor : D
Dμτ =
(5.5.1.4.)
Where D is defined by
⎟⎟⎠
⎜⎜⎝
⎛
∂∂
+∂
∂=
j
i
i
j
xu
xu
D⎞
(5.5.1.5.)
And μ is the viscosity, which is independent ofD . For some non-Newtonian
fluids, the shear stress can similarly be written in terms of a non-Newtonian
viscosity η:
212
( ) DDητ =
(5.5.1.6.)
In general, η is a function of all three invariants of the rate-of-deformation
tensorD . However, in the non-Newtonian models available in FLUENT, η is
considered to be a function of the shear rate γ& only. γ& is related to the second
invariant of D and is defined as
γ& DD :=
(5.5.1.7.)
FLUENT provides four options for modelling non-Newtonian flows:
• Power law
• Carreau model for pseudo-plastics
• Cross model
• Herschel-Bulkley model for Bingham plastics
In the simulations presented in this thesis, the non-Newtonian power law was
used and is described below.
Power Law for Non-Newtonian Viscosity
The non-Newtonian-power-law model was used in this study (FLUENT 6.0
Manual, Chapter 7.3.5), where the non-Newtonian viscosity is calculated as:
TTn ek 01−= γη &
(5.5.1.8.)
FLUENT allows upper and lower limits to be placed on the power law function,
yielding the following equation:
maxmin1 0 ηγηη <=< − TTn ek &
(5.5.1.9.)
where k, n, T0, ηmin, and ηmax are input parameters.
k is a measure of the average viscosity of the fluid (the consistency index); n is a
measure of the deviation of the fluid from Newtonian (the power-law index) (as
described below); T0 is the reference temperature; and ηmin and ηmax are the
lower and upper limits of non-Newtonian viscosity used in the power law,
respectively. If the viscosity computed from the power law is less than ηmin, the
value of ηmin will be used instead. Similarly, if the computed viscosity is greater
than ηmax, the value of ηmax will be used instead. Table 3.3.4.2 in the
Methodology chapter shows how viscosity is limited by ηmin and ηmax at low and
high shear rates in this model. The value of n determines the class of the fluid:
n = 1 Newtonian fluid
n > 1 shear-thickening (dilatant fluids)
n < 1 shear-thinning (pseudo-plastics)
The outflow at each outlet was predetermined, and the weighting percentage is as
follow: outlets 1 and 2, 15% each; outlets 3 and 4, 35% each. Outlets 1 and 2 are
on the left side and 3 and 4 on the right hand side in this model. This was done
not to represent medical conditions, but to show the possibilities available to
manipulate the model on demand (if severe pressure is applied in only one of the
branches, for example, due to stenosis, or fibrosis of one lobule only). From a
physiological point of view it is expected that the branch entering the larger liver
lobe will have a higher outflow rate.
213
5.6. FLUENT models: simulation results
5.6.1. First model visualization
Before the more complex and idealised model used in this thesis was developed,
the computer model was tested with a simple structure as presented below. The
purpose of this modelling was to explore possibilities and to show the multiple
applications of the work (i.e. that this model can be converted to any other part of
the circulatory system).
Figure 5.6.1.1. 3-D Grid of simplified blood vessel structure
The material (fluid) used was water-viscose (water with added viscosity).
To represent the viscosity, the non-Newtonian-power-law was used (and for
consistency with the realistic simplified model as shown above in this chapter)
with the following parameters: K=0.001 and N= 0.7, the temperature was set at
(K) = 310o, the minimum viscosity was 0.001 and the maximum viscosity was
0.01. The flow was assumed to be laminar, the energy equation was activated and
the default operating conditions of FLUENT were used.
214
Figure 5.6.1.2. Contours of static pressure (MPa) in a vessel, assuming
identical outflow from both branches
Figure 5.6.1.3. Velocity vectors coloured by velocity magnitude (m/s) in a
vessel, assuming identical outflow from both branches
215
Figure 5.6.1.4. Contours of wall shear stress (Pascal) in a vessel, assuming
identical outflow from both branches
Figure 5.6.1.5. Contours of boundary cell distance in a vessel, assuming
identical outflow from both branches
216
Depending on the information one desires to obtain from the model,
visualisations on a variety of planes “cutting through” the vessel can be made.
Those planes can be parallel to one of the axes, or under a customised angle if a
cutting plane representing the flow parameters in a different plane is required (for
example, a cutting plane under 45o to the x-axis).
Once created, the planes can be used to represent any given parameter such as
pressure, velocity or stress, or any combination of such parameters.
Below (Figs. 5.6.1.6 and 5.6.1.7) are examples of multiple planes parallel to the
x-axis for velocity vectors coloured by velocity magnitude.
Figure 5.6.1.6. Cutting planes parallel to the x-axis of velocity vectors
coloured by velocity magnitude
217
Figure 5.6.1.7. Cutting planes positioned within the model representing contours of velocity magnitude To achieve reliable results, a minimum number of 100 iterations are needed
before conversion of the results (i.e. before the equations are solved). Below (Fig
5.6.1.8) is a representation of the residual conversion for this model, including
continuum equation, the velocity in direction of the three axes and the energy
equation. In the other models discussed in this thesis similar numbers of
iterations were applied.
Figure 5.6.1.8. Scaled Residuals
218
219
5.6.2. Comparison between models with and without obstructions
The outflow was pre-defined with flow rate weightings of 0.15 for each of outlets
1 and 2 (left branch) and 0.35 for each of outlets 3 and 4 (right branch). In both
models the same parameters were used and the following results were obtained
(Figures 5.6.2.1-5.6.2.4). Differentiation of the flow rate in the outlets was used
based on the fact that the two lobes of the liver are of different size (the right
lobe is over 6 times larger than the left one (Gray 1995)), but mainly because the
effects of resistance from different parts of the liver on the blood flow were
considered.
For all figures below, (a) represents simple, non-obstructed vessel, whereas (b)
shows the flow in a model with additional obstructions.
Figures 5.6.2.1(a) and 5.6.2.1(b) show velocity contours on an x-y plane cutting
through the middle of the geometry. Figures 5.6.2.1(a) and 5.6.2.1(b) show that
there are significant changes in the velocity magnitude in the two models. This
model was created assuming that the problem causing portal hypertension
originated in the liver, such that the portal venous flow was diminished and the
portal vein pressure was 3922.66 Pascal (40 cm H2O). In this case there are two
possible conditions: Figure 5.6.2.1(a) shows the velocity magnitude if there were
no additional complications in the portal vein, whereas Figure 5.6.2.1(b) shows
the additional complication of portal vein thrombosis and the impact on blood
flow. In this model, there is no additional decrease in velocity, and so it gives the
most favourable picture of this condition (i.e. hypertension). Figure 5.6.2.1(c)
shows a closer view of the obstructed area of Figure 5.6.2.1(b).
Figure 5.6.2.1.(a) Contour of velocity magnitude on an x-y plane cutting through the middle of the geometry (Z=0 plane) without obstructions
Figure 5.6.2.1.(b) Contour of velocity magnitude on an x-y plane cutting through the middle of the geometry (Z=0) with obstructions
Figure 5.6.2.1.(c): Closer view of the contour of velocity magnitude on Z=0
plane with obstructions.
220
Figures 5.6.2.2(a) and 5.6.2.2(b), below, show pressure contours on an x-y plane
cutting through the middle of the geometry (Z=0 plane). These correlate with the
findings of Figures 5.6.2.1(a) and 5.6.2.1(b). The zones of low pressure, which
were typical for the two higher flow outlets (Figures 5.6.2.2(a) and 5.6.2.2(b)),
have “moved” to the areas of the obstructions, thus increasing the pressure in the
low flow outlets (Figure 5.6.2.4(b)).
Figure 5.6.2.2.(a) Contour of static pressure on Z=0 plane of the geometry
without obstructions
Figure 5.6.2.2.(b) Contour of static pressure on Z=0 plane of the geometry
with obstructions
Figures 5.6.2.3(a) and 5.6.2.3(b), below, show the contour of strain rate on an x-y
plane, cutting through the middle of the portal vein. In Figure 5.6.2.3(a) the strain
221
rate is uniform through the portal vein according to the velocity distribution,
whereas in Figure 5.6.2.3(b) the contour of strain rate is not uniform because of
the existence of obstacles in the blood flow. As expected, the strain rates are
higher in the constriction created by the obstruction. The colour charts are not
similar, and it needs to be noted that the top red area of the scale in figure (a) has
the same value as the middle of the scale (light blue) in figure (b), showing that
the high strain rates in the normal vessel are in the middle range of the model
with the obstructions.
Figure 5.6.2.3.(a) Contour of strain rate on Z=0 plane cutting through the
middle of the geometry without obstructions
Figure 5.6.2.3.(b) Contour of strain rate on Z=0 plane of the geometry with
obstructions
222
Figure 5.6.2.4.(a) Contour of wall shear stress on Z=0 plane without
obstructions
Figures 5.6.2.4(a) and 5.6.2.4(b) show contour of wall shear stress near the wall
on an x-y plane cutting through the middle of the geometry (Z=0). The
obstructions in the portal vein decrease the available cross section area. This
reduction of available cross section area ultimately introduces higher strain rates
around the obstructions. The higher strain rates around the obstructions results in
significantly higher shear stress near the wall, as presented in Figure 5.6.2.4(b).
The portal vein without obstructions (Figures 5.6.2.3(a) and 5.6.2.4(a)) show
much lower values of strain rates and wall shear stress.
Figure 5.6.2.4.(b): Contour of wall shear stress on Z=0 plane with
obstructions
223
5.6.3. FLUENT comparisons of different velocities
Below is an example of Newtonian flow through the normal (no obstructions)
model visualized for two different velocities. Strain rate is higher the higher the
velocity in the not obstructed model.
Figure 5.6.3.(a). Contours of strain rate when velocity is set at 0.0015m/s
Figure 5.6.3.(b). Contours of strain rate when velocity is set at 0.0225m/s
224
5.6.3.1. Velocity magnitude
One example with the obstructed vessel is shown below, where all model
parameters are identical and only the velocity has been varied. Figure 5.6.3.1.1
represents simulations in which the velocity is 0.07m/s while the velocity in
Figure 5.6.3.1.2 is less than 25% of the first – 0.015m/s. Some of the readily seen
differences are in the left sub-branches, and the outer right sub-branch, next to
the obstruction in the left main branch, and the obstructed area in the main
vessel.
The scale on the left hand side of the graph is considerably different for both
velocities, and is not uniform. This was done to make full use of the colours that
represent the velocity.
Figure 5.6.3.1.1. Z=0 plane contours of velocity magnitude when velocity is
simulated at 0.07m/s
225
Figure 5.6.3.1.2. Z=0 plane velocity magnitude when velocity is simulated at
0.015m/s
5.6.3.2. Visualization opportunities with FLUENT
Below are some examples of the different visual representations achieved using
FLUENT. Depending on the objective, one can visualize close to the vessel wall
(with or without the grid being displayed), in a “cutting plane” (horizontal,
vertical or under a certain angle) or in the default interior of the model vessel. In
this thesis, the Z=0 plane was used most often, as it “cuts” the vessel through the
middle, dividing it onto anterior and posterior parts. Each parameter can be
varied individually to represent different pressure, velocity or outflow. Another
available option is to vary the two parameters one wants to compare (for
example, velocity vectors coloured by X, Y or Z velocity, or by static pressure,
or stress). Figures 5.6.2.3.1 to 5.6.2.3.5 have used a velocity assumption of
0.07m/s, and the following figures (5.6.2.3.6 and 5.6.3.2.7) represent a velocity
assumption of 0.015m/s.
226
Figure 5.6.3.2.1. Velocity vectors at the wall with grid coloured by velocity
magnitude (m/s) when velocity is simulated at 0.07m/s
Figure 5.6.3.2.2. Velocity vectors at the wall without the grid coloured by
velocity magnitude (m/s) when velocity is simulated at 0.07m/s
227
Figure 5.6.3.2.3. Contours of the wall shear stress (Pascal) (at the wall)
Figure 5.6.3.2.4. Velocity vectors in the Z=0 plane coloured by Y velocity
(m/s)
228
Figure 5.6.3.2.5. Velocity vectors in Z=0 plane coloured by static pressure
(Pascals)
Figure 5.6.3.2.6. Histogram of frequency of velocity magnitude (with velocity assumption of 0.015m/s) This histogram represents the distribution (frequency) of the velocity magnitude
through the vessel, showing that although a velocity of 0.015m/s has been set, it
is not uniform throughout the vessel. The most common velocity was the one
predetermined, but a lower velocity of 0.005 to 0.0075 is frequently observed.
229
There are areas of higher than the average velocity, with some areas where the
velocity is double that of the average (i.e. 0.03 and above).
(mm) Formatted: Font: 9 pt
Figure 5.6.3.2.7. Static pressure verses position in the model in the Z=0 plane Figure 5.6.3.2.7 represents the static pressure (y axis) within the vessel (x axis) at
Z=0 plane. The 0 point on the x axis is the middle of the vessel, and as most of
the outflow has been set to flow through the right branch, there is an increase in
static pressure in the left branch. Those figures are given only to illustrate the
range of information which can be obtained using the developed computer
model.
5.6.4. Particle tracking
FLUENT allows for particle tracking after the flow has been set as multiphase
and the size and concentration of each phase has been determined. For the
purpose of these simulations, the size and frequencies were determined based on
data from the literature as presented in the methodology chapter (3.4.1
Rheological properties of human blood). Four phases were simulated –
leukocytes, erythrocytes, platelets and plasma - and particle tracking for each of
them (when the velocity is assumed at 0.07m/s) is presented below. The particles
are assumed to enter from the inlet, and the wall is assumed to be a non-moving,
230
rigid surface. The colour scheme is uniform to highlight the differences between
the four phases.
Figure 5.6.4.1. Leukocyte particle traces coloured by velocity fraction
Figure 5.6.4.2. Erythrocyte particle traces coloured by velocity fraction
231
Figure 5.6.4.3. Platelet particle traces coloured by velocity fraction
Figure 5.6.4.4. Plasma particle traces coloured by velocity fraction
The four figures above (5.6.4.1 to 5.6.4.4) are using the same scale and show the
difference in the velocity and particle distribution of each of the four phases.
They illustrate that each phase has a different behavior, for example the
leukocytes show highest overall velocity than the other phases, and they are
232
moving fastest to the left side of the vessel and slowest on the right side. The
platelets and erythrocytes have their respective highest velocities through the
right side of the main vessel and the two right oriented sub-branches of the left
and right branches. Such information can be useful if a particular phase (be that
blood cells of other particles carried by the blood) are know to ‘stick’ to the
vessel wall and cause obstructions.
If needed, each particle in each phase can be tracked separately, and below is an
example of five different particle streams, all of which are leukocytes.
Enlargements of these figures are presented enlarged in Appendix 3, and all have
uniform scale (left hand side).
(a) (b) (c)
(d) (e)
Figure 5.6.4.5. Single stream particles
5.6.5. Newtonian verses non-Newtonian flow
Below (Figs 5.6.5.1-5.6.5.4) are comparisons between Newtonian and non-
Newtonian flow when all the other parameters are equal. For the Newtonian flow
the viscosity is assumed constant at a value of the average of the minimum and
maximum viscosities used in the non-Newtonian model.
The visualisations below are presented in a Z=0 plane, which is the cutting plane
along the Y coordinate at the Z=0 position. As a result, flow near to the walls can
be visualised and a more comprehensive picture of the overall flow pattern can
be obtained.
233
The main difference between the contours of velocity magnitude (Figures
5.6.5.1. and 5.6.5.2.(c)) is the intensity of the colours, which is higher in the non-
Newtonian model.
Figures 5.6.5.2. (a) and (b) are given to illustrate the importance of using the
correct scaling for obtaining the maximum information about the flow behaviour.
The auto scale in this case allows the available information to be visualised so as
to provide information about the regions of highest velocity magnitude, whilst
the scaled (figure 5.6.5.2. (b)) shows the velocity magnitude in the branches and
trunk but not at the critical areas where it displays black fields (at the
obstructions).
Figure 5.6.5.1. Contours of velocity magnitude (m/s) for a Newtonian flow in
Z=0 plane
234
Figure 5.6.5.2.(a) Profiles of velocity magnitude (m/s) non-Newtonian auto scale
Figure 5.6.5.2.(b) Profiles of velocity magnitude (m/s) scaled for non-
Newtonian flow
235
Figure 5.6.5.2.(c) Contours of velocity magnitude (m/s) for non-Newtonian
flow in Z=0 plane (same as Fig. 5.6.3.1.1.) (Velocity = 0.07m/s)
Figure 5.6.5.3. Contours of velocity (m/s) in Z=0 plane for Newtonian flow
236
Figure 5.6.5.4.(a) Profiles of velocity magnitude (m/s) for non-Newtonian
flow in Z=0 plane
Figure 5.6.5.4.(b) Contours of velocity (m/s) in Z=0 plane for non-Newtonian
flow (same as Figure 5.6.2.(b) in section 6 of this chapter)
237
238
From Figures 5.6.5.3 and 5.6.5.4 (a & b), it can be seen that the flow direction
has not changed, and the velocity intensity is similar. The differences lie in the
following areas:
• Left outer sub-branches – in the Newtonian model the dark blue zone on
the outside is thick, but on the inner (top) side it is not present at all. In
contrast, the non-Newtonian model displays a more even distribution on
the bottom and top of the sub-branch.
• In the beginning of the left branch, just before the obstruction, the low
velocity zone is larger and easier to identify in the non-Newtonian model.
At the same time, the middle (inner) wall of the Newtonian model shows
a small area of low velocity close to the branching point, which is not
present in the other model.
• Further down, examination of the left side of the main vessel in the
Newtonian model reveals an area of higher velocity behind the
obstruction, which is limited to the area closest to the thrombi, whereas in
the non-Newtonian it spreads out upwards to the right branch.
• The differences in the right branch are again the elongation of the higher
velocity area in the non-Newtonian model behind the obstruction. That
area of higher velocity has a definite upwards direction towards the point
of sub-branching in the non-Newtonian model, whereas in the Newtonian
model it ‘points’ towards the right (outer) sub-branch. Overall, low
velocities can be seen close to the wall throughout the non-Newtonian
model, possibly indicating areas of a wall slip fluid layer.
Figure 5.6.5.5. Contours of wall shear stress (Pascal) for Newtonian flow in
Z=0 plane
Figure 5.6.5.6. Contours of wall shear stress (Pascal) for non-Newtonian flow in Z=0 plane
5.6.6. Idea for portal vein shunt
In this section the visualization of simulations were carried out using a novel
shaped shunt, designed to by-pass areas within the portal vein which are
obstructed. The inlet of the shunt is to be connected to the main portal vein,
whilst the outlets are to connect to the respective sub-branches. This shunt has
239
not been manufactured and the results of the simulations have not been verified
using LDA or another technique and those would constitutes part of future work
as separate research. Neither the anastomoses upstream and downstream, nor the
bending (angle) of the shunt between the inlet and outlet have been modelled.
This modelling was done to illustrate the versatility of the model and to provide
ideas for developing shunts for patients with severe occlusions of the portal vein.
5.6.6.1. Non-Newtonian flow visualization
In the results presented in this section the following parameters have been used
to solve the non-Newtonian power law equations: pressure = 26664.478 Pascal,
velocity magnitude 0.07 m/s, density 1050 (kg/m3), viscosity range 0.0125 to
0.03 (kg/m-s). The temperature, consistency coefficient and all other parameters
used are the same as in all simulations presented in the FLUENTS models
section of this chapter. It needs to be noted that the outflow in this shunt is the
same as in the modelled vessel above, i.e. the majority of outflow is through the
right branches (Figure 5.6.6.1.1.). The flow at the inlet is fully developed. The
figures below, unless otherwise stated, are representation of the flow in Z=0
plane.
Figure 5.6.6.1.1. Contours of velocity magnitude (m/s)
240
Figure 5.6.6.1.2. Contours of wall shear stress (Pascal)
Figure 5.6.6.1.3. Contours of wall shear stress (MPa)
Figure 5.6.6.1.3 above is a representation of the wall shear stress at the wall, and
shows the areas of branching as higher shear stress area, with the highest values
recorded at the right branch junction.
241
The following are visualizations of the static pressure in different planes. First is
the Z=0 plane, followed by the wall and the default interior of the vessel. From
here the non-uniform behaviour of the flow and the difference between the
middle and wall of the vessels is clear.
Figure 5.6.6.1.4. Contours of static pressure
242
5.6.6.2. Newtonian flow visualisation
In this section all parameters are the same as in section 5.6.6.1, except for the
viscosity, which is set at constant 0.02 (kg/m-s) and the deactivation of the power
law equations.
Figure 5.6.6.2.1. Velocity vectors coloured by velocity magnitude (m/s) in the
default interior
Figure 5.6.6.2.2. Contours of velocity magnitude (m/s) at Z=0 plane
243
Figure 5.6.6.2.3. Contours of Static pressure (Pascal) in the default interior
5.7. Visualisation of LDA measurements
5.7.1. Measurements and different visualisation opportunities
For LDA measurements, the laser beam was focused at the Z=0 plane (the
middle imaginary plane which ‘splits’ the vessel into front and back equal parts).
The laser head had to moved manually to measure different points at selected
Y=0 planes, i.e. along the x-coordinate. After completing measurements on each
plane, the laser head was elevated along the y-coordinate and the measurements
along the x-coordinate were conducted for that plane. Due to the very low rate of
the measurements (on average one hour per point), the points were substantially
distributed, and measurements conducted on few planes only (i.e. those where
244
245
differences in the flow were expected to occur according to the computer
models). The reasons for doing so were that:
• Each time the laser beam was focused, there was no absolute guarantee
that it was precisely positioned on the same plane as the one on which it
was focused earlier (resulting from the unavailability of a movable stent
to which the laser head needs to be attached and the manual positioning,
which accounts for human error)
• The water in the tank deteriorated in terms of transparency because of
dust particles and needed changing every 24 hours; and
• The formation of air bubbles inside the bioreactor as the device was not
airtight so as to allow exchange of the viscous fluid and of the vessel
model used (normal or obstructed). The bubbles needed to be removed
regularly as they scattered the light, which made it difficult to conduct the
measurements.
Following the recording of velocities in different points of both normal and
obstructed models, the files were extracted and plotted with Tecplot software.
Some of the Figures below have different vector lengths, to better show the flow
direction at the point. A smaller vector length of 300 is sufficient to provide
information on the flow in the normal vessel, but insufficient for the obstructed
vessel. Some points of reversed flow have been recorded predominantly in the
left branch, with the overall flow direction towards the right branch, the latter
being consistent with the computer simulated model.
Figure 5.7.1.1. LDA measurements of mean velocity in the normal vessel
(without obstructions). Vector representation and contours with vector
length of 300
The zero point on the x-axis is the middle of the vessel, but the zero point on the
y-axis is few centimetres above the Y=0 of the real model. As no measurements
were carried out very close to the inlet of the vessel, it is not represented in the
LDA visualizations. The drawing of the walls is an approximation, and should be
considered as a guide only. As stated earlier, the measurements were conducted
at the Z=0 plane, cutting through the middle of the model, hence the z-axis is not
presented in these results.
In Figure 5.7.1.1, the contours of the velocity vectors are presented to give an
indication of the flow pattern, strength and direction within the measured planes
of the studied model.
246
Figure 5.7.1.2. Representation of mean velocity vectors only (vector length 300) without the contours in a normal vessel (without obstructions).
Figure 5.7.1.3. Representation of the obstructed vessel with vector length of 300
247
As this software package does not depend on the geometry, it represents points
only relative to their spatial position, even when there are no boundaries or grid
created. The geometry is hand-drawn and is presented for easier interpretation of
the measured results; hence the obstructions are not shown. The shape of the
vessel is relative to the space coordinates and is not identical to the computer
model.
Figure 5.7.1.4. Close up of the obstructed vessel with vector length of 300
It can be seen that the vectors are much smaller for the same vector length as the
one used in Figures 5.7.1.1. and 5.7.1.2. Also, the direction of the vectors is quite
different within the obstructed vessel. Some curving of the vectors also occurs
(see left hand side vector at y=26; the middle one at y=22) due to the multiple
measurements taken at each point (an average of 1000 particle detections at each
point). The flow in the obstructed vessel was harder to detect and measurements
took longer to conduct, showing that obstructions have a definite effect on the
viscous flow.
248
Figure 5.7.1.5. A representation of rake of stream traces in a close up view of
the obstructed vessel with vector length of 300
It can be seen that the vector length of 300 is insufficient to show the vectors,
and a five-fold increase was needed in order to be able to see the direction and
magnitude of the flow. The only problem with the increase in the vector length is
that the scale vector (right hand side, parallel to the x-axis, showing the size of
velocity of 0.01m/s) also increases and in Figure 5.7.1.6, it is outside of the
visual area and cannot be used for scaling purposes. So depending on whether
information on the size or direction of velocity vectors is needed, the lengths will
need to be adjusted.
249
Figure 5.7.1.6. A representation of rake of stream traces in a close up view of
the obstructed vessel with vector length of 1500
5.7.2. Comparison between visualization using different vector lengths
The following Figures represent the measurements in the normal vessel with
different vector lengths of 300 and 500. Figure 5.7.2.1 shows information on the
flow direction, with vector flow of 300. Figure 5.7.2.2., although showing the
overall flow in a clearer way, loses some of the information at points of fast flow,
i.e. at the right branches. A balance is needed between visualising sufficient
information and not losing important information.
250
Figure 5.7.2.1. Normal vessel with vector length of 300 (same as Figure
5.7.1.2.)
The vector length is a ‘post-processing’ tool option, meaning that it has no
bearing on the results, it just makes it easier to visualise the flow direction and
the differences in the velocities around the vessel.
As the pointer might be outside of the area of visualization (as is the case with
the vectors in the right sub-branches in Figure 5.7.2.2. the velocity cannot be
determined), the vector length should not be too high. The faster the flow at a
point, the larger the vector and conversely, the smaller the vector, the slower the
flow. Looking at the different vectors, areas of slow and fast flow can be located.
As each point in this visualisation is an average representation of around one
thousand measured particles, it can be taken as a realistic representation of the
flow in the vessel during the experiments. This means that at each point an
average of one thousand reflecting particles have been registered.
251
Figure 5.7.2.2. Normal vessel with vector length of 500
5.7.3. Comparison between normal and obstructed models
The next two Figures compare the normal and the obstructed vessels using a
vector length of 700 in both cases. Both models with vector lengths of 300 were
presented earlier (Figures 5.7.1.2 and 5.7.1.3). In the Figures below it becomes
clearer that for better representation of the flow in the obstructed model, the
vectors length needs to be substantially larger than the one needed for the normal
vessel.
252
Figure 5.7.3.1 Normal model with vector length of 700
Figure 5.7.3.2. Obstructed model with vector length of 700
253
5.7.4. The most appropriate vector length
The Figures presented in this chapter showed the need for use of different vector
lengths to represent the flow in the two models. A vector length of 300 was
sufficient to provide information on the flow direction and magnitude in the
normal model. For the obstructed model it was found that a vector length of 1500
was more appropriate. One needs to be mindful when increasing the vector
length that the velocity is not the same at all points, and at some points (see the
outer right sub-branch point in the obstructed model in Figures 5.7.4.2. and
5.7.4.3) the vectors can increase disproportionately to the others. This has no
implication for the accuracy of the measurements or the visualisation, and clearly
shows that the flow is not uniform. Thus, learning more about the behaviour of
the flow allows predictions to be made which can possibly lead to improvements
to the flow.
Figure 5.7.4.1. Normal model with vector length of 300
254
Figure 5.7.4.2. Obstructed model with vector length of 1000
Figure 5.7.4.3. Obstructed model with vector length of 1500
255
5.8. Comparison between CFD and LDA models
The next four figures represent velocities at different cross sections and heights
for both the Fluent (Figures 5.8.1.-3.) and LDA measured points (Figure 5.8.4).
Where L and R are written after the mm in the Figures below, they represent L=
left branch and R= right branch. The flow profile for the normal model had its
peak in the middle of the vessel in both the CFD and LDA measurements, but in
the experimental measurements, due to noise, human error and equipment
shortcomings there is some extra disturbance in the flow, which are not present
in the computer generated model. Similar behaviour is observed in the obstructed
model. The interesting part is the flow in the right branch in the obstructed
model, where a clear peak can be seen in both models.
The CFD model does not account for any of the errors due to the nature of the
experimental work carried out with the LDA.
256
Velocity at different heights
0
0.02
0.04
0.06
0.08
0.1
0.12
0.14
-5 -4 -3 -2 -1 0 1 2 3 4 5
Radii mm
Vel
oci
ty m
/s
at y=10mm
at y=20mmat y=32mm at y=60mm
Figure 5.8.1. CFD points from the inlet to the middle of the branching in normal vessel
257
Velocity at different heights
0.06
0
0.01
0.02
0.03
0.04
0.05
-17 -16 -15 -14 -13 -12 -11 -10 -9 -8
Radii mm
Vel
oci
ty m
/s at y=74mmL
Velocity at different heights
0
0.02
0.04
0.06
0.08
0.1
0.12
0.14
8 9 10 11 12 13 14 15 16 17
Radii mm
Vel
oci
ty m
/s
at y=74mmR
Figure 5.8.2. CFD points in the right branch in normal vessel
Figure 5.8.3. CFD points in the left branch in normal vessel
258
Velocity at different heights
0
0.005
0.01
0.015
0.02
0.025
0.03
0.035
-5 -4 -3 -2 -1 0 1 2 3 4 5
Radii mm
Vel
oci
ty m
/s
at y = 10 mmat y = 20 mmat y = 32 mm
Figure 5.8.4. Points measured using LDA in normal vessel
The red line showing y=32mm is at height equal to 72mm in the CFD
simulations, i.e. in the area of branching and shows twice higher velocities in the
right branch compared to the left (velocity of around 0.013 m/s in the left and
0.024 m/s in the right branches). This is consistent with the velocities simulated
in the CFD model at height of 74mm shown on Figure 5.8.2 for the right branch
measuring velocities of 0.12 m/s compared to the left branch shown on Figure
5.8.3 measuring less than 0.06 m/s.
The velocity just below from the branching is represented by the blue line of y
=20mm (add 40 mm to equal the hight in the CFD model) in Figure 5.8.4 and the
shape of the velocity profile fits well with the one represented in Figure 5.4.1 by
the dark blue line at y=60mm.
Figure 5.8.5. Height of vessels used in LDA and CFD simulations
259
Velocity at different heights
0.00E+00
2.00E-02
4.00E-02
6.00E-02
8.00E-02
1.00E-01
1.20E-01
1.40E-01
1.60E-01
-5 -4 -3 -2 -1 0 1 2 3 4 5
Radii mm
Vel
oci
ty m
/s
9.1cm Height
of vessel
0cm
at y = 20 mmat y = 21 mmat y = 22 mmat y = 26 mmat y = 32 mmat y = 37 mmat y = 42.5 mmat y = 48 mm
Figure 5.8.6. Height points in the vessels from the CFD simulations up until just after the trunk obstruction
Velocity at different heights
0
0.02
0.04
0.06
0.08
0.1
0.12
0.14
0.16
0.18
0.2
-7.5 -5.5 -3.5 -1.5 0.5 2.5 4.5 6.5 8.5 10.5 12.5
Radii mm
Vel
oci
ty m
/s
at y = 67 mmLat y = 67 mmRat y = 61 mm
Figure 5.8.7. Height points in the vessels from the CFD simulations at the area of branching (blue line) and just below the obstructions in the branches (red and black lines for right and left branches respectively)
Velocity at different heights
0.16
0
0.02
0.04
0.06
0.08
0.1
0.12
0.14
-20 -15 -10 -5 0 5 10 15 20
Radii mm
Vel
oci
ty m
/s
at y = 74 mmLat y = 74 mmR
Figure 5.8.8. Velocity in the Left and Right branches around obstructions for the CFD simulated model
260
Velocity at different heights
0
0.0002
0.0004
0.0006
0.0008
0.001
0.0012
0.0014
0.0016
0.0018
-5 -4 -3 -2 -1 0 1 2 3 4 5
Radii mm
Vel
oci
ty m
/s at y = 20 mm
at y = 21 mmat y = 22 mmat y = 26 mmat y = 32 mm
Figure 5.8.9. LDA measured points in the obstructed model below the branching (add 20mm to compare with similar heights in the CFD model)
Velocity at different heights
0.008
-0.001
0
0.001
0.002
0.003
0.004
0.005
0.006
0.007
-21 -16 -11 -6 -1 4 9 14 19 24
Radii mm
Vel
ocit
y m
/s
at y = 43 mmat y = 47 mmat y = 51 mm
Figure 5.8.10. LDA measured points in the obstructed model in the area of branching (add 20mm to compare with similar heights in the CFD model) The velocity at height y=20 and y=22 in Figure 5.8.9 is comparable to the
velocity measured at y=20÷32 in Figure 5.8.6., both of which are below the
261
262
obstruction of the vessel. The profile is parabolic and the pattern starts to change
at around y=37mm in the CFD simulations and y=26 for the LDA measurements.
The peak that can be seen in Figure 5.8.10 represented by the purple line is the
same peak as CFD simulated and shown on Figure 5.8.7 by the red line. They
show the logical increase in velocity due to the narrowing cause by the
obstruction in the branch.
There are some abnormal particle measurements due to noise and human error as
the point measured by LDA and shown by the high velocity point on the right
side of the blue line in Figure 5.8.9 at y=21mm. Similar case with the lowest
velocity in the middle of the dark blue line at y=26 of the same Figure 5.8.9.
The model obstructed vessel in the LDA measurements was attached to the
bottom ring inside the bioreactor and so the measurements height has to be
adjusted by 20mm to compare with the CFD simulations. For the normal, non-
obstructed vessel the adjustment needs to be 40mm as the first plane for
measurements was higher than in the obstructed, the later having the trunk
obstruction lower than 40mm above the fixation point for the model vessel to the
bioreactor ring.
The tables with the data on which the above 10 Figures are based are represented
in Appendix 5 of this thesis.
5.9. Conclusions
These simulations confirmed the expectation that obstructions would have an
effect on the blood flow in the portal vein in the situation of diminished blood
supply to the liver due to disease of the organ. The simulations used low flow
velocity, although increased compared to normal portal vein flow, and an
average pressure of 40 cm H2O column. The clinical condition this presentation
is based on is portal hypertension and the impact of obstructions on this
condition was examined. Future studies will need to investigate variables such as
the size and location of the obstructions and their impact on the flow to the liver
263
under the same conditions. The thrombogenic effect of the blood in portal
hypertension with obstructions will also be investigated.
The experimental measurements are in agreement with the computer model
showing higher velocity towards the right branches and different flow in normal
and obstructed models. This computer model can aid medical practitioners in
understanding the blood flow and possibly predicting complications in patients
with portal hypertension with or without the added complications of obstructions.
The input data for this CFD model can be taken from in vivo measurements (e.g.
Duplex Ultrasound, Magnetic Resonance, echo-Doppler and Doppler Duplex
sonography) for individual patients. This model can be adjusted for a variety of
flow parameters and can assist medical practitioners, in conjunction with the
patient-based measurements, to predict the degree of risk to the patient. This type
of model may potentially be used to predict the chances of survival and the risks
of liver failure and mortality in patients with portal hypertension.
Some limitations in the measurements of the physical experiments were
discussed in section 5.7. above and need to be taking in consideration in future
studies.
CHAPTER 6
Conclusions
In this thesis, blood flow through the portal vein of the liver in an idealised
model under portal hypertension conditions was investigated. To achieve this,
research in a variety of areas had to be carried out as outlined below.
All methods currently used or under development to deal with portal
hypertension and/or its complications have their advantages and disadvantages,
and there are examples in which they are more or less effective than other
methods. It will be beneficial and logical for all treatment methods to be
available in any medical centre undertaking hepatic surgery, so that in each
individual case the appropriate approach can be chosen. In some cases a
combination of different treatments might be the correct approach, and the search
for novel methods needs to continue. Part of the aim of the research undertaken
in this thesis was to investigate ways to deal with the diminished blood flow and
look into possible ways to either improve the flow or to predict the impact of the
condition on the outcome for the patient. There has been no attempt to propose
medication or surgical procedures to cure or improve the survival rate in patients
with cirrhosis, as no part of this research had a clinical component.
The practical application of the research undertaken in this thesis will be in the
area of tissue engineering utilising endothelial cell on biodegradable scaffolds.
As the materials used for scaffold fabrication and the manufacturing process are
very complex issues, they have been consequently converted into separate
projects and are under investigation at the time of completing the research done
for this thesis. From independent experiments carried out to investigate possible
scaffold materials and best conditions for cell seeding, it became clear that the
relationship between scaffold porosity and cell adhesion in both steady and
pulsatile flow conditions and the development of more suitable biodegradable
polymers need further investigation and research.
264
The experimental component of this project involved both physical and computer
modelling. The computer model generated for this research was used to enhance
our knowledge on the features of the blood flow through the portal vein. This
model was created before manufacturing the vessel for the physical experiments,
and was used as reference in setting up the LDA measurements. The vessel used
in the physical experiments was fabricated to replicate the computer-generated
portal vein model, and the flow velocity and viscosity, as well as the pressure in
the vessel were a close approximation to the one set for the computer model.
Rheological and hemodynamic characteristics of the blood flow in healthy and
cirrhotic patients and the impact of portal hypertension on these characteristics
was combined with general theories and studies of the blood flow in the human
cardiovascular system. Comparison between assumed Newtonian behaviour of
the flow and the non-Newtonian realistic representation of blood flow were
given.
The physical model closely replicated the computer created model, and the
experiments were done mainly to verify the validity of the computer simulations.
Apart from velocity, no other measurements were carried in the LDA phase of
this work. The vessel used in those measurements was fabricated based on the
size and shape of the computer model, and the fluid used had the same viscosity
as the one set in the FLUENT model. The same velocity was achieved and
measurements for the normal and the obstructed models were carried out. Some
problems were identified during the physical experiments as either sources of
error or inaccurate representation of the physiological or in vitro tissue
engineering conditions and these are discussed below. The physical experiment
was reasonably easy to perform and repeat, and the cost of all components
(excluding the Laser itself, but including the pulsatile pump) is under AUD 1000.
If the same measurements need to be performed for individual patient portal vein
geometry, only the vessel model would need to be replaced and some
modifications to the bioreactor carried out, but the cost of those would not
exceed AUD 200.
For the purposes of both the physical measurements and the tissue culturing
components of this work (the latter being the basis for future research) a novel
design bioreactor was developed.
265
The simplified design of the bioreactor has a small number of parts allowing for
easy assembling of the device by non-experts – total time for assembling and
disassembling can be less than 10 minutes (including the tubing and reservoirs).
The materials used (glass, medical grade stainless steel and silicone tubing) can
be sterilised by commonly used methods, including autoclaving, thus providing
safe and easy accessible work conditions for tissue engineering. One of the main
problems in tissue engineering is the sterility of the materials used and this
concern has been addressed as part of the bioreactor design. The next
requirement is for compatibility so the device can be incubated, which has also
been taken into account with this design, as it allows for the bioreactor together
with all tubing and reservoirs to fit into commercially available incubators.
The device can be easily modified to accommodate vessels of different size,
diameter and branching geometry, by simply exchanging the top cylinder to
which the scaffold is attached.
The maximum cost to manufacture the bioreactor, including labour costs, is
AUD 500. This cost will be lower if a larger quantity is manufactured. The
bioreactor can be easily decontaminated, washed and re-used as required.
The design allows for control of the flow conditions and media mixture
(including media exchange). There are no moving parts within the bioreactor,
hence there is no need for externally generated momentum to keep the system
working.
The durability of the medical grade stainless steel parts have been tested by being
left in a glycerol solution for over two weeks without any indication of rusting.
Nevertheless, there are several areas, as outlined in the next section on future
work, which need to be investigated further and could lead to improved
performance of the apparatus.
The computer modelling and simulations presented in Chapter 5 confirmed the
expectation that obstructions would have an effect on the blood flow in the portal
vein in the situation of diminished blood supply to the liver due to disease of the
organ. The clinical condition this presentation is based on is portal hypertension
and the impact of obstructions on this condition was examined.
266
The visualisation opportunities using FLUENT software were given for
illustration purposes and are not restrictive to the visualisation of other
parameters. Those include such flow parameters as velocity vectors and
magnitude, static pressure, wall shear stress and stain rate, as well as
combinations of these parameters in different parts of the vessel and histograms
of velocity distribution or a visual representation of the static pressure for each
point inside the vessel.
Comparisons of the flow between the obstructed and normal vessels showed
significant differences due to the obstructions, as both simulations were
performed using the same flow parameters and operational and boundary
conditions. These differences were also observed when the same parameters
were used under Newtonian or non-Newtonian conditions or when different flow
velocities were simulated.
In the simulations where the fluid was set as a multiphase flow, comprising of
plasma, erythrocytes, leukocytes and platelets, a very distinct pattern between the
flows of the phases was seen. This visualisation can be very useful in individual
cases, as it shows that the blood will have a different behaviour depending on the
percentage and size of the various blood components. Thus, in patients with low
erythrocyte counts the flow will behave differently than in patients with normal
or high counts of those particles. This can possibly help explain the flow
disturbance in patients with blood disorders.
The experimental measurements were in agreement with the computer model
showing higher velocity towards the right branches and different flow in normal
and obstructed models. This model can aid medical practitioners in
understanding the blood flow and possibly predicting complications in patients
with portal hypertension with or without the added complications of obstructions.
The input data for this CFD model can be taken from in vivo measurements for
each individual patient. This model can be adjusted for a variety of flow
parameters and could assist medical practitioners, in conjunction with the
patient-based measurements, to predict the degree of risk to the patient. This type
of model, after clinical validation, may potentially be used to predict the chances
of survival and the risks of liver failure and mortality in patients with portal
hypertension.
267
Future work The research carried out in this thesis revealed some areas for future work and
improvements, all of which can assist in either deepening our understanding or
discovering new ways to solve specific problems. Below are some of the most
significant areas for future work identified in this thesis.
For the purpose of maintaining the viscosity of the fluid in the bioreactor during
LDA flow measurements, heating options need to be investigated for either the
viscous fluid or the water in the tank. Heating plates around the tank, a separate
heating chamber for the blood-like solution or alternative methods need to be
explored to resolve this problem. For the purpose of tissue engineering, the
growth medium can be heated by covering the silicone tubes with a warming
jacket, or simply placing the whole system into an incubator (if there is no need
of constant monitoring). The bioreactor and its tubing and reservoirs developed
in this work can be safely placed inside commercially available incubators.
In the present study the model blood vessel was made of glass to allow for a high
degree of transparency, but the limitations of this material are its stiffness and
lack of flexibility, both of which do not correctly represent the native vessel or
the ideal scaffold. However, the same properties were assigned to the computer
model to allow for verification of the results. Alternative materials, with similar
transparency but also with a reasonable degree of flexibility, need to be
developed for LDA measurements. It is worth remembering that transparency is
not an issue for biodegradable scaffolds used for tissue engineering, as the cells
growing on the material will not permit LDA measurements.
In this thesis the changing behaviour of blood has not been taken into account,
but this needs to be considered for future research. It is important to individualise
each model to represent the specific changes occurring in patients’ blood flow so
realistic modelling can be done. A better understanding of the blood flow pattern
using modelling and visualization may help to minimize the thrombogenesis of
artificial blood vessels and organs. The importance of the deformation
characteristics of blood cells on the flow has been demonstrated in many studies
268
over the last three decades and can be implemented in the model developed in
our research if required.
To improve the bioreactor design, tests with different angles of the inlets (if
shunts are considered) need to be carried out to determine the best one from a
hemodynamic point of view. Different joining angles between the lower and
perfusion chambers of the bioreactor and their effect on the flow need to be
studied to determine the most suitable one. Creating a screw lid for the
inlet/outlet of the bioreactor’s lid will allow for better sealing. Silicone sealant
can be used on the outer surface of the connection between the body of the
bioreactor and the lid to prevent leaks after the system has been sterilized and
assembled.
A pulsatile pump can be introduced to the system while the bioreactor is inside
the incubator using the linking outlets most incubators have (so the pump stays
outside the incubator). Alternatively, a new pulsatile pump needs to be designed,
which can operate safely in an environment that is warmer and more humid than
room temperature.
Computational Fluid Dynamics studies using the models developed in this thesis
should include an investigation of variables such as the size and location of the
obstructions and their impact on the flow to the liver under the same conditions.
The thrombogenic effect of the blood in portal hypertension with obstructions
will also need to be investigated. Multiphase flow clinical studies of the impact
of the concentration of blood cell types on the behaviour of the flow and whether
they have consequences to the patients’ survival should also be considered.
Other areas that could be investigated to better improve the utility of the
bioreactor in tissue engineering applications include:
• The development of biodegradable porous materials, which do not release
any toxic particles when degrading in tissue culturing conditions and are
strong enough and suitable for grafting of blood vessels.
269
• Testing of different cell seeding methods to find the appropriate ones for
the chosen scaffolding material.
As can be seen from the list above, there are various areas in which further
research can be done following the work conducted in this thesis. Each topic
requires specialists in that area to conduct the investigations and as all require
extensive research that is beyond the scope of the current study.
Finally, due to the wide variation in the structure, function and pathology of
patients, portal vein and blood flow, each computer model needs to take into
account the clinically measured parameters (derived from imaging), the patient
disease history, and other conditions the individual might suffer from (i.e. kidney
failure, elevated blood pressure, etc.). Computer modelling therefore has to be
carried out on a case-by-case basis with regards to each patient. Until a system is
developed that allows rapid modelling (within minutes) while still being based
on the individual patient, computer models cannot be used for emergency
procedures. With the development of new software, the possibility of changing
the model by simply changing the input parameters may arise. So far, and from
our experience, the simulation requires re-drawing of the grid and mesh for each
case, which is in essence the most time-consuming part of the modelling process.
270
References Absolom, D R, Hawthorn, L A & Chang, G 1988, 'Endothelialization of polymer
surfaces', J Biomed Mater Res, vol. 22, no. 4, Apr, pp. 271-285. Aderriotis, D & Sandor, G K B 1999, 'Outcomes of Irradiated Polyglactin 910
Vicryl Rapide Fast-Absorbing Suture in Oral and Scalp Wounds', Journal de l’Associacion dentaire canadienne, vol. 65, no. 6, June.
Agarwala, S, Dutta, H, Bhatnagar, V, Gulathi, M, Paul, S & Mitra, D 2000, 'Congenital hepatoportal arteriovenous fistula: report of a case', Surg Today, vol. 30, no. 3, pp. 268-271.
Alanazi, A, Nojiri, C, Kido, T, Noguchi, T, Ohgoe, Y, Matsuda, T, Hirakuri, K, Funakubo, A, Sakai, K & Fukui, Y 2000, 'Engineering Analysis of Diamond-Like Carbon Coated Polymeric Materials for Biomedical Applications', Artificial Organs, vol. 24, no. 8, pp. 624-627.
Almoudarres, M, Vega, K J & Trotman, B W 1998, 'Noncirrhotic portal hypertension in the adult: case report and review of the literature', J Assoc Acad Minor Phys, vol. 9, no. 3, pp. 53-55.
American Liver Foundation 2004, Shunt Surgical Procedures – Liver Health Information viewed March 2004, www.liverfoundation.org/egi-bia/dbs/articles.egi?db-article&uid.
Amesur, N B, Zajko, A B, Orons, P D, Sammon, J K & Casavilla, F A 1999, 'Transjugular intrahepatic portosystemic shunt in patients who have undergone liver transplantation', Journal of Vascular and Interventional Radiology, vol. 10, no. 5, May 1999, pp. 569-73.
Amitrano, L, Brancaccio, V, Guardascione, M A, Margaglione, M, Iannaccone, L, D'Andrea, G, Riccardo, M R, Ames, P R & Balzano, A 2000, 'Inherited Coagulation Disorders in Cirrhotic Patient With Portal Vein Thrombosis', Hepatology, vol. 31, no. 2, Feb 2000, pp. 345-348.
Amitrano, L, Brancaccio, V, Guardascione, M A, Margaglione, M, Sacco, M, Martino, R, De Nucci, C, Mosca, S, Iannaccone, L, Ames, P R, Romano, L & Balzano, A 2002, 'Portal vein thrombosis after variceal endoscopic sclerotherapy in cirrhotic patients: role of genetic thrombophilia', Endoscopy, vol. 34, no. 7, July 2002, pp. 535-8.
Anderson, C B, Sicard, G A & Etheredge, E E 1980, 'Bovine carotid artery and expanded polytetrafluoroethylene grafts for hemodialysis vascular access', Journal of Surgical Research, vol. 29, pp. 184-188.
Anthony, P P, Ishak, K G, Nayak, N C, Poulsen, H E, Scheuer, P J & Sobin, L H 1978, 'The morphology of cirrhosis. Recommendations on definition, nomenclature, and classification by a working group sponsored by the World Health Organization', Journal of Clinical Pathology, vol. 31, pp. 395-414.
Appel, H, Loddenkemper, C, Schirmacher, P, Dienes, H P, Sieper, J, Rudwaleit, M, Golder, W & Zeitz, M 2003, 'Congenital absence of the portal vein with splenomegaly and hypersplenism in a young woman', Digestion. Basel, vol. 67, no. 1/2, p. 105.
271
Arey, L B 1957, Human Histology. A Textbook in Outline Form, Forth Edition edn., WB Saunders Company
Armstrong, J, Narracott, A J, Milton, R, Galea, J, Cooper, G J, Lawford, P V, Hose, D R, Cumberland, D C & Holt, C M 2000, 'Development of an ex vivo model to investigate the effects of altered haemodynamics on human bypass grafts', Journal of Medical Engineering & Technology, vol. 24, no. 5, Sept-Oct 2000, pp. 183-191.
Aroesty, J & Gross, J F 1972, 'The mathematics of pulsatile flow in small vessels I. Casson theory', Microvascular Research, vol. 4, no. 1, Jan 1972, pp. 1-12.
Asakura, T, Ohkohchi, N, Orii, T, Koyamada, N, Tsukamoto, S, Sato, M, Enomoto, Y, Usuda, M & Satomi, S 2003, 'Portal vein pressure is the key for successful liver transplantation of an extremely small graft in the pig model', Transpl Int, vol. 16, pp. 376-382.
Atri, M, Bret, P M & Fraser-Hill, M A 1992, 'Intrahepatic portal venous variations: prevalence with US', Radiology, vol. 184, pp. 157-158.
Augst, A D, Barratt, D C, Glor, F P, McG Thom, S A & Xu, X Y 2003, 'Accuracy and Reproducibility of CFD Predicted Wall Shear Stress Using 3D Ultrasound Images', Journal of Biomechanical Engineering, vol. 125, no. 2, April 2003, pp. 218-222.
Bailly, A L 1999, 'Thrombosis of angiographic catheters in humans: experimental study', Int J Artif Organs, vol. 22, pp. 690-700.
Bakathavasalam, R, Marsh, C L, Perkins, J D, Levy, A E, Healey, P J & Kuhr, C S 2001, 'Rescue of acute portal vein thrombosis after liver transplantation using a cavoportal shunt at re-transplantation', Am J Transplant, vol. 1, no. 3, Sept 2001, pp. 284-7.
Balaz, S 2000, Drug Distribution: Elimination III, Ballermann, B J & Ott, M J 1995, 'Adhesion and differentiation of endothelial cells
by exposure to chronic shear stress: a vascular graft model', Blood Purif, vol. 13, no. 3-4, pp. 125-134.
Barakat, M 2002, 'Portal vein pulsatility and spectral width changes in patients with portal hypertension: relation to the severity of liver disease', The British Journal of Radiology, vol. 75, pp. 417-421.
Barakat, M 2003, 'Non pulsatile hepatic and portal vein waveforms in patients with liver cirrhosis: Concordant and disconcordant relationships', Ultrasound in Medicine and Biology, vol., no. 29, p. 5S.
Bardell, R L, Sharma, N R, Foster, F K, Afromowitz, M A & Penney, R J 1997, 'Designing high-performance micro-pumps based on no-moving-parts valves' MEMS ASME 1997, 47-53.
Barkat, M 2004, 'Non-pulsatile hepatic and portal vein waveforms in patients with liver cirrhosis: concordant and discordant relationships', British Journal of Radiology, vol. 77, no. 919, July 2004, pp. 547-551.
Barkell, H, Tripathi, D, Ferguson, J W, Ireland, H, Redhead, D N & Hayes, P C 2004, 'A single centre experience of PTFE covered stents: a TIPSS revolution', Gut, vol. 53, no. 4, Apr 2004, p. A91.
272
Bate, H 1977, 'Blood viscosity at different shear rates in capillary tubes', Biorheology, vol. 14, no. 5-6, pp. 267-75.
Becker, Y & Reed, G 1996, 'The role of elective operation in the treatment of portal hypertension', American Surgeon, vol. 62, no. 3, March, pp. 313-48.
Belanger, M-C, Marois, Y, Roy, R, Mehri, Y, Wagner, E, Zhang, Z, King, M W, Yang, M, Hahn, C & Guidoin, R 2000, 'Selection of Polyurethane Membrane for the Manufactire of Ventricles for a Totally Implantable Artificial Heart: Blood Compatibility and Biocompatibility Studies', Artificial Organs, vol. 24, no. 11, pp. 879-888.
Bhat, V D, Klitzman, B, Koger, K, Truskey, G A & Reichert, W M 1998, 'Improving endothelial cell adhesion to vascular graft surfaces: clinical need and strategies', J Biomater Sci Polym Ed, vol. 9, no. 11, pp. 1117-35.
Bolognese, M, Sacerdoti, D, Bomdonato, G, Merkel, C, Sartori, G, Merenda, R, Nava, V, Angeli, P, Feltracco, P & Gatta, A 2002, 'Change in portal flow after liver transplantation: effect on hepatic arterial resistance indices and role of spleen size', Hepatology, vol. 35, no. 3, March 2002, pp. 601-608.
Bolognesi, M, Sacerdoti, D, Merkel, C, Bombonato, G & Gatta, A 2001, 'Noninvasive grading of the severity of portal hypertension in cirrhotic patients by echo-color-Doppler', Ultrasound in Medicine and Biology, vol. 27, no. 7, July 2001, pp. 901-907.
Bolondi, L, Bassi, S L, Gaiani, S & Barbara, L 1990, 'Doppler flowmetry in portal hypertension', Journal of Gastroenterology and Hepatology, vol. 5, pp. 459-467.
Bonert, M, Leask, R L, Butany, J, Etheir, C R, Myers, J G, Johnston, K W & Ojha, M 2003, in BioMedical Engineering OnLine, www.biomedical-engineering-online.com/content/2/1/18, Vol. Nov 2003.
Bonner, R & Nossal, R 1981, 'Model for laser Doppler measurements of blood flow in tissue', Applied Optics, vol. 20, no. 12, 15 June 1981, pp. 2097-2107.
Bordenave, L, Remy-Zolghadri, M, Fernandez, P, Bareille, R & Midy, D 1999, 'Clinical performance of vascular grafts lined with endothelial cells', Endothelium, vol. 6, no. 4, pp. 267-275.
Bos, G W, Scharenborg, N M, Poot, A A, Engbers, G H, Beugeling, T, van Aken, W G & Feijen, J 1999, 'Blood compatibility of surfaces with immobilized albumin-heparin conjugate and effect of endothelial cell seeding on platelet adhesion', J Biomed Metar Res, vol. 47, no. 3, 5 Dec 1999, pp. 279-91.
Bos, G W, Scharenborg, N M, Poot, A A, Engbers, G H, Terlingen, J G, Beugeling, T, Van Aken, W G & Feijen, J 1998, 'Adherence and proliferation of endothelial cells on surface-immobilized albumin-heparin conjugate', Tissue Eng, vol. 4, no. 3, Fall, pp. 267-79.
Bosch, J & Garsia-Pagan, J C 2000, 'Complications of cirrhosis. I. Portal Hypertension', J Hepatol, vol. 32(1Suppl), pp. 141-56.
273
Botchwey, E A, Pollack, S R, Levine, E M & Laurencin, C T 2001, 'Bone tissue engineering in a rotating bioreactor using a microcarrier matrix system', Inc. J Biomed Mater Res, vol. 55, pp. 242-253.
Bowlin, G L & Rittgers, S E 1997, in Virtual Medical World – monthly, Brussels, Nov-DecBrossels.
Boyer, T D 2003, 'Transjugular intrahepatic portosystemic shunt: current status', Gastroenterology vol. 124, no. 6, May, pp. 1700-10.
BritishLiverTrust British Liver Trust, http://www.britishlivertrust.org.uk/home.aspx. Broering, D C, Hillert, C, Krupski, G, Fischer, L, Mueller, L, Achilles, E G,
Schulte, J, Esch, A & Rogiers, X 2002, 'Portal Vein Embolization vs. Portal Vein Ligation for Induction of Hypertrophy of the Future Liver Remnant', Journal of Gastrointestinal Surgery, vol. 6, no. 6, Nov-Dec 002, pp. 905-913.
Brown, R S J 1997, 'Transjugular intrahepatic portosystemic shunt as a form of treatment for portal hypertension: indications and contradictions', Adv Intern Med, vol. 42, pp. 485-504.
Bruden, J L, Hsieh, T-c, Lerea, K M, Olson, S C & Wu, J M 2001, 'Induced cytoskeletal changes in bovine pulmonary artery endothelial cells by resveratrol and the accompanying modified responses to arterial shear stress', BMC Cell Biology, vol. 2, no. 1, p. 1.
Brunn, P O & Vorwerk, J 1993, 'Determination of the steady-state shear viscosity from measurements of the apparent viscosity for some common types of viscosity', Rheologica Acta, vol. 32, pp. 380-198.
Buell, J F, Funaki, B, Cronin, D C, Yoshida, A, Perlman, M K, Lorenz, J, Kelly, S, Brady, L, Leef, J A & Millis, J M 2002, 'Long-term venous complications after full-size and segmental pediatric liver transplantation', Ann Surg, vol. 236, no. 5, Nov 2002, pp. 658-66.
Burg, K J, Culberson, C R, Greene, K G, Loebsack, A B, Roland, W R, Holder, W D J, Halberstadt, C R & Beiler, R J 1999, 'Absorbable Mesh Aids in Development of Discrete, Tissue-Engineered Constructs' 18-th Annual Southern Biomedical Engineering Conference, Crit Rev Biomed Eng.2000 383-387.
Burns, P N & Jaffe, C C 1985, 'Quantitative Flow Measurements with Doppler Ultrasound: Techniques, Accuracy, and Limitations' Symposium on Advances in Cardiac Imaging, Radiologic Clinic of North America Dec 1985,
Caps, M T, Helton, W S & Johansen, K 1996, 'Left-upper-quadrant devascularization for ‘unshuntable’ portal hypertension', Arch Surg, vol. 131, Aug 1996, pp. 834-838.
Capussotti, L, Vergara, V, Polastri, R, Bouzari, H & Galatola, G 2000, 'Liver function and encephalopathy after partial vs direct side-to-side portacaval shunt: a prospective randomized clinical trial', Surgery, vol. 127, no. 6, June 2000, pp. 614-621.
Cardoso, J E, Giroux, L, Kassissia, I, Houssin, D, Habib, N & Huet, P M 1994, 'Liver function improvement following increased portal blood flow in cirrhotic rats', Gastroenterology, vol. 107, pp. 460-467.
274
Carlisle, E S, Muthumarthanda, R M, Nelson, K D, Thomes, B E, Timmons, R B, Constantinescu, A, Eberhart, R C & Bankey, P E 2000, 'Enhancing Hepatocyte Adhesion by Pulsed Plasma Deposition and Polyethylene Glycol Coupling', Tissue engineering, vol. 6, no. 1, pp. 45-52.
Carter, J H, Welch, C S & Barron, R E 1961, 'Changes in the hepatic blood vessels in cirrhosis of the liver', Surgery Gynecology & Obstetrics, vol. 113, no. Number 2, Aug, pp. 133-137.
Chaib, E, Ribeiro, M A F J, Saad, W A & Gama-Rodrigues, J 2005, 'The main hepatic anatomic variations for the purpose of split-liver transplantation', Transplantation Proceedings, vol. 37, no. 2, March, pp. 1036-1066.
Chamuleau, R 2002, 'Bioartificial liver support anno 2001', Metabolic Brain Disease, vol. 17, no. 4, Dec, pp. 485-491.
Chamuleau, R A F M 2003, 'Artificial Liver Support in the Third Millennium', Artificial Cells, Blood Substitutes, and Biotechnology, vol. 31, no. 2, pp. 117-126.
Chamutal, G, Yaron, I & Oren, S 2004, 'Hepatic hydrothorax – pathophysiology, diagnosis and treatment – review of the literature', Liver, vol. 24, no. 4, Aug 2004, pp. 281-285.
Charara, J, Ruel, J, Lemay, J & Doillon, C J 1999, 'Development of a flow simulator to study haemodynamic behaviour of natural and artificial blood vessels under physiologic flow conditions', Journal of Medical Engineering & Technology, vol. 23, no. 3, May-June 1999, pp. 83-95.
Chatila, R, Ferayorni, L, Gupta, T & Groszmann, R J 2000, 'Local arterial vasoconstriction induced by octreotide in patients with cirrhosis', Hepatology, vol. 31, no. 3, March, pp. 572-6.
Chawla, Y, Santa, N, Dhiman, R K & Dilawari, J B 1998, 'Portal hemodynamics by duplex Doppler sonography in different grades of cirrhosis', DIGESTIVE DISEASES AND SCIENCES, vol. 43, no. 2, Feb 1998, pp. 354-357.
Chen, R Y Z, Carlin, R D, Simchon, S, Jan, K M & Chien, S 1989, 'Effects of dextran-induced hyperviscosity on regional blood flow and hemodynamics in dogs', Am J Physiol Heart Circ Physiol, vol. 256, pp. H898-H905.
Chen, S C, Mullon, C, Kahaku, E, Watanabe, F, Hewitt, W, Eguchi, S, Middleton, Y, Arkadopoulos, N, Rozga, J, Solomon, B & Demetriou, A A 1997, 'Treatment of severe liver failure with a bioartificial liver', Ann N Y Acad Sci, vol. 831, pp. 350-360.
Chen, V T, Wei, J & Liu, Y C 1992, 'A new procedure for management of extrahepatic portal obstruction. Proximal splenic-left intrahepatic portal shunt', Arch Surg, vol. 127, Nov 1992, pp. 1358-1360.
Chervu, A & Moore, W S 1990, 'An overview of intimal hyperplasia', Surg Gynecol Obstet, vol. 171, no. 5, Nov 1900, pp. 433-47.
Cheynel, N, Arnal, E, Rat, P, Benoit, L, Tronilloud, P & Favre, J-P 2001, 'Absence of the portal bifurcation at the hulim of the liver due to intrahepatic origin of the left branch of the portal vein', Surgical and Radiological Anatomy, vol. 23, no. 5, pp. 355-357.
275
Child, C G, Barr, D, Holswade, G R & Harrison, C S 1953, 'Liver Regeneration Following Portacaval Transposition in Dogs', Annals of Surgery, vol. 139, no. 4, Oct.
Children’sLiverDiseaseFoundation. Chojkier, M & Conn, H O 1980, 'Esophageal tamponade in the treatment of
bleeding varices. A decadel progress report', Dig Dis Sci, vol. 25, no. 4, April 1980, pp. 267-272.
Clark, R E, Goldstein, A H, Pacella, J J, Walters, R A, Moeller, F W, Cattivera, G R, Davis, S & Magovern, G J 1996, 'Small, low-cost implantable centrifugal pump for short-term circulatory assistance', Ann Thorac Surg, vol. 61, no. 1, Jan 1996, pp. 452-460.
Coherent Scientific & Dantec Dynamics www.dantecdynamics.com/LDA/Princip/Index.html viewed 2002, . www.dantecdynamics.com/LDA/Princip/Index.html.
Collart, F, Caus, T, Pomane, C, Pellissier, V, Chi, Y, Monties, J R & Mesana, T 2000, 'Clinical Evaluation of Heparin-Coated Circuits for Routine Coronary Artery Bypass Grafting Surgery: A Prospective Randomized Study', Artificial Organs, vol. 24, no. 8, pp. 611-613.
Collins, J C 1998, 'Partial Portacaval Shunt for Variceal Hemorrhage', The Journal of the American Medical Association, vol. 280, no. 8, 26 Aug1998, p. 678j(1).
Collins, J C, Rypins, E B & Sarfeh, I J 1994, 'Narrow-diameter portacaval shunt for management of variceal bleeding', World Journal of Surgery, vol. 18, no. 2, Mar-Apr, pp. 211-5.
Collins, J C & Sarfeh, I J 1995, 'Surgical management of portal hypertension', The Western Journal of Medicine, vol. 162, no. 6, June 1995, pp. 527-535.
Collins, J C & Sarfeh, I J 1998, 'Percutaneous or surgical shunts for varices caused by portal hypertension', The Western Journal of Medicine, vol. 168, no. 1, Jan, pp. 34-36.
Cooney, D O 1976, Biomedical Engineering and Instruments. Couinaud, C 1993, 'Absence of the portal bifurcation', Journal de Chirurgie, vol.
130, no. 3, pp. 111-115. Cwikiel, W 2006, 'Interventional Procedures Involving Portal Vein Circulation: A
Review ', Acta Radiologica, vol. 47, pp. 145-156. Dagenais, M, Pomier, L G, Dufresne, M P, Roy, A, Fenyves, D & Lapointe, R
1994, 'Experience with the transjugular intrahepatic portacaval shunt', Ann Chir, vol. 48, no. 8, pp. 671-8.
Dalinghaus, M, Knoester, H, Gratama, J W, Van der Meer, J, Zijlstra, W G & Kuipers, J R 1994, 'Effect of increased whole blood viscosity on regional blood flows in chronically hypoxemic lambs', Am J Physiol Heart Circ Physiol, vol. 267, pp. H471-H476.
Damen, B S 2003, Design, Delevopment and Optimisation of a tissue culture vessel system for tissue engineering applications, thesis, Industrial Research Institute Swinburne, Swinburne University of Technology, Melbourne.
276
Dark, G 2003, http://cancerweb.ncl.ac.uk/cgi-bin/omd?adhesion. Darnault, T P, Bretagne, J F, Fournier, V & Raoul, J 1989, 'Splenic
haemodynamic assessment in patients with liver cirrhosis and hypersplenism', Gastroenterology, vol. 96, p. A589.
De Jonge, J, Zondervan, P E, Kooi, P P, Izermans, J N, Metselaar, H J & Tilanus, H W 2003, 'Portal Flow Diversion is Essential for Graft Survival in Canine Auxiliary Partial Orthotopic Liver Transplantation', European Surgical Research, vol. 35, pp. 14-21.
De Ville, J d G, Alberti, D, Falchetti, D, Rigamonti, W, Matricardi, L, Clapuyt, P, Sokal, E M, Otte, J B & Caccia, G 1999, 'Treatment of Extrahepatic Portal Hypertension in Children by Mesenteric-to-left Portal Vein Bypass: a New Physiological Procedure', Eur J Surg, vol. 165, pp. 777-781.
Debski, R, Borovetz, H, Haubold, A & Hardesty, R 1982, 'Polytetrafluoroethylene grafts coated with ULTI carbon', Trans Am Soc Artif Intern Organs, vol. 28, pp. 456-8.
Del Rizzo, D F, Moon, M C, Werner, J P & Zahradka, P 2001, 'A Novel Organ Culture Method to Study Intimal Hyperplasia at the Site of a Coronary Artery bypass Anastomosis', Ann Thoracic Surg, vol. 71, pp. 1273-80.
Denk, H 2004, 'Pathology of portal hypertension', Journal of Gastroenterology and Hepatology, vol. 19, pp. S346-S348.
Deplano, A, Migaleddu, V, Pischedda, A, Garrucciu, G, Gregu, G, Multinu, C, Piredda, M, Tocco, A, Urigo, F, Cossu, P A & Solinas, A 1999, 'Portohepatic gradient and portal hemodynamics in patients with cirrhosis due to hepatitis C virus infection', Digestive Diseases and Sciences, vol. 44, no. 1, Jan, pp. 155-162.
Dept of Mechanical Engineering and Dept of Electrical Engineering, U o W, Seattle, U.S.A. 2000-ASME; 2000-uTAS; 1999-ASME; 1998-uTAS; 1997-ASME; 1995-ASME, 'Micro-Pumps With No-Moving-Part Valves' ASME,
Dhiman, R K, Chawla, Y, Vasishta, R K, Kakkar, N, Dilawary, J B, Trehan, M S, Puri, P, Mitra, S K & Suri, S 2002, 'Non-cirrhotic portal fibrosis (idiopathic portal hypertension): Experience with 151 patients and a review of the literature', Journal of Gastroenterology and Hepatology, vol. 17, pp. 6-16.
Dinnar, U & Raton, B 1981, Cardiovascular Fluid Dynamics. Dirchfield, M R, Gibson, R N, Donlan, J D & Gibson, P R 1992, 'Duplex Doppler
Ultrasound Signs of Portal Hypertension: Relative Diagnostic Value of Examination of Paraumbilical Vein, Portal Vein and Spleen', Australas Radiol, vol. 36, pp. 102-105.
Drelich-Zbroja, A, Jargiello, T, Krzyzanowski, W, Szymanska, A & Szczerbo-Trojanowska, M 2003, 'The diagnostic value of Levovist in Doppler imaging of renal arteries in patients with renovascular hypertension before and after angioplasty', Ultrasound in Medicine & Biology, vol. 29, Supplement 1, no. 5, p. S61.
Dumont, K, Stijnen, J M A, Vierendeels, J, Van de Vosse, F N & Verdonck, P R 2004, 'Validation of a Fluid-Structure Interaction Model of a Heart Valve using the Dynamic Mesh Method in Fluent', Computer Methods in
277
Biomechanics and Biomedical Engineering, vol. 7, no. 3, June 2004, pp. 139-146.
Dumont, K, Yperman, J, Verbeken, E, Segers, P, Meuris, B, Vandenberghe, S, Flameng, W & Verdonck, P 2002, 'Design of a New Pulastile Bioreactor for Tissue Engineered Aortic Heart Valve Formation', Artificial Organs, vol. 26, no. 8, August, pp. 710-714.
Dumont, K, Yperman, J, Verbeken, E, Segers, P, Meuris, B, Vandenberghe, S, Flameng, W & Verdonck, P R 2001, 'A new pulsatile bioreactor for tissue engineered heart valve formation' XXVI Congress de la Societe de Biomecanique, Marseille, France, 13-14 Sept, 74.
Dunkern, T R, Paulitschke, M, Meyer, R, Buttemeyer, R, Hetzer, R, Burmester, G & Sittinger, M 1999, 'A novel perfusion system for endothelialization of PTFE grafts under defined flow', Eur J Vasc Endovasc Surg, vol. 18, no. 2, Aug, pp. 105-110.
Durst, F, Melling, A & Whitelaw, J H 1976, Principles and Practice of laser-Doppler-anemometry, London: Academic Press, , London.
Easthrope, P L 1980, 'A comparison rheological constitutive functions for whole human blood', Biorheology, vol. 17, p. 235.
Ederer, I, Raetsch, P, Schullerus, W, Tille, C & Zech, U 1990-2000 multiple, Piezoelectric Micro-pump (FhG-IFT), viewed
Eguchi, A, Hashizume, M, Kitano, S, Tanoue, K, Wada, H & Sugimachi, K 1991, 'High rate of portal thrombosis after splenectomy in patients with esophageal varices and idiopathic portal hypertension', Arch Surg, vol. 126, June 1991, pp. 752-755.
Eguchi, S 1986, 'Portal systemic collateral flow patterns demonstrates by splenoportogram in portal hypertension and their clinical significance', Acta Hepatol Japonica, vol. 27, pp. 1589-1597.
Ellingsen, J E & Lyngstadaas, S P 2003, Bio-Implant Interface. Improving Biomaterials and Tissue Reactions, CRC PressLLC.
Encyclopaedia Britannica 2005. Ersoz, G & Ersoz, S 2003, 'Changes in portal blood flow following acute exercise
in liver transplant recipients', Transplant Proc, vol. 35, no. 4, June 2003, pp. 1456-7.
Esquivel, C O & Blaisdell, F W 1986, 'Why small caliber vascular grafts fail: a review of clinical and experimental experience and the significance of the interaction of blood at the interface', J Surg Res, vol. 41, no. 1, July 1986, pp. 1-15.
Finol, E A & Amon, C H 2002, 'Flow-induced Wall Shear Stress in Abdominal Aortic Aneurysms: Part I – Steady Flow Hemodynamics', Computer Methods in Biomechanics and Biomedical Engineering, vol. 5, no. 4, pp. 309-318.
Fischer, G, Rak, R & Sackmann, M 1998, 'Improved investigation of portal-hepatic veins by echo-enhanced Doppler sonography', Ultrasound in Medicine & Biology, vol. 24, no. 9, Dec 1998, pp. 1345-1349.
Fisher, J-G, Mewes, H, Hopp, H-H & Schubert, R 1996, 'Analysis of pressurized resistance vessel diameter changes with a low cost digital image
278
processing device', Computer Methods and Programs in Biomedicine, vol. 50, no. 1, June 1996, pp. 23-30.
Flemming, J S, Ackery, D M, Walmsley, B H & Karran, S J 1983, 'Scintigraphic estimation of arterial and portal blood supplies to the liver', J Nucl Med, vol. 24, pp. 1108-13.
FLUENT 2006, www.fluent.com. Fluent Europe Ltd 2002, 'Application Review of Computational Fluid Dynamics
for the Medical and Pharmaceutical Industry', Business Briefing: Medical device manufacturing & technology 2002,
Foster, F K, Bardell, R L, Afromowitz, M A, Sharma, N R & Blanchard, A 1995, 'Design, fabrication and testing of fixed-valve micro-pumps' 1995 IMECE & ASME 1995, Division, P o t A F E, 39-44
Fung, Y C 1993, Biomechanics.Mechanical Properties of Living Tissues, Second Edition edn., Springer-Verlag, New York.
Gaiani, S, Bolondi, L, Li, B S, Zironi, G, Siringo, S & Barbara, L 1991, 'Prevalence of spontaneous hepatofugal portal flow in liver cirrhosis. Clinical and endoscopic correlation in 228 patients', Gastroenterology, vol. 100, no. January, pp. 160-167.
Galambos, J T 1979, 'Major problems in internal medicine', Cirrhosis, the series, vol. XVII.
Galbraith, C G, Skalak, R & Chien, S 1998, 'Shear stress induces spatial reorganization of the endothelial cell cytoskeleton', Cell Motilily and the Cytoskeleton, vol. 40, no. 4, pp. 317-330.
Gallego, C, Velasco, M, Marcuello, P, Tejedor, D, De Campo, L & Friera, A 2002, 'Congenital and acquired anomalies of the portal venous system', Radiographics, vol. 22, no. 1, Jan-Feb 2002, pp. 141-59.
Gallix, B P, Reinhold, C, Dauzat, M & Bret, P M 2002, 'Streamlined flow in the portal vein: demonstration with MR angiography', J Magnetic Resonance Imaging, vol. 15, no. 5, May 2002, pp. 603-9.
Gallix, B P, Taourel, P, Dauzat, M, Bruel, J M & Lafortune, M 1997, 'Flow pulsatility in the portal venous system: a study of Doppler sonography in healthy adults', AJR Am J Roentgenol, vol. 169, no. 1, July 1997, pp. 141-4.
Garcia-Tsao, G, Wonggcharatrawee, S & VAHepatitisCResourceCenters 2003, Treatment of Patients withCirrhosis and Portal Hypertension, viewed 2004, www.va.gov/hepatitisc.
Gates, G F & Dore, E K 1973, 'Streamline flow in the human portal vein', Journal of Nuclear Medicine, vol. 14, no. 2, pp. 79-83.
Gauthier, F J, Goldsmith, H L & Mason, S G 1972, 'Flow of suspensions through Tubes-X liquid drops as models of erythrocytes', Biorheology, vol. 9, pp. 205-224.
Geraghty, J G, Angerson, W J & Carter, D C 1989, 'Portal venous pressure and portasystemic shunting in experimental portal hypertension', American Journal of Physiology: Gastrointestinal and Liver Physiology, vol. 257, no. 1, pp. 52-G57.
279
Gerbes, A L, Gulberg, V, Waggershauser, T, Holl, J & Reiser, M 1998, 'Renal effects of transjugular intrahepatic portosystemic shunt in cirrhosis: Comparison of patients with ascites, with refractory ascites, or without ascites', Hepatology, vol. 28, pp. 683-688.
Gibson, P R, Gibson, R N, Donlan, J D, Jones, P A, Colman, J C & Dudley, F J 1996, 'A comparison of Doppler flowmetry with conventional assessment of acute changes in hepatic blood flow', Journal of Gastroenterology and Hepatology, vol. 11, pp. 14-20.
Gibson, P R, Hebbard, G S, Gibson, R N, Firkin, A G & Bhathal, P S 1993, 'Percutaneous transhepatic measurement of the pressure gradient between the portal and hepatic veins', Aust NZ J Med, vol. 23, pp. 374-380.
Gibson, R N, Gibson, P R, Donlan, J D & Padmanabhan, R 1993, 'Modified Doppler Flowmetry in the Splanchnic Circulation', Gastroenterology, vol. 105, pp. 1029-1034.
Gill, R W 1979, 'Pulsed Doppler with B-mode imaging for quantitative blood flow measurement', Journal of Med and Biol, vol. 5, pp. 223-235.
Gill, R W 1985, 'Measurement of blood flow by ultrasound: accuracy and sources of error', Ultrasound in Med. &Biol., Clinical Supplement, vol. 11, no. 4, pp. 625-641.
Glanemann, M, Settmacher, U, Langrehr, J M, Kling, N, Hidajat, N, Stange, B, Staffa, G, Bechstein, W O & Neuhaus, P 2001, 'Portal vein angioplasty using a transjugular, intrahepatic approach for treatment of extrahepatic portal vein stenosis after liver transplantation', Transpl Int, vol. 14, pp. 48-51.
Goldsmith, H L 1971, 'Red cell motions and wall interactions in tube flow', Fed. Proc. 1971, vol. 30, pp. 1578-1590.
Goldstein, A H, Pacella, J J, Trumble, D R & Clark, R E 1992, 'Development of an implantable centrifugal blood pump', ASAIO J, vol. 38, no. 3, Jul-Sep 1992, pp. M362-5.
Gornam, J 2001, 'Liver cells thrive on novel silicon chips', Science News, vol. 159, no. 15, 04 Jan 2001, p. 230.
Grace, N D 1990, 'A hepatologist's view of variceal bleeding', Am J Surg, vol. 160, no. 1, July 1990, pp. 26-31.
Grace, N D 1993, 'Management of portal hypertension', Gastroenterologist, vol. 1, no. 1, March 1993, pp. 39-58.
Grace, N D, Groszmann, R J, Garcia-Tsao, G, Burroughs, A K, Pagliaro, L, Makuch, R W, Bosch, J, Stiegmann, G V, Henderson, J M, de Franchis, R, Wagner, J L, Conn, H O & Rodes, J 1996, 'Portal Hypertension and Variceal Bleeding: An AASLD Single topic Symposium1' An AASLD Single topic Symposium1, Hold over 2 days in June 1996, sponsored by the American Association for the Study of Liver Diseases
Gray, H 1959, 'Anatomy of the human body', in Gross, C M (ed) Anatomy of the human body, Lea&Febiger, Philadelphia, pp. 1294-1305.
Gray, H F R S 1995, Gray’s Anatomy, Fifth edition edn., Barnes and Noble Books, N.Y.
280
Grazi, G L, Cescon, M, Salvi, F, Ercolani, G, Ravaioli, M, Arpesella, G, Magelli, C, Grigioni, F & Cavallari, A 2003, 'Combined heart and liver transplantation for familial amyloidotic neuropathy: Considerations from the hepatic point of view', Liver Transplantation, vol. 9, pp. 986-992.
Greisler, H P, Endean, E D, Klosak, J J, Ellinger, J, Henderson, S C, Pham, S M, Durham, S J, Showalter, D P, Levine, J & Borovetz, H S 1989, 'Hemodynamic effects on endothelial cell monolayer detachment from vascular prostheses', Arch Surg, vol. 124, no. 4, Apr, pp. 429-33.
Grigioni, M, Amodeo, A, Daniele, C, D’Avenio, G, Formigari, R & Di Donato, R M 2000, 'Particle Image Velocimetry Analysis of the Flow Field in the Total Cavopulmonary Connection', Artificial Organs, vol. 24, no. 12, pp. 946-952.
Grosjean, C & Tai, Y-C 1999, 'A thermopneumatic peristaltic micropump' Technical Digest of Transducers’99, Seidai, Japan, Cathech EE 136-93.
Groszmann, R J, Kotelanski, B & Cohn, J N 1971, 'Hepatic Lobular Distribution of Splenic and Mesenteric blood flow in Man', Gastroenterology, vol. 60, no. 6, pp. 1047-1052.
Gruttadauria, S, Marino, G, Catalano, F, Sgroi, A V, Di Mauro, G L & F., B 2001, 'Porcine Orthotopic Liver Autotransplantation: Facilitated Technique', Journal of Investigative Surgery, vol. 14, pp. 79-82.
Gurlek, C, Guleren, K M, Aydin, K & Pinarbasi, A 2002, 'Steady laminar flow computation through vascular tube constrictions' 6th Biennial Conference on Engineering Systems Design and Analysis, Istanbul, Turkey, 8-11 July 2002, Proceedings of ESDA2002/BIO-010
Haag, K, Rossle, M, Ochs, A, Huber, M, Siegerstetter, V, Olschewski, M, Berger, E, Lu, S C & Blum, H E 1999, 'Correlation of duplex sonography findings and portal pressure in 375 patients with portal hypertension', American Journal of Roentgenology, vol. 172, no. 3, March, pp. 631-635.
Habib, N, Houssin, D, Carol, C, Cardoso, J, Boisseau, C, Calmus, Y, Ozier, Y, Quere, J, Wood, C & McIntyre, N 1991, 'Experimental reduction of portal hypertension by mechanical increase of liver portal flow', The Lancer, vol. 337, no. 8723, 5 Jan 1999, pp. 16-18.
Hacking, W J, Van Bavel, E & Spaan, J A 1996, 'Shear stress in not sufficient to control growth of vascular networks: a model study', American Journal of Physiology: Heart and Circulation Physiology, vol. 270, no. 1, pp. H364-H375.
Haidekker, M A, Tsai, A G, Brady, T, Stevens, H Y, Frangos, J A & Theodorakis, E A 2002, 'A Novel Approach to Blood Plasma Viscosity Measurement Using Fluorescent Molecular Rotors', Amer. J. Physiol. Heart Circ. Physiol. , vol. 282, pp. H1609-H1614.
Harry, R & Wendon, J 2002, 'Management of variceal bleeding', Current opinion in critical care, vol. 8, pp. 164-170.
Haskal, Z J, Weintraub, J L & Susman, J 2002, 'Recurrent TIPS thrombosis after polyethylene stent-graft use and salvage with polytetrafluoroethylene stent-grafts', Journal of Vascular and Interventional Radiology, vol. 13, no. 12, Dec 2002, pp. 1255-9.
281
Hay, P D, Veitch, A R, Smith, M D, Cousins, R B & Gaylor, J D S 2000, 'Oxygen Transfer in a Diffusion-Limited Hollow Fiber Bioartificial Liver', Artificial Organs, vol. 24, no. 4, pp. 278-288.
Haynes, W G, Clarke, J G, Cockcroft, J R & Webb, D J 1991, 'Pharmacology of endothelin-1 in vivo in humans', J Cardiovasc Pharmacol, vol. 17, no. Supl 7, pp. S284-6.
Hedeman, J P P, Verhagen, H J, Heijnen-Snyder, G J, van Vroonhoven, T J, Sixma, J J, de Groot, P G & Eikelboom, B C 1998, 'Thrombogenesis of different cell types seeded on vascular grafts and studied under blood-flow conditions', J Vasc Surg, vol. 28, no. 6, Dec, pp. 1094-1103.
Heilshorn, S C, DiZio, K A, Welsh, E R & Tirrell, D A 2003, 'Endothelial cell adhesion to the fibronectin CS5 domain in artificial extracellular matrix proteins', Biomaterials, vol. 24, no. 23, Oct 2003, pp. 4245-52.
Helton, W S, Maves, R, Wicks, K & Johansen, K 2001, 'Transjugular Intrahepatic Portasystemic Shunt vs Surgical Shunt in Good-Risk Cirrhotic Patients: A Case-Control Comparison', Arch Surg, vol. 136, Jan 2001, pp. 17-20.
Henderson, J M 1995, 'Portal hypertension – the surgical pendulum', The Western Journal of Medicine, vol. 162, no. 6, June, p. 554(2).
Henderson, J M, Gilmore, G T, Hooks, M A, Galloway, J R, Dodson, T F, Hood, M M, Kutner, M H & Boyer, T D 1992, 'Selective shunt in the management of variceal bleeding in the era of liver transplantation', Ann Surg, vol. 216, p. 248.
Henry, F S, Shortland, A P, Ludicello, F, Black, R A, Jarvis, J C, Collins, M W & Salmons, S 1997, 'Flow in simple model skeletal muscle ventricle: comparison between numerical and physical simulations', Journal of Biomechanical Engineering, vol. 119, no. 1, Feb 1997, pp. 13-19.
Henry, P D & Chen, C H 1993, 'Inflammatory mechanisms of atheroma formation. Influence of fluid mechanics and lipid-derived inflammatory mediators', Am J Hypertens, vol. 6, no. 11 Pt 2, Nov 1993, pp. 328S-334S.
Herring, M, Smith, J, Dalsing, M, Glover, J, Compton, R, Etchberger, K & Zollinger, T 1994, 'Endothelial seeding of polytetrafluoroethylene femoral popliteal bypasses: the failure of low-density seeding to improve patency', J Vasc Surg, vol. 20, no. 4, Oct 1994, pp. 650-5.
Hidajat, N, Vogl, T, Stobbe, H, Schmidt, J, Wex, C, Lenzen, R, Berg, T, Neuhaus, P & Felix, R 2000, 'Transjugular intrahepatic portosystemic shunt. Experiences at a liver transplantation center', Acta Radiologica, printed in Denmark, vol. 41, pp. 474-478.
Hinds, M T, Park, Y J, Jones, S A, Giddens, D P & Alevriadou, B R 2001, 'Local hemodynamics affect monocytic cell adhesion to a three-dimensional flow model coated with E-selectin', Journal of Biomechanics, vol. 34, pp. 95-103.
Hirakawa, S, Rothe, C F, Shoukas, A A & Tyberg, J V 1993, Veins. Their Functional Role in the Circulation, Springer-Verlag Tokyo.
Hiromichi, S, Takamizawa, K, Nakayama, Y, Yasui, H & Matsuda, T 2001, 'Small-diameter compliant arterial graft prosthesis: Design concept of coaxial
282
double turbular graft and its fabrication', Inc J Biomed Mater Res, vol. 55, pp. 266-276.
Hoerstrup, S P, Sodian, R, Sperling, J S, Vacanti, J P & Mayer, J E J 2000, 'Technical report: New Pulsatile Bioreactor for In Vitro Formation of Tissue Engineered Heart Valves', Tissue Engineering, vol. 6, no. 1, pp. 75-79.
Hűbner, G, Steudel, N, Kleber, G, Behrmann, C, Lotterer, E & Fleig, W E 2000, 'Hepatic arterial blood flow velocities: assessment by transcutaneous and intravascular Doppler sonography', Journal of Hepatology vol., no. 32, June 2000, pp. 893-899.
Huff, M 1997, MEMS pump control and monitors drug flow in the body cat. no. Case Western Reserve University, Cleaveland, OH, USA,. Viewed 07/07/2000, www.spie.org/web/oer/august/aug97/instrument.html
Hutmacher, D W, Schantz, T, Zein, I, Ng, K W, Teoh, S H & Tan, K C 2001, 'Mechanical properties and cell cultural response of polycaprolactone scaffolds designed and fabricated via fused deposition modelling', Inc. J Biomed Mater Res, vol. 55, pp. 203-216.
In Seok Kang 2002, 'A microscopic study on the rheological properties of human blood in low concentration limit', Korea-Australia Rheology Journal, vol. 14, no. 2, June, pp. 77-86.
Institute fur Mikrotechnik Mainz, G 1999, 'Injection Moulded Micropumps', Liga news Fourth Issue, Institute fur Mikrotechnik Mainz GmbH, April 1999, Viewed 17/07/2000,
Iredale, J P 2003, 'Cirrhosis: new research provides a basis for rational and targeted treatments', BMJ, vol. 327, 19 July 2003, pp. 143-147.
Jalan, R, Sen, S & Williams, R 2004, 'Prospects for Extracorporeal Liver Support. Recent Advances in Clinical Practice', Gut, vol. 53, issue 6, pp. 890-898.
Jang, L-S, Morris, C J, Sharma, N R, Bardell, R L & Foster, F K 1999, 'Transport of particle-laden fluids through fixed-valve micropumps' ASME 1999,
Jang, L-S, Sharma, N R & Foster, F K 2000, The effect of particles on performance of fixed-valve micropumps, cat. no. Netherlands Viewed
Jarrell, B E, Williams, S K, Rose, D, Garibaldi, D, Talbot, C & Kapelan, B 1991, 'Optimization of human endothelial cell attachment to vascular graft polymers', J Biomed Eng, vol. 113, no. 2, May, pp. 120-2.
Jensen, J A 1996, Estimation of Blood Velocities Using Ultrasound. A Signal Processing Approach, Cambridge University Press.
Jiao, L R, Seifalian, A M, Davidson, B R & Habib, N 2000, 'In vivo evaluation of an implantable portal pump system for augmenting liver perfusion', British Journal of Surgery, vol. 87, pp. 1024-1029.
Jin, G L & Rikkers, L F 1991, 'The significance of portal vein thrombosis after distal splenorenal shunt', Archives of Surgery, vol. 126, no. 8, August 1991.
Johnson & Johnson Gateway Ethicon, viewed www.ecatalog.ethicon.com. Kaibara, M, Iwata, H, Wada, H, Kawamoto, Y, Iwaki, M & Suzuki, Y 1996,
'Promotion and control of selective adhesion and proliferation of endothelial cells on polymer surface by carbon deposition', J Biomed Mater Res, vol. 31, no. 3, July, pp. 429-435.
283
Kaihara, S, Borenstein, J, Koka, R L, S., Ochoa, E R, Ravens, M, Pien, H, Cunningham, B & Vacanti, J 2000, 'Silicon micromachining to tissue engineer branched vascular channels for liver fabrication', Tissue Eng, vol. 6, no. 2, Apr, pp. 105-117.
Kam, L & Boxer, S G 2001, 'Cell adhesion to protein-micropatterned-supported lipid bilayer membranes', Inc. J Biomed Mater Res, vol. 55, pp. 487-495.
Kane, R A & Katz, S C 1982, 'The spectrum of sonographic findings in portal hypertension: a subject review and new observations', Radiology, vol. 142, no. 453-458.
Kaneko, Z, Kotani, H, Komuta, K & Satomura, S 1961, 'Studies on peripheral circulation by Ultrasonic Blood-Rheograph', Jpn Circ.J, vol. 25, pp. 203-213.
Karau, K L, Krenz, G S & Dawson, C A 2001, 'Branching exponent heterogeneity and wall shear stress distribution in vascular trees', American Journal of Physiology: Heart and Circulation Physiology, vol. 280, no. 3, March 2001, pp. H1256-H1263.
Karimov, S, Borovskii, S P, Rakhmanov, S U & Atakhanov, D A 2003, 'Endovascular embolization of the hepatolienal vessels in the treatment of portal hypertension', Angiol Sosud Khir., vol. 9, no. 2, pp. 40-46.
Kawasaki, T, Itani, T, Mimura, J & Komori, H 1999, 'Portal venous volume flow: in vivo measurement by time-domain color-velocity imaging', Ultrasound in Medicine & Biology, vol. 25, no. 6, July 1999, pp. 915-921.
Kawashita, Y, Ohtsuru, A, Fujioka, H, Kamohara, Y, Kawazoe, Y, Sugiyama, N, Eguchi, S, Kuroda, H, Furui, J, Yamashita, S, Kaneda, Y & Kanematsu, T 2000, 'Safe and Efficient Gene Transfer into Porcine Hepatocytes Using Sendai Virus-Cationic Liposomes for Bioartificial Liver Support', Artificial Organs, vol. 24, no. 12, pp. 932-938.
Kaymakoglu, S, Kahraman, T, Kudat, H, Demir, K, Cakaloglu, Y, Adalet, I, Dincer, D, Besisik, F, Boztas, G, Sozen, A B, Mungan, Z & Okten, A 2003, 'Hepatopulmonary syndrome in noncirrhotic portal hypertensive patients', Digestive Diseases & Sciences, vol. 48, no. 3, March, pp. 556-560.
Keiding, S, Solvig, J, Gronbaek, H & Vilstrup, H 2004, 'Combined liver vein and spleen pulp pressure measurements in patients with portal or splenic vein thrombosis', Scandinavian Journal of Gastroenterology, vol. 39, no. 6, June 2004, pp. 594-600.
Keller, M S, Taylor, K J & Riely, C A 1989, 'Pseudoportal Doppler signal in the partially obstructed inferior vena cava', Radiology, vol. 170, no. 2, Feb 1989, pp. 475-7.
Kent Scientific Corporation, U S A 1997, www.kentscientific.com/99WebCatalog/97.html
Khalili, T M, Navarro, A, Ting, P, Kamohara, Y, Arkadopoulos, N, Solomon, B A, Demetriou, A A & Rozga, J 2001, 'Bioartificial Liver Treatment Prolongs Survival and Lowers Intracranial Pressure in Pigs with Fulminant Hepatic Failure', Artificial Organs, vol. 25, no. 7, July 2001, p. 566.
Kim, S S 2000, 'Dynamic seeding and in vitro culture of hepatocytes in a flow perfusion system', Tissue Eng, vol. 6, no. 1, Feb, pp. 39-44.
284
Kinley, C E, Paasche, B E, MacDonald, A S & Marble, A E 1974, 'Stress at vascular anastomosis in relation to host artery: Synthetic graft diameter', Surgery, vol. 75, pp. 128-130.
Kiserud, T, Kilavuz, O & Hellevik, R L 2003, 'Venous pulsation in the fetal left portal branch: the effect of pulse and flow direction', Ultrasound in Obstetrics & Gynecology, vol. 21, no. 4, Apr, pp. 359-364.
Klabunde, R E 2003, Cardiovascular Physiology Concepts, viewed Feb 2003, www.oucom.ohiou.edu/cvphysiology.
Kobashi, T & Matsuda, T 2000, 'Branched Hybrid Vessel: In Vitro Loaded Hydrodynamic Forces Influence the Tissue Architecture', Cell Transplantation, vol. 9, pp. 93-105.
Koda, M, Murawaki, Y & Kawasaki, H 2000, 'Renovascular resistance assessed by color Doppler ultrasonography in patients with chronic liver diseases', Journal of Gastroenterology and Hepatology, vol. 15, pp. 1424-1429.
Kofidis, T, Lenz, A, Boublik, J, Akhyari, P, Wachsmann, B, Mueller-Stahl, K, Hofmann, M & Haverich, A 2003, 'Pulsatile perfusion and cardiomyocyte viability in a solid three-dimensional matrix', Biomaterials, vol. 24, no. 27, Dec 2003, pp. 5009-5014.
Kofidis, T, Lenz, A, Boublik, J, Akhyari, P, Wachsmann, B, Stahl, K M, Haverich, A & Leyh, R G 2003, 'Bioartificial grafts for transmural myocardial restoration: a new cardiovascular tissue culture concept', European Journal of Cardio-thoracic Surgery, vol. 24, pp. 906-911.
Kofodis, T, Lenz, A, Boublik, J, Akhyari, P, Wachsmann, B, Mueller-Stahl, K, Haverich, A & Leyh, R G 2003, 'Bioartificial grafts for transmural myocardial restoration: a new cardiovascular tissue culture concept', European Journal of Cardio-Thoracic Surgery, vol. 24, no. 6, Dec 2003, p. 906.
Kok, T, van der Jagt, E J, Haagsma, E B, Bijleveld, C M, Jansen, P L & Boeve, W J 1999, 'The value of Doppler ultrasound in cirrhosis and portal hypertension', Scandinavian Journal of Gastroenterology, vol. 34, no. Suppl.230, pp. 82-88.
Koslin, D B, Mulligan, S A & Berland, L L 1992, 'Duplex assessment of the portal venous system', Semin Ultrasound CT MR, vol. 13, no. 1, Feb 1992, pp. 22-33.
Krahenbuhl, L, Seiler, C A & Buchler, M W 1999, 'Variceal hemorrhage in portal hypertension: role of surgery in the acute and elective situation', Schweiz Med Wochenschr, in German, vol. 129, no. 16, 24 April 1999, pp. 631-638.
Krieger, I M & Elrod, H 1953, 'Direct determination of the flow curves of non-Newtonian fluids. II. Shearing rate in the concentric cylinder viscometer', J Applied Phys, vol. 24, pp. 134-136.
Krieger, I M & Maron, S H 1954, 'Direct determination of the flow curves of non-Newtonian fluids. II. Standardized treatment of viscometric data', J Applied Phys, vol. 25, pp. 72-75.
Krupski, G, Broring, D C, Wittkugel, O, Muller, L, Nicolas, V, Rogiers, X, Adam, G & Bucheler, E 2002, 'Formation of portal venous collaterals after ligation of
285
the portal vein for induction of liver regeneration', Rofo Fortschr Geb Ronrgenstr Neuen Bildged Verfahr, vol. 174, no. 10, Oct 2002, pp. 1281-4.
Kubo, S, Shiomi, S, Tanaka, H, Shuto, T, Takemura, S, Mikami, S, Uenishi, T, Nishino, Y, Hirohashi, K, Kawamura, E & Kinoshita, H 2002, 'Evaluation of the effect of portal vein embolization on liver function by (99m)tc-galactosyl human serum albumin scintigraphy', J Surg Res, vol. 107, no. 1, Sept 2002, pp. 113-8.
LaBerge, J M, Ferrell, L D & Ring, E J 1993, 'Histopathologic study of stenotic and occluded transjugular intrahepatic portosystemic shunts', J Vasc Interv Radiol, vol. 4, no. 6, Nov-Dec 1993, pp. 779-786.
Lafortune, M, Martinet, J P, Denys, A, Patriquin, H, Dauzat, M, Dufresne, M P, Colombato, L & Pomier-Layrargues, G 1995, 'Short- and long-term hemodynamic effects of transjugular intrahepatic portosystemic shunts: a Doppler/manometric correlative study', American Journal of Roentgenology, vol. 164, pp. 997-1002.
Lai, S T 1997, 'Management of the complications of chronic liver disease', HKMJ, vol. 3, no. 3, Sept.
Lake, J R 2000, 'The Role of Transjugular Portosystemic Shunting in Patients with Ascites', The New England Journal of Medicine, vol. 342, no. 23, 8 June 2000, pp. 1745-1747.
Langer, R & Vacanti, J P 1993, 'Tissue engineering (methods of replacing or substituting for damaged or diseased tissues)', Science, vol. 260, no. 5110, 14 May 1993, p. 920(7).
Lappin, P B, Ross, K L, King, L E, Fraker, P J & Roth, R A 1998, 'The response of pulmonary vascular endothelial cells to monocrotaline pyrrole: cell proliferation and DNA synthesis in vitro and in vivo', Toxicol Appl Pharmacol, vol. 150, no. 1, May 1998, pp. 37-48.
Lata, J, Hulek, P & Vanasek, T 2003, 'Management of acute variceal bleeding', Dig. Dis., vol. 21, no. 1, pp. 6-15.
Legallais, C, Doré, E & Paullier, P 2000, 'Design of a Fluidized Bed Bioartificial Liver ', Artificial Organs, vol. 24, no. 7, pp. 519-525.
Levesque, M J, Liepsch, D, Moravec, S & Nerem, R M 1986, 'Correlation of endothelial cell shape and wall shear stress in a stenosed dog aorta', Arteriosclerosis, vol. 6, pp. 220-229.
Levick, J R 1995, An Introduction to Cardiovascular Physiology, Second edition edn., Butterworth-Heinemann, Oxford; Boston.
Li, J, Pan, J, Zhang, L & Yu, Y 2003, 'Culture of hepatocytes on fructose-modified chitosan scaffolds', Biomaterials, vol. 24, no. 13, June, pp. 2317-2322.
Liepsch, D 1978, 'Berührungsloses Messen von Strömungsgeschwindigkeiten mittels laser-Doppler-anemometry', HLH, vol. 29, pp. 341-344.
Lin, B 2000, 'Development of a Flow Chamber for Cell Adhesion Studies,' Journal of Undergraduate Research University of Florida vol., Oct.
Lin, C H, Lo, G H, Lai, K H, Mok, K T, Pan, H B & Teng, W B 1996, 'Obstructive jaundice caused by calcified portal venous thrombosis: a case report', Zhonghua Yo Xue Za Zhi, vol. 58, no. 1, Jul 1996, pp. 50-3.
286
Lind, C D, Malisch, T W, Chong, W K, Richards, W O, Pinson, C W, Meranze, S G & Mazer, M 1994, 'Incidence of shunt occlusion or stenosis following transjugular intrahepatic portosystemic shunt placement', Gastroent vol. 106, pp. 1277-1283.
Liu, J C, Heilshorn, S C & Tirrell, D A 2004, 'Comparative cell response to artificial extracellular matrix proteins containing the RGD and CS5 cell-binding domains', Biomacromolecules, vol. 5, no. 2, Mar-Apr 2004, pp. 497-504.
Long, Q 1999, The study of blood flow in human arterial bifurcations by the combination of CFD and MRI (BL), thesis, G6h, Ph.D., City, 49-9011.
Long, Q, Xu, X Y, Ariff, B, Thom, S A, Hughes, A D & Stanton, A V 2000, 'Reconstruction of blood flow patterns in a human carotid bifurcation: a combined CFD and MRI study', Journal of Magnetic Resonance Imaging, vol. 11, no. 3, March 2000, pp. 299-311.
Lowe, K C 1999, 'Perfluorinated blood substitutes and artificial oxygen carriers', Blood Rev, vol. 13, no. 3, Sept, pp. 171-84.
Luca, A, García-Pagán, J C, De Lacy, A M, Escorsell, A, Feu, F, Visa, J & Bosch, J 1999, 'Systemic and pulmonary haemodynamics in cirrhotics. Effects of end-to-side portacaval shunt and distal splenorenal shunt on systemic and pulmonary haemodynamics in patients with cirrhosis', Journal of Gastroenterology and Hepatology vol. 14, no. 11, Nov, p. 1112.
Mack, C L, Superina, R A & Whitington, P F 2003, 'Surgical restoration of portal flow corrects procoagulant and anticoagulant deficiencies associated with extrahepatic portal vein thrombosis', REVISTA - The Journal of Pediatrics, vol. 142, no. 2, February 2003, p. JP6327.
Maisaia, K, Mizandari, M, Mtvaradze, A, Tatishvili, D, Mamukashvili, G & Tatanashvili, D 2001, 'Doppler Criteria of Portal Hypertension in Liver Cirrhosis', Annals of Biomedical Research and Education, vol. 1, no. 1, Jan-March 2001, pp. 39-43.
Malagó, M, Hertl, M, Rogiers, X, Burdelski, M & Broelsch, C E 1998, 'Chirurgische Therapie der portalen Hypertension, in German', Der Internist, Springer-Verlag, vol. 39, pp. 280-285.
Malassagne, B, Soubrane, O, Dousset, B, Legmann, P & Houssin, D 1998, 'Extrahepatic portal hypertension following liver transplantation: a rare but challenging problem', HPB Surg, vol. 10, no. 6, pp. 357-63, discussion 363-4.
Mancuso, A, Watkinson, A, Tibballs, J, Patch, D & Burroughs, A K 2003, 'Budd-Chiari syndrome with portal, splenic, and superior mesenteric vein thrombosis treated with TIPS: who dares wins', GUT, vol. 52, p. 438.
Marques, P F, Oliveira, M E C, France, A S & Pinotti, M 2003, 'Modeling and Simulation of Pulsatile Blood Flow with a Physiologic Wave Pattern', Artificial Organs, vol. 27, no. 5, pp. 478-485.
Marseille, O, Habib, N, Reul, H & Rau, G 1988, 'Implantable Micropump System for Augmented Liver Perfusion', Artificial Organs, vol. 22, no. 6, pp. 458-460.
287
Marsutani, S, Maruyama, H, Akiike, T, Sudhmushu, K, Fukuzawa, T, Mizumoto, H & Saisho, H 2003, 'Assessment of hemodynamic response of the left gastric vein to propranolol with Doppler sonography in pharmacotherapy of portal hypertension', Ultrasound in Medicine & Biology, vol. 29, Supplement 1, no. 5, May 2003, p. S60.
Marsutani, S, Maruyama, H, Sato, G, Fukuzawa, T, Mizumoto, H & Saisho, H 2003, 'Hemodynamic response of the left gastric vein to glucagon in patients with portal hypertension and esophageal varices', Ultrasound in Med. &Biol., , vol. 29, no. 1, January, pp. 13-7.
Martin, I, Wendt, D & Heberer, M 2004, 'The role of bioreactors in tissue engineering', Trends in Biotechnology, vol. 22, no. 2, Feb 2004, p. 80.
Masayuki, Y, Kwon, O H, Hirose, M, Kikuchi, A & Okano, T 2001, 'Novel patterned cell coculture utilizing thermally responsive grafted polymer surfaces', Inc J Biomed Mater Res, vol. 55, pp. 137-140.
Mathie, R T & Alexander, B 1990, 'The role of adenosine in the hyperaemic response of the hepatic artery to portal vein occlusion (the “buffer response”)', Br J Pharmacology, vol. 100, pp. 626-630.
Mathie, R T & Blumgart, L H 1983, 'The hepatic haemodynamic response to acute portal venous blood flow reduction in the dog', Pfülgers Arch, vol. 399, pp. 223-227.
Mathur, S K, Shah, S R, Nagral, S S & Soonawala, Z F 1999, 'Transabdominal extensive esophagogastric devascularization with gastroesophageal stapling for management of noncirrhotic portal hypertension: long-term results', World Journal of Surgery, vol. 23, no. 11, Nov 1999, pp. 1168-75.
Matsumoto, S, Klein, A & Maeda, R 1999, 'Valve-less Micropump Based on Temperature Dependence of Liquid Viscosity' MEMS, www.aist.go.jp/MEL/soshiki/kyoku/ryoshi/matsumoto/pump.html
Matsumoto, T, Kobayashi, S, Shimizu, H, Nakajima, M, Watanabe, S, Kitami, N, Sato, N, Abe, H, Aoki, Y, Hoshi, T & Hashimoto, H 2000, 'The liver in collagen diseases: pathologic study of 160 cases with particular reference to hepatic arteritis, primary biliary cirrhosis, autoimmune hepatitis and nodular regenerative hyperplasia of the liver', Liver, vol. 20, pp. 366-373.
Matsutani, S, Furuse, J, Ishii, H, Mizumoto, H, Kimura, K & Ohto, M 1993, 'Hemodynamics of the left gastric vein in portal hypertension', Gastroenterology, vol. 105, no. 2, 2 August 1993, pp. 513-518.
McKeown, N B, Chang, G, Niven, J, Romaschin, A D, Wilson, G J, Thompson, M & Kalman, P G 1991, 'Facilitation of endothelial cell growth on hydroxylated ePTFE vascular grafts', ASAIO Trans, vol. 37, no. 3, Jul-Sept, pp. M477-8.
MedicineNet.com 2003, http://www.medterms.com/. MEMS, G 1997, University of Southampton, Department of Electronics and
Computer Science, UK. Menegaux, F, Keeffe, E B, Baker, E, Egawa, H, Concepcion, W, Russell, T R &
Esquivel, C O 1994, 'Comparison of transjugular and surgical portosystemic shunts on the outcome of liver transplantation', Arch Surg, vol. 129, Oct 1994, pp. 1018-1023.
288
Meng-Yen, T & Jui-Che, L 2001, 'Surface characterization and platelet adhesion studies of self-assembled monolayer with phosphonate ester and phosphonic acid functionalities', Inc. J Biomed Mater Res, vol. 55, pp. 554-565.
Merkel, C, Schipilliti, M, Bighin, R, Bellini, B, Angeli, P, Bolognesi, M, Vescovi, F & Gatta, A 2003, 'Portal hypertension and portal hypertensive gastropathy in patients with liver cirrhosis: a haemodynamic study', Digestive and Liver Disease, vol. 35, no. 4, April, pp. 269-274.
Merzhanov, A & Damianov, D 1989, 'Sugiura’s operation in portal hypertension', Khirurgiia (Sofia) in Bulgarian, vol. 42, no. 3, pp. 1-7.
Micro Infusions Systems, G 2000, Totally Implantable Drug Delivery System (TI-DDS), viewed 06/07/2000 www.mis-gmbh.com/mds_e.htm
Mid-America Interventional Radiological Society 'Other Vascular Interventions', in VASCULAR AND INTERVENTIONAL RADIOLOGY: A Syllabus for House Officers and Medical Students, Mid-America Interventional Radiological Society, http://www.mirs.org/syllabus/vasdx3c.htm.
Miller, M A & Leavell, L C 1972, Kimber-Gray-Stackpole’s Anatomy and Physiology, 16th edn., Macmillan Publishing Co Inc, N.Y.
Mischiati, C, Puviani, A C, Brogli, M, Guarniero, S, Sereni, A, Breda, L, Ricci, D, Galavotti, D, Morsiani, E & Gambari, R 2003, 'Modulation of pro-apoptotic (Bax) and anti-apoptotic (Bcl-2) gene expression in isolated porcine hepatocytes perfused within a radial-flow bioreactor after low-temperature storing', International Journal of Artificial Organs, vol. 26, no. 2, Feb 2003, pp. 139-148.
Mitchel, A, Mirza, D, De Ville de Goyet, J & Buckels, J 2000, 'Absence of the left portal vein: A difficulty for reduction of liver grafts?' Transplantation, vol. 69, no. 8, 27 April, pp. 1731-1732.
Mitchell, S L & Niklason, L E 2003, 'Requirements for growing tissue-engineered vascular grafts', Cardiovasc. Pathol., vol. 12, no. 2, Mar-Apr, pp. 59-64.
Miyamoto, Y, Oho, K, Kumamoto, M, Toyonaga, A & Sata, M 2003, 'Balloon-occluded retrograde transvenous obliteration improves liver function in patients with cirrhosis and portal hypertension', Journal of Gastroenterology and Hepatology, vol. 18, no. 8, Aug, pp. 934-942.
Mooney, M 1931, 'Explicit Formulas for Slip and Fluidity', J Rheology, vol. 2, pp. 210-222.
Moore, J A, Steinman, D A & Ethier, C R 1998, 'Computational blood flow modelling: Errors associated with reconstructing finite element models from magnetic resonance images', Journal of Biomechanics, vol. 31, pp. 179-184.
Morales-Ruiz, M, Cejudo-Martin, P, Fernando-Varo, G, Tugues, S, Ros, J, Angeli, P, Rivera, F, Arroyo, V, Rodes, J, Sessa, W C & Jimenez, W 2003, 'Transduction of the liver with activated AKT normalized portal pressure in cirrhotic rats', Gastroenterology, vol. 125, no. 2, August 2003, pp. 522-31.
Moreno, A H, Burchell, A R, Reddy, R V, Steen, J A, Panke, W F & Nealon, T F J 1975, 'Spontaneous Reversal of Portal Blood flow: The case For and
289
Against Its Occurrence in Patients With Cirrhosis of the Liver', Ann. Surg, vol. 181, no. 3, March, pp. 346-357.
Mori, Y 1989, Vascular Dynamics. Physiological Perspectives, New York Plenum Press, New York.
Moriyasu, F, Nishida, O, Ban, N, Nakamura, T, Miura, K, Sakai, M, Miyake, T & Uchino, H 1986, 'Measurement of Portal Vascular Resistance in Patients With Portal Hypertension', Gastroenterology, vol., no. 90, pp. 710-7.
Morsi, Y S, Damen, B S, Bos, S, Palombo, E, Narasimhan, V & Tavrou, S 2001, 'Preliminary study on multifunctional bio reactor for in vitro tissue engineering' First Asian-Pacific Congress on Computational Mechanics, Sydney Australia,, 20-23 November,
Moser, M A, Kneteman, N M & Minuk, G Y 2000, 'Research toward safer resection of the cirrhotic liver', HPB Surg, vol. 11, no. 5, pp. 285-97.
MPRSGGroup 2001, Multifunctional bioreactor for in vitro tissue engineering, cat. no. School of Engineering and Science, Swinburne University of Technology, Melbourne. Viewed
Munk, O L, Keiding, S & Bass, L 2003, 'Impulse-response function of splanchnic circulation with model-independent constraints: theory and experimental validation', Am Journal Physiology: Gastrointest Liver Physiology, vol. 285, pp. G671-680.
Nadal, C 2000, 'Nonregenerative stimulation of hepatocyte proliferation in the rat: Variable effects in relation to spontaneous liver growth; a possible link with metabolic induction', Cell Proliferation October, vol. 33, no. 5, Oct 2000, pp. 287-300.
Nagafuchi, Y, Watanabe, H, Kitahara, K, Okamoto, K, Hirata, K, Tabaru, A, Otsuki, M, Nakata, H & Itho, H 1996, 'Intrahepatic portosystemic venous shunt: report of two cases assessed by helical computed tomography', Scand J Gastroenterol, vol. 31, no. 11, Nov 1996, pp. 1132-5.
Nakamura, M & Sawada, T 1988, 'Numerical study on the flow of a non-Newtonian fluid through an axisymmetric stenosis', J. Biomech. Engng, vol. 110, pp. 137-143.
Nakata, K, Yoshikawa, M, Takano, T, Maeda, T, Nonaka, K, Linneweber, J, Kawahito, S, Glueck, J, Fujisawa, A, Makinouchi, K, Yokokawa, M & Nose, Y 2000, 'Antithrombogenicity Evaluation of a Centrifugal Blood Pump', Artificial Organs, vol. 24, no. 8, pp. 667-670.
Nam, K J, Lee, J H, Choi, J C, Oh, J Y, Yoon, S K, Shin, T B & Park, B H 2003, 'Congenital arterioportal fistula and portovenous shunt associated with intrahepatic portal vein', Ultrasound in Med. &Biol.,, vol. 29, no. 5, Suppl. 1, May, p. S162.
Naruse, K, Sakai, Y, Lei, G, Sakamoto, Y, Kobayashi, T, Puliatti, C, Aronica, G, Morale, W, Leone, F, Qiang, S, Ming, S G, Ming, S, Li, Z, Chang, S J, Suzuki, M & Makuuchi, M 2001, 'Efficacy of Nonwoven Fabric Bioreactor Immobilized with Porcine Hepatocytes for Ex Vivo Xenogeneic Perfusion Treatment of Liver Failure in Dogs', Artificial Organs, vol. 25, no. 4, pp. 273-280.
290
Nazarian, G K, Ferral, H, Castaneda-Zuniga, W R, Bjarnason, H, Foshager, M C, Rank, J M, Anderson, C A, Rengel, G J, Herman, M E & Hunter, D W 1994, 'Development of stenoses in transjugular intrahepatic portosystemic shunts', Radiology, vol. 192, pp. 231-234.
Nerem, R M 1991, 'Cellular engineering', Annals of Biomedical Engineering, vol. 19, pp. 529-545.
Nerem, R M, Alexander, R W, Chappell, D C, Medford, R M, Varner, S E & Taylor, W R 1998, 'The study of the influence of flow on vascular endothelial biology', American Journal of Medical Sciences, vol. 316, no. 3, Sep 1998, pp. 169-75.
Netter, F H 1964, 1957, The CIBA Collection of Medical Illustrations, Digestive System part III, Liver, Biliary Tract and pancreas.
Neuhaus, P & Blumhardt, G 1991, 'Surgery for portal hypertension', Hepatogastroenterology, vol. 38, no. 5, Oct 1991, pp. 355-359.
Neuzil, D F, Rozga, J, Moscioni, A D, Man-Soo, R, Hakim, R, Arnaout, W S & Demetriou, A A 1993, 'Use of a novel bioartificial liver in a patient with acute liver insufficiency', Surgery, vol. 113, no. 3, pp. 340-343.
Newby, D E & Hayes, P C 2002, 'Hyperdynamic circulation in liver cirrhosis: not peripheral vasodilatation but ‘splanchnic steal’', QJMed, vol. 95, pp. 827-830.
Niklason, L E, Abbott, W, Gao, J, Klagges, B, Hirschi, K K, Ulubayram, K, Conroy, N, Jones, R, Vasanawala, A, Sanzgiri, S & Langer, R 2001, 'Morphologic and mechanical characteristics of engineered bovine arteries', J Vascular Surgery, vol. 33, no. 3, March, pp. 628-638.
Niklason, L E, Gao, J, Abbott, W M, Hirschi, K K, Houser, S, Marini, R & Lager, R 1999, 'Functional Arteries Grown in Vitro', Science, vol. 284, no. 5413, 16 April 1999, p. 489(1).
Nikolaychik, V V, Samet, M M & Lelkes, P I 1994, 'A new, cryoprecipitate based coating for improved endothelial cell attachment and growth on medical grade artificial surfaces', ASAIO J, vol. 40, no. 3, Jul-Sept, pp. M846-52.
Nishimura, Y, Meyns, B, Ozaki, S, Racz, R, Dohmen, P & Flameng, W 1999, 'The enabler cannula pump: a novel circulatory support system', Int J Artif Organs, vol., no. 22, pp. 317-323.
Niu, Y-Y, Wei-Kuang, C, Shou-Cheng, T, Lung-Cheng, L & Hsi-Yu, Y 2002, 'Numerical Simulations of Aortic Flow Problems Based on a Parallelized Artificial Compressibility Solver' Proceedings.of the Secound International Conference on Computational Fluid Dynamics, ICCFD, Sydney, Australia, 15-19 July 2002, 815-816.
Niwa, T, Aida, N, Tachibana, K, Shinkai, M, Ohhama, Y, Fujita, K, Abe, A, Lee, J, Ozawa, Y & Inoue, T 2002, 'Congenital absence of the portal vein: clinical and radiological findings', J Comput Assist Tomogr, vol. 26, no. 5, Sep-Oct 2002, pp. 681-686.
Noishiki, Y & Miyata, T 1986, 'A simple method to heparinize biological materials', J Biomed Mater Res, vol. 20, no. 3, Mar 1986, pp. 337-46.
Northrup, M, Mendez-Castillo, A, Sethi, Y & Churchill, R 2002, 'Congenital absence of the portal vein with an intrahepatic inferior vena cava branch
291
showing hepatopetal flow', Journal Of Ultrasound In Medicine: Official Journal Of The American Institute Of Ultrasound In Medicine, vol. 21, no. 5, May 2002, pp. 569-572.
Nosaka, T, Teramoto, K, Tanaka, Y, Igari, T, Takamatsu, S, Kawamura, T, Inoue, Y, Goseki, N, Arii, S, Iwai, T, Inomata, Y & Tanaka, K 2003, 'Varicose bleeding after liver transplantation in a patient with severe Portosystemic shunts', Journal of Gastroenterology, vol. 38, no. 7, pp. 700-3.
Nose, Y, Yoshikawa, M, Murabayashi, S & Takano, T 2000, 'Development of rotary blood pump technology: past, present and future', Artificial Organ vol., no. January.
Ochs, A, Rössle, M, Haag, K, Hauenstein, K-H, Deibert, P, Siegerstetter, V, Huonker, M, Langer, M & Blum, H E 1995, 'The Transjugular Intrahepatic Portosystemic Stent-Shunt Procedure for Refractory Ascites', The New England Journal of Medicine, vol. 332, no. 18, 4 May 1995, pp. 1192-1197.
Ogawa, A, Takamori, S, Kojima, K, Fukasawa, M, Beppu, T & Futagawa, S 2002, 'Hemodynamics in extrahepatic portal vein obstruction and its changes during long-term follow-ups', Hepatology Research, vol. 24, no. 2, Oct 2002, pp. 141-151.
Ohashi, K, Kojima, K, Fukazawa, M, Beppu, T & Futagawa, S 1998, 'Long-Term prognosis of non-shunt operation for idiopathic portal hypertension', Journal of Gastroenterology, vol. 33, pp. 241-246.
Ohnishi, K, Chin, N, Sugita, S, Saito, M, Tanaka, H, Terabayashi, H, Saito, M, Iida, S, Nomura, F & Okuda, K 1987, 'Quantitative aspects of portal-systemic and arteriovenous shunts within the liver in cirrhosis', Gastroenterology, vol. 93, pp. 129-134.
Oikawa, H, Masuda, T, Sato, S, Yashima, A, Suzuki, K, Sato, S & Satodate, R 1998, 'Changes in Lymph Vessels and Portal Veins in the Portal Tract of Patients With Idiopathic Portal Hypertension: A Morphometric Study', Hepatology, vol. 27, no. 6, June 1998, pp. 1607-1610.
Oka, S 1980, Cardiovascular Hemorheology, Cambridge University Press. Okazaki, K, Miyazaki, M, Onishi, S & Ito, K 1986, 'Effects of food intake and
various extrinsic hormones on portal blood flow in patients with liver cirrhosis demonstrated by pulsed Doppler with the Octoson', Scand J Gastroenterol, vol. 21, no. 9, pp. 1029-1038.
Okuda, K, Suzuki, K, Musha, H & Arimizu, N 1977, 'Percutaneous transhepatic catheterisation of the portal vein from the study of portal hemodynamics and shunts', Gastroenterology, vol. 73, pp. 279-284.
Olsen, D B 2000, 'The History of Continuous-Flow Blood Pumps', Artificial Organs, vol. 24, no. 6, pp. 491-404.
Oosterbroek, R E, Lammerink, T S J, van den Berg, A & Elwenspoek, M C 2000, Principles and methods for micro fluid handling system, viewed
Orozco, H, Mercado, M A, Martinez, R, Tielve, M, Chan, C, Vasquez, M, Zenteno-Guichard, G & Pantoja, J P 1998, 'Is Splenectomy Necessary in Devascularization Procedures for Treatment of Bleeding Portal Hypertension', Arch, vol. 133, Jan 1998, pp. 36-38.
292
Ortale, J R, Naves De Freitas, A C H & Mello De Castro, C 2000, 'Anatomy of the Intrahepatic Ramification of the Portal Vein in the Right Hemiliver', Cells Tissues Organs, vol. 166, no. 4, pp. 378-387.
Ott, M, Schmidt, H H-J, Cichon, G & Manns, M P 2000, 'Emerging Therapies in Hepatology: Liver-Directed Gene Transfer and Hepatocyte Transplantation', Cells Tissues Organs, vol. 167, no. 2-3, pp. 81-87.
Ozaki, C F, Anderson, J C, Lieberman, R P & Rikkers, L F 1988, 'Duplex Ultrasonography as a Noninvasive Techniques for Assessing Portal Hemodynamics', The American Journal of Surgery, vol. 155, Jan 1988, pp. 70-75.
Pahernik, S A, Thasler, W E, Doser, M, Gomez-Lechon, M J, Castell, M J, Planck, H & Koebe, H G 2001, 'High Density Culturing of Porcine Hepatocytes Immobilized on Nonwoven Polyurethane-Based Biomatrices', Cells Tissues Organs, vol. 168, no. 3, pp. 170-177.
Palder, S B, Kirkman, R L, Whittemore, A D, Hakim, R M, Lazarus, J M & Tilney, N L 1985, 'Vascular access for hemodialysis. Patency rates and results of revision', Ann Surg, vol. 202, no. 2, Aug 1985, pp. 235-9.
Palmer, M 2000/2004, Dr. Melissa Palmer's Guide To Hepatitis and Liver Disease, viewed www.liverdisease.com/
Park, T G 2002, 'Perfusion culture of hepatocytes within galactose-derivatized biodegradable poly(lactide-co-glycolide) scaffolds prepared by gas foaming of effervescent salts', J Biomed Mater Res, vol. 59, no. 1, Jan 2002, pp. 127-35.
Partiquin, H, Lafortune, M, Burns, P & Dauzat, M 1987, 'Duplex Doppler examination in portal hypertension: technique and anatomy', AJR Am J Roentgenol, vol. 149, pp. 71-6.
Patel, N H, Sasadeusz, K J, Seshadri, R, Chalasani, N, Shah, H, Johnson, M S, Namyslowski, J, Moresco, K P & Trerotola, S O 2001, 'Increase in Hepatic Arterial Blood Flow after Transjugular Intrahepatic Portosystemic Shunt Creation and its Potential Predictive Value of Postprocedural Encephalopathy and Mortality', Journal of Vascular and Interventional Radiology, vol. 12, pp. 1279-1284.
Pedley, T J 1980, The Fluid Mechanics of Large Blood Vessels, Cambridge University Press, Cambridge, New York
Petkova, S, Hossain, A, Naser, J & Palombo, E 2003, 'CFD modelling of blood flow in the portal vein hypertension with and without thrombosis' Third International Conference on CFD in the minerals and Process Industries, CSIRO, Melbourne, Australia, 10-12 Dec 2003, 527-530.
Petrov, V 1994, Lectures in Biodynamics and Biorheology, in Bulgarian, UZU Neofit Rilski, Blagoevgrad
Petrovic, L M 2003, 'Liver Transplantation Worldwide Can early liver biopsies predict long-term outcome of the graft?' Liver Transplantation, vol. 9, no. 1, Dec 2003, pp. 99-100.
Poepping, T L, Nikolov, H N, Thorne, M L & Holdsworth, D W 2004, 'A thin-walled carotid vessel phantom for Doppler ultrasound flow studies', Ultrasound Med Biol, vol. 30, no. 8, Aug 2004, pp. 1067-78.
293
Pomier-Layrargues, G, Villeneuve, J P, Deschênes, M, Bui, B, Perreault, P, Fenyves, D, Willems, B, Marleau, D, Bilodeau, M, Lafortune, M & Dufresne, M-P 2001, 'Transjugular intrahepatic portosystemic shunt (TIPS) versus endoscopic variceal ligation in the prevention of variceal rebleeding in patients with cirrhosis: a randomised trial', Gut, vol. 48, March 2001, pp. 390-396.
Pozler, O, Krajina, A, Vanicek, H, Hulek, P, Zizka, J & Michl, A E, P. 2003, 'Transjugular intrahepatic portosystemic shunt in five children with cystic fibrosis: Long-term results', Hepatogastroenterology, vol. 50, no. 52, Jul-Aug, pp. 1111-4.
Pozrikidis, C 1990, 'The axisymmetric deformation of a red blood cell in uniaxial straining flow', J. Fluid Mech, vol. 216, pp. 231-254.
Prakash, K, Regimbeau, J M & Belghiti, J 2003, 'Reconstruction of portal vein using a hepatic vein patch graft after combined hepatectomy and portal vein resection', Am J Surg, vol. 185, no. 3, March 2003, pp. 230-231.
Primignani, M, Carpinelli, L, Preatoni, P, Battaglia, G, Carta, A, Prada, A, Cestari, R, Angeli, P, Gatta, A, Rossi, A, Spinzi, G & De Franchis, R 2000, 'Natural History of Portal Hypertensive Gastropathy in Patients with Liver Cirrhosis', Gastroenterology, vol. 119, no. July, pp. 181-187.
Puel, J, Juillière, Y, Bertrand, M E, Rickards, A F, Sigwart, U & Serruys, P W 1988, 'Early and late assessment of stenosis geometry after coronary arterial stenting', Am J Cardiol, vol. 61, pp. 546-553.
Puviani, A C, Ottolenghi, C, Tassinari, B, Pazzi, P & Morsiani, E 1998, 'An update on high-yield hepatocyte isolation methods and on the potential clinical use of isolated liver cells', Comparative Biochemistry and Physiology – Part A: Molecular and Integrative Physiology, vol. 121, no. 2, pp. 99-109.
Quinn, S F, Kim, J & Sheley, R C 2002, 'Percutaneous inferior vena cava-to-portal vein shunt (PIPS) using a stent-graft: preliminary results', J Endovasc Ther, vol. 9, no. 4, Aug 2002, pp. 503-510.
Ramesh, H, Prakash, K, Kuruvilla, K, Philip, M, Jacob, G, Venugopal, B, Lekha, V & Varma, D 2003, 'Biliary access loops for intrahepatic stones: results of jejunoduodenal anastomosis', ANZ Journal of Surgery, vol. 73, no. 5, May, pp. 306-312.
Ramm, G A, Carr, S C, Bridle, K R, Li, L, Britton, R S, Crawford, D H G, Vogler, C A, Bacon, B R & Tracy, T F, Jr. 2000, 'Morphology of liver repair following cholestatic liver injury: resolution of ductal hyperplasia, matrix deposition and regression of myofibroblasts', Liver, vol. 20, pp. 387-396.
Rapaport, A, Noon, G P & McCollum, C H 1981, 'Polytetrafluoroethylene (PTFE) grafts for hemodialysis in chronic renal failure: Assessment of durability and function at three years', Aust N Z Surg, vol. 51, p. 562.
Reichen, J 1998, http://www.ikp.unibe.ch/lab2/ppnew/pp7/text.htm, viewed 2002, http://www.ikp.unibe.ch/lab2/ppnew/pp7/text.htm
Reidy, M A & Schwartz, S M 1981, 'Endothelial regeneration. III. Time course of intimal changes after small defined injury to rat aortic endothelium', Lab Invest, vol. 44, pp. 301-308.
294
Rengo, C, Brevetti, G, Sorrentino, G, D'Amato, T, Imparato, M, Vitale, D F, Acanfora, D & Rengo, F 1998, 'Portal Vein Pulsatility Ratio Provides a Measure of Right Heart Function in Chronic Heart Failure', Ultrasound in Medicine & Biology, vol. 24, no. 3, March 1998, pp. 327-332.
Reynolds, T B 1955, 'Comparison of wedged hepatic vein pressure with portal vein pressure in human subjects with cirrhosis', The journal of clinical investigation, vol. 34, no. 2, pp. 213-218.
Reynolds, T B, Balfour, D C, Levinson, D C, Magidson, O, Mikkelsen, W P & Pattison, A C 1954, 'Measurement of portal venouspressure by hepatic vein catheterization', AMERICAN JOURNAL OF MEDICINE vol. 17, no. 1, pp. 124-124.
Rhee, P & Sapfeh, I J 1993, 'Portal hypertension', Curr Opin Gen Surg, vol., pp. 249-56.
Richardson, E 1974, 'Deformation and haemolysis of red cells in shear flow' Proceedings of the Royal Society of London,
Richardson, P D I & Withrington, P G 1978, 'Pressure-flow relationships and effects of noradrenaline and isoprenaline on the hepatic arterial and portal venous vascular beds of the dog', J Physiol (London), vol., no. 282, pp. 451-70.
Richardson, P D I & Withrington, P G 1981, 'Liver blood flow', Gastroenterology, vol., no. 81, pp. 159-73.
Richardson, P D I & Withrington, P G 1981, 'Liver Blood Flow. I. Intrinsic and Nervous Control of Liver Blood Flow', Gastroenterology, vol. 81, pp. 159-73.
Richter, S 2000, 'Impact of intrinsic blood flow regulation in cirrhosis: maintenance of hepatic arterial buffer response', Am J Physiol Gastrointest Liver Physiol, vol. 279, no. G, pp. G454-G462.
Rikkers, L F, Burnett, D A, Volentine, G D, Buchi, K N & Cormier, R A 1987, 'Shunt surgery versus endoscopic sclerotherapy for long-term treatment of variceal bleeding. Early results of a randomized trial', Ann Surg, vol. 206, no. 3, Sept 1987, pp. 261-271.
Robertson, M B, Köhler, U, Hoskins, P R & Marshall, I 2001, 'Flow in Elliptic Vessels Calculated for Physiological Waveform', Journal of Vascular Research, vol. 38, no. 1, pp. 73-82.
Rosado, B & Kamath, P S 2003, 'Transjugular intrahepatic portosystemic shunts: An update', Liver Transplantation, vol. 9, no. 3, March, pp. 207-217.
Rosemurgy, A S, Bloomston, M, Ghassemi, P & Serafini, F 2002, 'Preshunt and postshunt portal vein pressure and portal vein-to-inferior vena cava pressure gradients do not predict outcome following partial portal decompression', Am Surg, vol. 68, no. 1, Jan 2002, pp. 70-4.
Rőssle, M, Haag, K, Ochs, A, Sellinger, M, Noldge, G, Peratnau, J-M, Berger, E, Blum, U, Gabelmann, A, Hauenstein, K, Langer, M & Gerok, W 1994, 'The Transjugular Intrahepatic Portosystemic Stent-Shunt Procedure for Variceal Bleeding', The New England Journal of Medicine, vol. 330, no. 3, 20 Jan 1994, pp. 165-171.
295
Rőssle, M, Ochs, A, Gülberg, V, Siegerstetter, V, Holl, J, Deibert, P, Olschewski, M, Reiser, M & Gerbes, A L 2000, 'A comparison of paracenetsis and transjugular intrahepatic portosystemic shunting in patients with ascites', The New England Journal of Medicine, vol. 342, no. 23, June, pp. 1701-1707.
Rowan, J O 1981, Physics and the Circulation Medical Physics Handbook 9, Department of Clinical Physics, Institute of Neurological Sciences, Glasgow.
Sacerdoti, D, Merkel, C, Bolognesi, M, Amodio, P, Angeli, P & Gatta, A 1995, 'Hepatic Arterial Resistance in Cirrhosis With and Without Portal Vein Thrombosis: Relationship with Portal Hemodynamics', Gastroenterology vol. 108, pp. 1152-1158.
Saito, N, Motoyama, S & Sawamoto, J 2000, 'Effects of New Polymer-Coated Extracorporeal Circuits on Biocompatibility During Cardiopulmonary Bypass', Artificial Organs, vol. 24, no. 7, pp. 547-554.
Salwei, R M, O'Brien, R T & Matheson, J S 2003, 'Use of contrast harmonic ultrasound for the diagnosis of congenital portosystemic shunts in three dogs', Vet Radiol Ultrasound, vol. 44, no. 3, May-Jun 2003, pp. 301-5.
Sanyal, A J 2000, 'The use and misuse of transjugular intrahepatic portasystemic shunts', Curr Gastroenterol Rep, vol. 2, no. 1, Feb, pp. 61-71.
Sarin, S K, Sollano, J D, Chawla, Y K, Amarapurkar, D, Hamid, S, Hashizume, M, Jafri, W, Kumar, A, Lesmana, L A, Sharma, B C, Shisha, G, de Silva, H J & APASL WOrking Party on Portal Hypertension 2006, 'Consensus on extra-hepatic portal vein obstruction', Liver International, vol. 26, pp. 512-519.
Sauer, I M & Gerlach, J C 2002, 'Modular Extracorporeal Liver Support', Artificial Organs, vol. 26, no. 8, August, pp. 703-706.
Sauer, I M, Neuhaus, P & Gerlach, J C 2002, 'Concept for modular extracorporeal liver support for the treatment of acute hepatic failure', Metabolic Brain Disease, vol. 17, no. 4, Dec, pp. 477-484.
Scantlebury, V P, Zajko, A B, Esquivel, C O, Marino, I R & Starzt, T E 1989, 'Successful reconstruction of late portal vein stenosis after hepatic transplantation', Arch Surg, vol. 124, Apr 1989, pp. 503-505.
Schalina, L & Liepschb, D 2001, 'Laser-Doppler-anemometry studies of flow behavior in rigid and elastic varicose vein models', Technology and Health Care, IOS Press, vol. 9, pp. 378-402.
Schenk, W G J, McDonald, J C, McDonald, K & Drapanas, T 1962, 'Direct measurement of hepatic blood flow in surgical patients: with related observations on hepatic flow dynamics in experimental animals', Ann Surg, vol. 156, Sep 1962, pp. 463-71.
Schiedermaier, P 2004, 'Splanchnic hemodynamics: cirrhotic versus non-cirrhotic portal hypertension', Journal of Gastroenterology and Hepatology, vol. 19, pp. S150-S154.
Schneider, P A, Hanson, S R, Price, T M & Harker, L A 1988, 'Preformed confluent endothelial cell monolayers prevent early platelet deposition on
296
vascular prostheses in baboons', J Vasc Surg, vol. 8, no. 3, Sep 1988, pp. 229-35.
Scott, W J & Mann, P 1990, 'Substrate effects on endothelial cell adherence rates', ASAIO Trans, vol. 36, no. 3, Jul-Sept, pp. M739-41.
Seflon, M V, Sawyer, A, Gorbet, M, Black, J P, Cheng, E, Gemmell, C & Pottinger-Cooper, E 2001, 'Does surface chemistry affect thrombogenicity of surface modified polymers?' Inc. J Biomed Mater Res, vol. 55, pp. 447-458.
Seifalian, A M, Salacinski, H J, Punshon, G, Krijgsman, B & Hamilton, G 2001, 'A new technique for measuring the cell growth and metabolism of endothelial cells seeded on vascular prostheses', Inc. J Biomed Mater Res vol. 55, pp. 637-644.
Seiler, R W & Mariani, L 2000, 'Sellar reconstruction with resorbable Vicryl patches, gelatin foam, and fibrin glue in transsphenoidal surgery: a 10-year experience with 376 patients', Preview for J Neurosurg, vol. 93, pp. 772-775.
Sekido, H, Matsuo, K-i, Takeda, K, Morioka, D, Kubota, T, Tanaka, K, Endo, I, Togo, S, Tanaka, K & Shimada, H 2002, 'Severe fatty change of the graft liver caused by a portosystemic shunt of mesenteric varices', Transpl Int, vol. 15, pp. 259-262.
Senderos, A B, Garcia Sanz, M C & Casanova, M A 2001, 'Portal vein thrombosis. Review of the disease and report of two cases', An Esp Pediatr, vol. 55, no. 6, Dec 2001, pp. 565-8.
Shah, S R, Deshmukh, H L & Mathur, S K 2003, 'Extensive portal and splenic vein thrombosis: Differences in hemodynamics and management', Hepato-Gastroenterology, vol. 50, no. 52, Jul-Aug 2003, pp. 1085-1089.
Shah, S R, Nagral, S S & Mathur, S K 1999, 'Results of a modified sugiura’s devascularisation in the management of “unshuntable” portal hypertension', Hepatic, pancreatic and biliary surgery (HPB Surg), vol. 11, no. 4, pp. 235-9.
Shapiro, R S, Ramos, R, Stancato-Pasik, A, Glajchen, N, Parsons, R, Mitty, H A & H.C., Y 1995, 'Transjugular intrahepatic portosystemic shunt: correlation of portal vein velocity measurements and portosystemic pressure gradients', Computerized Medical Imaging and Graphics, vol. 19, no. 4, July-Aug 1995, pp. 343-350.
Shatford, R A, Nyberg, S L, Meier, S J, White, J G, Payne, W D, Hu, W S & Cerra, F B 1992, 'Hepatocyte function in a hollow fiber bioreactor: a potential bioartificial liver', J Surg Res, vol. 53, pp. 549-557.
Sherlock, S 1974, 'Classification and function aspects of portal hypertension', Am J Surg, vol., no. 127, pp. 121-8.
Sherlock, S 1974, 'Classification and Functional Aspects of Portal Hypertension', The American Journal of Surgery, vol. 127, Feb, pp. 121-128.
Sherlock, S 1978, 'Portal circulation and portal hypertension', GUT, vol. 19, pp. 70-83.
297
Shin, S & Keum, D Y 2002, 'Measurement of blood viscosity using mass-detecting sensor', Biosens Bioelectron, vol. 17, no. 5, May 2002, pp. 383-8.
Shinoka, T 2002, 'Tissue Engineered Heart Valves: Autologous Cell Seeding on Biodegradable Polymer Scaffold', Artificial Organs, vol. 26, no. 5, pp. 402-406.
Shiomi, S, Iwata, Y, Sasaki, N, Morikawa, H, Tamori, A, Habu, D, Takeda, T, Nishiguchi, S, Kuroki, T & Ochi, H 2000, 'Assessment of hepatic blood flow by PET with 150 water: correlation between per-rectal portal scintigraphy with 99Tc(m)-pertechnetate and scintigraphy with 99Tc(m)-GSA', Nucl Med Commun vol. 21, no. 6, June, pp. 533-8.
Shito, M, Tilles, A W, Tompkins, R G, Yarmush, M L & Toner, M 2003, 'Efficacy of an extracorporeal flat-plate bioartificial liver in treating fulminant hepatic failure', Journal of Surgical Research, vol. 111, no. 1, May 2003, pp. 53-62.
Sikuler, E, Kravetz, D & Groszmann, R J 1985, 'Evolution of portal hypertension and mechanisms involved in its maintenance in a rat model', American Journal of Physiology: Gastrointestinal and Liver Physiology, vol. 248, no. 6, pp. 618-G625.
Simonneau, G, Galie, N, Rubin, L J, Langleben, D, Seeger, W, Domenighetti, G, Gibbs, S, Lebrec, D, Speich, R, Beghetti, M, Rich, S & Fishman, A 2004, 'Clinical classification of pulmonary hypertension', J Am Coll Cardiol, vol. 43, no. 12_Suppl_S, June 16, 2004, pp. 5S-12.
Sinoka, T, Imai, Y & Matsumura, G 2000, 'Current status of tissue engineering for cardiovascular structures', J Artif Organs, vol. 3, pp. 102-106.
Sipehia, R, Martucci, G & Lipscombe, J 1996, 'Transplantation of human endothelial cell monolayer on artificial vascular prosthesis: the effect of growth-support surface chemistry, cell seeding density, ECM protein coating, and growth factors', Artif Cells Blood Substit Immobil Biotechnol, vol. 24, no. 1, Jan 1996, pp. 51-63.
Siro, K, Yoshinory, I, Takehisa, I & Junichi, O 2002, 'Numerical Simulations of Blood Flow at the Distal Anastomosis between a Prosthetic Graft and a Native Popliteal Artery' Proceedings of the Secound International Conference on Computational Fluid Dynamics, ICCFD, Sydney, Australia, 15-19 July 2002, 185-190.
Sirose, E, Cote, M F & Doillon, C J 1993, 'Growth factors and biological supports for endothelial cell lining: in vitro study', Int J Artif Organs, vol. 16, no. 8, Aug, pp. 609-19.
Smatr Measurement 2004, Efunda.com. Smith, J J & Kampine, J P 1984, Circulatory Physiology – the essentials /second
edition, Second Edition edn., Williams & Wilkins, Baltimore, MD USA. Smyrniotis, V, Kostopanagiotou, G, Kondi, A, Gamaletsos, E, Theodoraki, K,
Kehagias, D, Mystakidou, K & Contis, J 2002, 'Hemodynamic interaction between portal vein and hepatic artery flow in small-for-size split liver transplantation', Transpl Int vol. 15, pp. 355-360.
298
Society Interventional Radiology 2004, Interventional Treatment for Liver Disease, viewed March 2004, www.sirweb.org/patPub/liverDisease.shtml.
Song, M-H, Sato, M & Ueda, Y 2000, 'Three-Dimensional Simulation of Coronary Artery Bypass Grafting with the Use of Computational Fluid Dynamics', Surg Today, vol. 30, pp. 993-998.
Sovak, M, Soulen, R L & Reichle, F A 1999, 'Blood flow in the human portal vein', Radiology, vol. 99, no. 19th June
pp. 531-536. Spirchez, Z, Stoian, A, Zaharia, T & Badea, R 2003, 'Ultrasound-guided biopsy of
portal vein thrombosis. Value in diagnosis of occult hepatocellular carcinoma in cirrhosis', Ultrasound in Medicine & Biology, vol. 29, Supplement 1, no. 5, May 2003, p. S177.
St Moravec 1987, Strömungsuntersuchungen mit newtonschen und viskoelastischen Flüssigkeiten in Modellen menschlicher Nierenarterien unter Verwendung eines Laser-Doppler-Anemometers, viewed 01/03/2005 http://tumb1.biblio.tu-muenchen.de/publ/diss/mw/2000/weigand.pdf.
Stanley, J C, Burkel, W E, Ford, J W, Vinter, D W, Kahn, R H, Whitehouse, W M J & Graham, L M 1982, 'Enhanced patency of small-diameter, externally supported Dacron iliofemoral grafts seeded with endothelial cells', Surgery, vol. 92, no. 6, Dec 1982, pp. 994-1005.
Starmans-Kool, M J, Stanton, A V, Zhao, S Z, Xu, X Y, Thom, S A & Hughes, A D 2002, 'Measurement of hemodynamics in human carotid artery using ultrasound and computer fluid dynamics', Journal of Applied Physiology, vol. 92, March 2002, pp. 957-961.
Stary, H C, Blankenhorn, D H, Chandler, A B, Glagov, S, Insull, W, Jr., Richardson, M, Rosenfeld, M E, Schaffer, S A, Schwartz, C J, Wagner, W D & Wissler, R W 1992, 'A definition of the intima of human arteries and of its atherosclerosis-prone regions: a report from the Committee on Vascular Lesions of the Council on Arteriosclerosis, American Heart Association
', Circulation, vol. 85, pp. 391-405. Strackee & Westerhof 1993, The physics of heart and circulation, Institute of
Physics Pub, Philadelphia, PA, University of Amsterdam, The Netherlands.
Strandell, T, Erwald, R, Kulling, K G, Lundbergh, P, Marions, O & Wiechel, K L 1973, 'Measurement of dual hepatic blood flow in awake patients', Journal of Applied Physiology, vol. 35, no. 5, pp. 755-761.
Suda, M, Eder, O J, Kunsch, B, Magometschnigg, D & Magometschnigg, H 1993, 'Preoperative assessment and prediction of postoperative results in an artificial arterial network using computer simulation', Computer Methods and Programs in Biomedicine, vol. 41, no. 2, Dec 1993, pp. 77-87.
Sugawara, Y, Yamamoto, J, Higashi, H, Yamasaki, S, Shimada, K, Kosuge, T, Takayama, T & Makuuchi, M 2002, 'Preoperative portal embolization in patients with hepatocellular carcinoma', World J Surg, vol. 26, no. 1, Jan 2002, pp. 105-110.
299
Sugita, S, Ohnishi, K, Saito, M & Okuda, K 1987, 'Splanchnic hemodynamics in portal hypertensive dogs with portal fibrosis', American Journal of Physiology: Gastrointestinal and Liver Physiology, vol. 252, no. 1, pp. 748-G754.
Sugiura, Y 1988, 'A method for analyzing non Newtonian blood viscosity data in low shear rates', Biorheology, vol. 13, pp. 201-218.
Sule, T, Kivanc, C, Sukru, P & Sezai, V 2002, 'Rheological properties of blood in patients with chronic liver disease', Clinical Hemorheology and Microcirculation, vol. 26, pp. 9-14.
Sulica, R, Emree, S & Poom, M 2004, 'Medical management of porto-pulmonary hypertension and right heart failure prior to living-related liver transplantation', Congestive heart failure (Greenwich ,Conn.), vol. 10, no. 4, July, pp. 192-194.
Svoboda, P, Kantorova, I, Brhelova, H, Vasickova, J & Ochmann, J 1997, 'Recent position of transjugular intrahepatic portosystemic shunt in the treatment of portal hypertension', Hepatogastroenterology, vol. 44, no. 15, May-June, pp. 647-655.
Svoboda, P, Kantorova, I, Brhelova, H, Vasickova, J & Ochmann, J 1997, 'Transjugular intrahepatic portosystemic shunt in the treatment of complications in portal hypertension', (in Czech) Rozhl Chir. , vol. 76, no. 2, Feb, pp. 97-102.
Syoten, O 1981, Cardiovascular Hemorheology, First Edition edn., Cambridge University Press, Cambridge (Eng.), New York.
Szilagyi, D E, Whitcomb, J & Schenker, W 1960, 'The laws of fluid flow and arterial grafting', Surgery, vol. 47, pp. 55-67.
Takashinizu, S, Shiraishi, K, Watanabe, N, Numata, M, Kawazoe, K, Miyachi, H, Tokunaga, M, Akatuka, A & Matsuzaki, S 2000, 'Scanning electron microscopic studies on morphological abnormalities of erythrocytes in alcoholic liver diseases', Alcohol Clin Exp Res., vol. 24, no. 4Suppl, April, pp. 81S-86S.
Takayama, Y, Moriura, S, Nagata, J, Akutagawa, A, Hirano, A, Ishiguri, S, Matsumoto, T & Sato, T 2001, 'Embolization of the left portal vein to inferior vena cava shunts for chronic recurrent hepatic encephalopathy via the mesenteric vein', Journal of Gastroenterology and Hepatology, vol. 16, no. 12, Dec 2001, pp. 1425-8.
Tanaka, M & Wanless, I R 1998, 'Pathology of the Liver in Budd-Chiari Syndrome: Portal Vein Thrombosis and the Histogenesis of Veno-Centric Cirrhosis, Veno-Portal Cirrhosis, and Large Regenerative Nodules', Hepatology, vol. 27, no. 2, Feb 1998, pp. 488-496.
Tarazov, P G, Sergeev, V I, Granov, D A, Dolikarpov, A A, Polysalov, V N & Rosenhaus, E V 2002, 'Preoperative embolization of the portal vein in malignant tumors of the liver: preliminary results', Vestn Khir Im I Grek, vol. 161, no. 3, pp. 29-33.
Tasu, J-P, Rocher, L, Peletier, G, Kuoch, V, Kulh, E, Miquel, A, Buffet, C & Biery, M 2002, 'Hepatic Venous Pressure Gradients Measured by Duplex Ultrasound', Clinical Radiology, vol. 57, pp. 746-752.
300
Tayama, E, Hayashida, N, Akasu, K, Kosuga, T, Fukunaga, S, Akashi, H, Kawara, T & Aoyagi, S 2000, 'Biocompatibility of Heparin-Coated Extracorporeal Bypass Circuits: New Heparin Bonded Bioline System', Artificial Organs, vol. 24, no. 8, pp. 618-623.
Taylor, K J, Burns, P N, Woodcock, J P & Wells, P N 1985, 'Blood flow in deep abdominal and pelvic vessels: ultrasonic pulsed-Doppler analysis', Radiology, vol. 154, pp. 487-93.
Tevaearai, H T, Mueller, X M, Jegger, D, Augsburger, M, Stumpe, F & von Segesser, L K 1999, 'Optimization of the pump driven venous return for minimally invasive open heart surgery', Int J Artificial Organs, vol., no. 22, pp. 684-9.
Tevaearai, H T, Mueller, X M, Seigneul, I, Burki, M, Horisberger, J, Stumpe, F & Von Segesser, L K 1999, 'Trillium coating of cardiopulmonary bypass circuits improves biocompatibility', Int J Artif Organs, vol. 22, no. September, pp. 629-34.
The East Development Group, I 2000 Totally Implantable Assist Device (VAD) The Primary Biliary Cirrhosis (PBC) Foundation, U K 2000, UK, ‘Giving support to
PBC Sufferers and their families’, http://www.pbcfoundation.org.uk/. Thilmany, J 2003, 'Computational fluid medicine', Mechanical Engineering, vol.
125, no. 9, Sept 2003, p. 22. Thomford, N R, Sirinek, K R & Martin, E W J 1975, 'A series of 20 successful
Warren shunts', Arch Surg, vol. 110, May 1975, pp. 584-587. Thompson, M A 1981, 'Managing the patient with liver dysfunction', Nursing81,
vol. 11, 11 Nov pp. 101-7. Thomson, A B R & Shaffer, E A 2006, AstraZeneca Canada Inc. Thuluvath, P J, Bal, J S, Mitchell, S, Lund, G & Venbrux, A 2003, 'TIPS for
management of refractory ascites: response and survival are both unpredictable', Dig Dis Sci, vol. 48, no. 3, March 2003, pp. 542-450.
Toso, C, Oberholzer, J, Ceausoglu, I, Ris, F, Rochat, B, Rehor, A, Bucher, P, Wandrey, C, Schuldt, U, Belenger, J, Bosco, D, Morel, P & Hunkeler, D 2003, 'Intra-portal injection of 400-μm microcapsules in a large-animal model', Transpl Int vol. 16, pp. 405-410.
Tsuchiya, Y, Suzuki, S, Inaba, K, Sakaguchi, T, Baba, S, Miwa, M, Konno, H & Nakamura, S 2003, 'Impact of endothelin-1 on microcirculatory disturbance after partial hepatectomy under ischemia/reperfusion in thioacetamide-induced cirrhotic rats', J Surg Res, vol. 111, no. 1, 1 May 2003, pp. 100-108.
Tunstall, A, Eberhart, R C & Prager, M D 1995, 'Endothelial cells on Dacron vascular prostheses: adherence, growth, and susceptibility to neutrophils', J Biomed Mater Res, vol. 29, no. 10, Oct, pp. 1193-9.
Tygstrup, N, Winkler, K, Mellemgaard, K & Andreassen, M 1962, 'Determination of the hepatic arterial blood flow and oxygen supply in man clamping the hepatic artery during surgery', The journal of clinical investigation, vol. 41, no. 3, March, pp. 447-454.
Ueda, H, Kitani, K, Kameda, H, Takeda, T, Chiba, K, Nagatani, M, Yamada, H & Iio, M 1971, 'Splenic blood flow in idiopathic portal hypertension in Japan
301
measured by 85 Kr clearance method', Acta Hepato-splenol, vol., no. 18, pp. 28-40.
Uflacker, R 2003, 'Applications of percutaneous mechanical thrombectomy in transjugular intrahepatic portosystemic shunt and portal vein thrombosis', Tech Vasc Interv Radiol, vol. 6, no. 1, March 2003, pp. 59-69.
Undar, A & Vaughn, W K 2002, 'Effects of mild hypothermic cardiopulmonary bypass on blood viscoelasticity in coronary artery bypass grafting patients', Artificial Organs, vol. 26, pp. 964-966.
University Of Cincinnati 1994, The Mont Reid Surgical Handbook, Forth Edition edn., The University of Cincinnati Residents, Cincinnati.
Urban, C H, Deutschmann, A, Kerbl, R, Lackner, H, Schwinger, W, Konigsrainer, A & Margreiter, R 2002, 'Organ tolerance following cadaveric liver transplantation for chronic graft-versus-host disease after allogeneic bone marrow transplantation', Bone Marrow Transplantation, vol. 30, no. 8, 15 Oct 2002, p. 535.
USA 2003, National Vital Statistics Reports, cat. no., p. Vol 52 N 9. Viewed Van Kooten, T G & Von Recum, A F 1999, 'Cell adhesion to textures silicone
surfaces: the influence of time of adhesion and texture on focal contact and fibronectin fibrin formation', Tissue Eng, vol. 5, no. 3, June, pp. 223-240.
Van Splunder, I P, Huisman, T W, Stijnen, T & Wladimiroff, J W 1994, 'Presence of pulsations and reproducibility of waveform recording in the umbilical and left portal vein in normal pregnancies', Ultrasound Obstet Gynecol, vol. 4, no. 1, 1 Jan 1994, pp. 49-53.
Van Wachem, P B, Beugeling, T, Mallens, B W, Dekker, A, Feijen, J, Bantjes, A & van Aken, W G 1988, 'Deposition of endothelial fibronectin on polymeric surfaces', Biomaterials, vol. 9, no. 1, Jan 1988, pp. 121-3.
Varghese, S S & Frankel, S H 2003, 'Numerical Modeling of Pulsatile Turbulent Flow in Stenotic Vessels', Journal of Biomechanical Engineering, vol. 125, no. 4, Aug 2003, pp. 445-460.
Versweyveld, L 1997, 'Smart Blood Micro-pump improve cirrhotic liver perfusion', Virtual Medical World – monthly, Brussels, vol., 26 November.
Viallet, A, Joly, J G, Marleau, D & Lavoie, P 1970, 'Comparison of free portal venous pressure and wedged hepatic venous pressure in patients with cirrhosis of the liver ', Gastroenterology, vol. 59, no. 3, Sept, pp. 372-375.
Virtual Hospital 2003, Virtual Hospital: Atlas of Liver Pathology: Vascular Alteration in the Liver, viewed 16 Sept 2003, www.vh.org/
Vorobioff, J, Bredfeldt, J E & Groszmann, R J 1983, 'Hyperdynamic circulation in portal-hypertensive rat model: a primary factor for maintenance of chronic portal hypertension', American Journal of Physiology: Gastrointestinal and Liver Physiology, vol. 244, no. 1, pp. 52-G56.
Walburn, F J & Schneck, D J 1976, 'A constitutive equation for whole human blood', Biorheology, vol. 13, pp. 201-218.
Waller, J, Brook, N R & Nicholson, M L 2003, 'Cardiac allograft vasculopathy: current concepts and treatment', Transpl Int, vol. 16, no. 6, Jun 2003, pp. 367-75. Viewed Epub 2003 May 17,
302
Walser, E M, DeLa Pena, R, Villanueva-Meyer, J, Ozkan, O & Soloway, R 2000, 'Hepatic Perfusion before and after the Transjugular intrahepatic Portosystemic Shunt Procedure: Impact on Survival', Journal of Vascular and Interventional Radiology, vol. 11, pp. 913-918.
Wampler, R K, Moise, J C, Frazier, O H & Olsen, D B 1988, 'In vivo evaluation of a peripheral vascular access axial flow blood pump', Trans Am Soc Artif Intern Organs, vol. 34, p. 450.
Wang, X & Stolz, J F 1994, 'Characterization of pathological blood with a new rheological relationship', Clinical Hemorheology, vol. 13, pp. 237-245.
Weiskircher, R, Jens, K, Weiskirchen, S, van de Leur, E, Dagmar, K & Gressner, A M 2000, 'Comparative evaluation of gene delivery devices in primary cultures of rat hepatic stellate cells and rat myofibroblasts ', BMC Cell Biology, vol. 1, p. 4.
Wexler, M J & MacLean, L D 1975, 'Massive spontaneous portal-systemic shunting without varices', Arch Surg, vol. 110, pp. 996-1003.
Worman, H J 1995, Cirrhosis, viewed 2002, http://cpmcnet.columbia.edu/dept/gi/cirrhosis.html.
Worman, H J 1999, The Liver Disorders sourcebook, Lowell House, LA. Wright, K C, Wallace, S, Charnsangavej, C, Carrasco, C H & Gianturco, C 1985,
'Percutaneous endovascular stents: an experimental evaluation', Radiology, vol. 156, no. 69-72.
Xu, X Y, Long, Q, Collins, M W, Bourne, M & Griffith, T M 1999, 'Reconstruction of blood flow patterns in human arteries', Proceedings of the Institution of Mechanical Engineers Part H- Journal of Engineering in Medicine, vol. 213, pp. 411-421.
Yamaguchi, T, Hayasaka, T & Mori, D 2002, 'Towards Computational Biomechanics Based Cardiovascular Medical Practice' Riken Symposium, 46-61.
Yambe, T, Owada, N, Kobayashi, S-I, Tanaka, A, Yoshizawa, M, Abe, K-I, Tabayashi, K, Takeda, H, Hashimoto, H & Nitta, S-I 2000, 'Totally Implantable Ventricular Assist System that Can Increase Brain Blood Flow', Artificial Organs, vol. 24, no. 8, pp. 644-647.
Yang, T H, Miyoshi, H & Ohshima, N 2001, 'Novel cell immobilization method utilizing centrifugal force to achieve high-density hepatocyte culture in porous scaffold', Inc. J Biomed Mater Res, vol. 55, pp. 379-386.
Yang, W, Koti, R, Glantzounis, G, Davidson, B R & Seifalian, A M 2003, 'Arterialization of the portal vein improves hepatic microcirculation and tissue oxygenation in experimental cirrhosis', Br J Surg, vol. 90, no. 10, Oct 2003, pp. 1232-9.
Yang, Y, Harvey, S A K & Gandhi, C R 2003, 'Kupffer cells are a major source of platelet-activating factor in the CCl4-induced cirrhotic rat liver', J Hepatol, vol. 39, pp. 200-207.
Yedavalli, R V, Loth, F, Yardimci, A, Pritchard, W F, Oshinski, J N, Sadler, L, Charbel, F & Alperin, N 2001, 'Construction of a Physical Model of the Human Carotid Artery Based Upon In Vivo Magnetic Resonance Images', Journal of Biomechanical Engineering, vol. 123, no. 4, Aug 2001, p. 372.
303
Youk, J H, Lee, J M & Kim, C S 2003, 'Therapeutic Response Evaluation of Malignant Hepatic Masses Treated by Interventional Procedures With Contrast-Enhanced Agent Detection Imaging', J Ultrasound Med, vol. 22, no. 9, Sep 2003, pp. 911-20.
Yousef, K, Yang, W, Mazumber, J & Morsi, Y 2001, (ed Technology, S U o) www.swin.edu.au/iris/pdf/profiles.KhaledYousif.pdf.
Yu, J S, Kim, K W, Park, M S & Yoon, S W 2001, 'Bile duct injuries leading to portal vein obliteration after transcatheter arterial chemoembolization in the liver: CT findings and initial observations', Radiology, vol. 221, no. 2, Nov 2001, pp. 429-36.
Žáček, M & Krause, E 1996, ' Numerical simulation of the blood flow in the human cardiovascular system', Journal of Biomechanics, vol. 29, no. 1, pp. 13-20.
Zardi, E M, Picardi, A, Ambrosino, G, Serio, A, Costantino, S & Dobrina, A 2003, 'Hemodynamic effects of a prostacyclin analog (iloprost) on portal flow velocity and volume and visceral artery circulation in patients with lower limb arteriopathy', Ultrasound in Medicine & Biology, vol. 29, no. 2, Feb 2003, pp. 209-213.
Zemel, G, Katzen, B T, Becker, G J, Benenati, G F & Sallee, D S 1991, 'Percutaneous transjugular portosystemic shunt', JAMA, The Journal of the American Medical Association, vol. 226, no. 3, 17 July 1991, p. 390(4).
Zeng, H, Fei, D-Y, Fu, C-T & Kraft, K A 2003, 'Internet (WWW) based system of ultrasonic image processing tools for remote image analysis', Computer Methods and Programs in Biomedicine, vol. 71, no. 3, July 2003, pp. 235-241.
Zent, T, Rossle, M, Bertz, H, Siegerstetter, V, Ochs, A & Finke, J 2001, 'Severe veno-occlusive disease after allogeneic bone marrow or peripheral stem cell transplantation – role of transjugular intrahepatic portosystemic shunt (TIPS)', Liver, vol. 21, pp. 31-36.
Zhang, J-B & Kuang, Z-B 2000, 'Study on blood constitutive parameters in different blood constitutive equations', Journal of Biomechanics, vol. 33, pp. 355-360.
Zhao, S Z, Xu, X Y, Hughes, A D, Thom, A S, Stanton, A V, Ariff, B & Long, Q 2000, 'Blood flow and vessel mechanics in a physiologically realistic model of a human carotid arterial bifurcation', Journal of Biomechanics, vol. 33, pp. 975-984.
Zilla, P 1991, 'Endothelialization of vascular grafts', Current Opinion in Cardiology, vol. 6, no. 6, Dec 1991, pp. 877-886.
Žižka, J, Eliáš, P, Krajina, A, Michl, A, Lojík, M, Ryška, P, Mašková, J, Hůek, P, Šafka, V, Vaňásek, T & Bukač, J 2000, 'Value of Doppler Sonography in Revealing Transjugular Intrahepatic Portosystemic Shunt Malfunction', American Journal of Roentgenology, vol. 175, pp. 141-148.
304
Appendix 1 MIXTURE MODEL – FOUR PHASES
Material: blood-alike (fluid)
Property Units Method Value(s) ---------------------------------------------------------------------------------------------- Density (kg/m3) constant 1070 Cp (Specific Heat) (j/kg-k) constant 4182 Thermal Conductivity (w/m-k) constant 0.6 Viscosity (kg/m-s) non-Newtonian (0.2073 0.4851 310
-power-law 0.00125 0.003) Molecular Weight (kg/kgmol) constant 18.0152 Standard State Enthalpy (j/kgmol) constant 0 Reference Temperature (k) constant 298.15 FLUENT Version: 3d, segregated, mixture, lam (3D, segregated, Mixture, laminar) Release: 6.1.22 Model Settings ------------------------------------- Space 3D Time Steady Viscous Laminar Heat Transfer Enabled Solidification and Melting Disabled Radiation None Species Transport Disabled Coupled Dispersed Phase Disabled Pollutants Disabled Soot Disabled
Boundary Conditions Zones Name ID Type -------------------------------------- fluid 2 fluid wall 3 wall outlet4 4 outflow outlet3 5 outflow outlet2 6 outflow outlet1 7 outflow inlet 8 velocity-inlet default-interior 10 interior
i
Boundary Conditions of Fluid
Condition Value ------------------------------------------------------------------------------------- Material Name blood-alike Specify source terms? no Source Terms ((x-momentum (inactive . #f) (constant . 0)
(profile )) (y-momentum (inactive . #f) (constant . 0) (profile )) (z-momentum (inactive . #f) (constant . 0) (profile )) (energy (inactive . #f) (constant . 0) (profile )))
Specify fixed values? no Local Coordinate System for Fixed Velocities no Fixed Values ((x-velocity (inactive . #f) (constant . 0) (profile ))
(y-velocity (inactive . #f) (constant . 0) (profile )) (z-velocity (inactive . #f) (constant . 0) (profile )) (temperature (inactive . #f) (constant .0) (profile)))
Motion Type 0 X-Velocity Of Zone 0 Y-Velocity Of Zone 0 Z-Velocity Of Zone 0 Rotation speed 0 X-Origin of Rotation-Axis 0 Y-Origin of Rotation-Axis 0 Z-Origin of Rotation-Axis 0 X-Component of Rotation-Axis 0 Y-Component of Rotation-Axis 0 Z-Component of Rotation-Axis 1 Deactivated Thread no Porous zone? no Porosity 1 Solid Material Name aluminum
Boundary conditions at the Wall
Condition Value ------------------------------------------------------------------------------------ Wall Thickness 0 Heat Generation Rate 0 Material Name aluminum Thermal BC Type 1 Temperature 300 Heat Flux 0 Convective Heat Transfer Coefficient 0 Free Stream Temperature 300 Enable shell conduction? no Wall Motion 0 Shear Boundary Condition 0 Define wall motion relative to adjacent cell zone? yes
ii
Apply a rotational velocity to this wall? no Velocity Magnitude 0 X-Component of Wall Translation 1 Y-Component of Wall Translation 0 Z-Component of Wall Translation 0 Define wall velocity components? no X-Component of Wall Translation 0 Y-Component of Wall Translation 0 Z-Component of Wall Translation 0 External Emissivity 1 External Radiation Temperature 300 Rotation Speed 0 X-Position of Rotation-Axis Origin 0 Y-Position of Rotation-Axis Origin 0 Z-Position of Rotation-Axis Origin 0 X-Component of Rotation-Axis Direction 0 Y-Component of Rotation-Axis Direction 0 Z-Component of Rotation-Axis Direction 1 X-component of shear stress 0 Y-component of shear stress 0 Z-component of shear stress 0 Surface tension gradient 0 Outlet4 Condition Value --------------------------- Flow rate weighting 0.35 Outlet3 Condition Value --------------------------- Flow rate weighting 0.35 Outlet2 Condition Value --------------------------- Flow rate weighting 0.15 Outlet1 Condition Value --------------------------- Flow rate weighting 0.15 Inlet Condition Value ------------------------------------------- Temperature 310 Is zone used in mixing-plane model? no
iii
Default-interior Condition Value Material Properties
Material: blood_cells_third (fluid) Property Units Method Value(s) --------------------------------------------------------------------------------------------------- Density kg/m3 constant 2500 Cp (Specific Heat) j/kg-k constant 4182 Thermal Conductivity w/m-k constant 0.60000002 Viscosity kg/m-s non-Newtonian-power-law (0.2073 0.4851 310
0.00125 0.003) Molecular Weight kg/kgmol constant 18.0152 Standard State Enthalpy j/kgmol constant 0 Reference Temperature k constant 298.14999 L-J Characteristic Length angstrom constant 3.711 L-J Energy Parameter k constant 78.6 Thermal Expansion Coefficient 1/k constant 0 Degrees of Freedom constant 0
Material: blood_cells_second (fluid)
Property Units Method Value(s) ---------------------------------------------------------------------------------------------- Density kg/m3 constant 2050 Cp (Specific Heat) j/kg-k constant 4182 Thermal Conductivity w/m-k constant 0.60000002 Viscosity kg/m-s non-Newtonian-power-law (0.2073 0.4851 310
0.00125 0.003) Molecular Weight kg/kgmol constant 18.0152 Standard State Enthalpy j/kgmol constant 0 Reference Temperature k constant 298.14999 L-J Characteristic Length angstrom constant 3.711 L-J Energy Parameter k constant 78.6 Thermal Expansion Coefficient 1/k constant 0 Degrees of Freedom constant 0
Material: blood_cells_first (fluid) Property Units Method Value(s) ------------------------------------------------------------------------------------------------- Density kg/m3 constant 2000 Cp (Specific Heat) j/kg-k constant 4182 Thermal Conductivity w/m-k constant 0.60000002 Viscosity kg/m-s non-Newtonian-power-law (0.2073 0.4851 310
0.00125 0.003 ) Molecular Weight kg/kgmol constant 18.0152 Standard State Enthalpy j/kgmol constant 0 Reference Temperature k constant 298.14999 L-J Characteristic Length angstrom constant 3.711 L-J Energy Parameter k constant 78.6 Thermal Expansion Coefficient 1/k constant 0 Degrees of Freedom constant 0
iv
Material: plasma_blood (fluid)
Property Units Method Value(s) ------------------------------------------------------------------- Density kg/m3 constant 1010 Cp (Specific Heat) j/kg-k constant 4182 Thermal Conductivity w/m-k constant 0.60000002 Viscosity kg/m-s constant 1.7894001e-05 Molecular Weight kg/kgmol constant 18.0152 Standard State Enthalpy j/kgmol constant 0 Reference Temperature k constant 298.14999 L-J Characteristic Length angstrom constant 3.711 L-J Energy Parameter k constant 78.6 Thermal Expansion Coefficient 1/k constant 0 Degrees of Freedom constant 0
Material: aluminum (solid)
Property Units Method Value(s) ---------------------------------------------------------------------------------------- Density kg/m3 constant 2719 Cp (Specific Heat) j/kg-k constant 871 Thermal Conductivity w/m-k constant 202.4
Material: air (fluid)
Property Units Method Value(s) ---------------------------------------------------------------------------------------- Density kg/m3 constant 1.225 Cp (Specific Heat) j/kg-k constant 1006.43 Thermal Conductivity w/m-k constant 0.0242 Viscosity kg/m-s constant 1.7894e-05 Molecular Weight kg/kgmol constant 28.966 Standard State Enthalpy j/kgmol constant 0 Reference Temperature k constant 298.15 L-J Characteristic Length angstrom constant 3.711 L-J Energy Parameter k constant 78.6 Thermal Expansion Coefficient 1/k constant 0 Degrees of Freedom constant 0
v
Contours of Static Pressure (Pascal) for the mixture
vi
The scales used for all velocity visualisations are identical to show the differences between the phases
Velocity Vectors Coloured by Velocity Magnitude (m/s) of the mixture
Phase 1 plasma Velocity Coloured by Velocity Magnitude (m/s) in the mixture
vii
Phase 2 erythrocytes Coloured by Velocity Magnitude (m/s) in the mixture
Phase 3 Leukocytes Velocity Coloured by Velocity Magnitude (m/s) in the mixture
viii
Phase 4 Platelets Coloured by Velocity Magnitude (m/s) in the mixture VELOCITY/VELOCITY MAGNITUDE HISTOGRAM MIXTURE 0 cells below 0.0011 (0 %) 3603 cells between 0.0011 and 0.021 (6.57 %) 9954 cells between 0.021 and 0.041 (18.17 %) 10240 cells between 0.041 and 0.061 (18.69 %) 10846 cells between 0.061 and 0.081 (19.8 %) 10716 cells between 0.081 and 0.10 (19.56 %) 6157 cells between 0.10 and 0.121 (11.24 %) 2042 cells between 0.12 and 0.14 (3.7 %) 680 cells between 0.14 and 0.16 (1.24 %) 426 cells between 0.16 and 0.18 (0.777 %) 117 cells between 0.18 and 0.20 (0.21 %) 1 cells above 0.201 (0.0018 %) PLASMA 0 cells below 0 (0 %) 3401 cells between 0 and 0.02 (6.2 %) 9705 cells between 0.02 and 0.04 (17.7 %) 10397 cells between 0.04 and 0.06 (18.97 %) 10699 cells between 0.06 and 0.08 (19.5 %) 11011 cells between 0.08 and 0.1 (20.1 %) 6480 cells between 0.1 and 0.12 (11.8 %) 1873 cells between 0.12 and 0.14 (3.4 %) 667 cells between 0.14 and 0.16 (1.2 %) 425 cells between 0.16 and 0.18 (0.77 %) 123 cells between 0.18 and 0.2 (0.22 %) 1 cells above 0.2 (0.0018 %)
ix
ERYTROCYTES 0 cells below 0 (0 %) 3315 cells between 0 and 0.02 (6.05 %) 9780 cells between 0.02 and 0.04 (17.8 %) 10392 cells between 0.04 and 0.06 (18.97 %) 10699 cells between 0.06 and 0.08 (19.5 %) 11010 cells between 0.08 and 0.1 (20.1 %) 6489 cells between 0.1 and 0.12 (11.8 %) 1889 cells between 0.12 and 0.14 (3.45 %) 663 cells between 0.14 and 0.16 (1.2 %) 423 cells between 0.16 and 0.18 (0.77 %) 121 cells between 0.18 and 0.2 (0.22 %) 1 cells above 0.2 (0.0018 %) LEUKOCYTES 0 cells below 0 (0 %) 18997 cells between 0 and 0.015 (34.67 %) 3004 cells between 0.015 and 0.0316 (5.5 %) 4095 cells between 0.0316 and 0.047 (7.47 %) 5527 cells between 0.047 and 0.063 (10.09 %) 6452 cells between 0.063 and 0.079 (11.77 %) 5578 cells between 0.079 and 0.095 (10.18 %) 5926 cells between 0.095 and 0.11 (10.8 %) 3910 cells between 0.11 and 0.126 (7.1 %) 1193 cells between 0.126 and 0.14 (2.17 %) 99 cells between 0.14 and 0.158 (0.18 %) 1 cells above 0.158 (0.0018 %) PLATELETS 0 cells below 0 (0 %) 3305 cells between 0 and 0.02 (6.03 %) 9735 cells between 0.02 and 0.04 (17.77 %) 10456 cells between 0.04 and 0.06 (19.08 %) 10689 cells between 0.06 and 0.08 (19.5 %) 11044 cells between 0.08 and 0.1 (20.16 %) 6472 cells between 0.1 and 0.12 (11.8 %) 1865 cells between 0.12 and 0.14 (3.4 %) 666 cells between 0.14 and 0.16 (1.2 %) 426 cells between 0.16 and 0.18 (0.78 %) 123 cells between 0.18 and 0.2 (0.22 %) 1 cells above 0.2 (0.0018 %) Blood-alike (mixture) Density kg/m3 = 1070 Plasma Density =1010, constant viscosity Erythrocytes Density=2000 Leukocytes Density=2050 Platelets Density=2500 All cells are described via non-Newtonian power law
x
Appendix 2
This appendix contains the technical drawings used for the manufacturing of the
Bioreactor and some photographic images not included in the body of this thesis.
Figures 1-4 drawn by Peter Robb, Engineering workshop at Swinburne University of
technology, based on the design provided by the author of this thesis.
Figure 1 Lower ring (outside) (size is accurate if the image is enlarged to fit
properly into A4 size paper)
xi
Figure 2 Upper ring
xii
Figure 3 Inner hub (cylinder) with holes
Figure 4 Inner hub (cylinder)
xiii
Figure 5 Inner cylinder used in the bioreactor
Figure 6 Inner hub attaching holes view with one spike
xiv
Figure 7 Both outer rings - lower in front, upper in the back
xv
Figure 8 Both rings assembled bottom view
xvi
Appendix 3
Bioreactor glass sections
Figure 1Smooth angle of fluid inlet
Figure 2 Lid rough area to connect to the body andf prevent slipping
xvii
Appendix 4
If needed, each particle in each phase can be tracked separately, and below is an example of five different particle streams, all of which are leukocytes.
Figure 6.6.3.5. Single particle stream ID 1
Figure 6.6.3.6. Single particle stream ID 5
xviii
Figure 6.6.3.7. Single particle stream ID 7
Figure 6.6.3.8. Single particle stream ID 10
xix
Figure 6.6.3.9. Single particle stream ID 31
xx
Appendix 5
Normal Geometry points measured by LDA 10 20
run Ch1 Vel
Mean Point X Y Vel (m/s)
X Vel (m/s)
Mean Velocity
Point
Y Vel (m/s)
X Vel (m/s)
Mean Velocity
5 -
0.00278 9 -3 0.0117 0.000652 0.01176 19 -
0.00189 0.00037 0.001929 4 0.0011 7 -2 0.02358 0.000556 0.023589 17 0.00983 0.00078 0.009869 1 0.0314 1 0 0.03147 0.001688 0.03152 11 0.02877 0.00077 0.028785 2 0.00168 3 2 0.00985 0.0011 0.00991 13 0.0203 0.0010 0.020337 3 0.0098 5 3 -0.002 -0.016 0.01677 15 0.0099 0.0008 0.009945
10 0.0006 9 0.0117 6 -0.016 7 0.023 8 0.0005
15 0.0099 14 0.0010 11 0.0287 12 0.0007 13 0.0203 21 0.0086 20 0.0003 19 -0.0018 16 0.0008 17 0.0098 18 0.0007
32 52
Mean Velocity Point
Y Vel (m/s)
X Vel (m/s)
Mean Velocity Point
X points
Y Vel (m/s) X Vel (m/s)
Mean Velocity
0.00192 29 0.01040 0.00187 0.01057 41 -17 0.00066 0.00816 0.0082 0.00987 27 0.01206 0.0039 0.01268 39 -14 -0.0128 0.00178 0.013 0.02878 21 0.0086 0.0083 0.0120 37 -11 -0.025 0.00527 0.025
0.0203 23 0.0215 0.009 0.0234 31 11 0.0024 0.0022 0.0032 0.00994 25 0.0216 0.0083 0.023 33 14 0.029 0.0140 0.0323
35 17 0.020 3.42E-05 0.0200
Table 1 LDA measured points Normal geometry
xxi
Points measured with LDA in obstructed model
point X Y Z X mean velocity Y mean velocity Mean Velocity
5 -3 3 0 -0.000793489 2.75E-05 0.000793965 1 0 3 0 -0.00144936 6.12E-05 0.001450651 3 3 3 0 -0.000675162 7.57E-05 0.000679396
11 -2 4 0 0.000101825 0.00094458 0.000950052 7 0 4 0 -0.000975743 0.000609822 0.001150633 9 3 4 0 -0.000547796 0.00156395 0.001657112
17 -4 5 0 -0.000268072 0.000702149 0.000751582 13 0 5 0 -0.00106129 0.000273288 0.001095912 15 4 5 0 -0.000551111 0.000180593 0.000579946 23 -1 9 0 -0.00131735 0.000310778 0.001353512 19 0 9 0 0.000344948 0.000177235 0.000387816 21 3 9 0 -0.000471501 0.000434746 0.00064134 25 0 15 0 -0.000840722 0.000671132 0.001075747 27 3 15 0 -0.000554916 0.000367572 0.000665613 29 4 15 0 -0.000619239 -0.000164478 0.000640711 41 -13 26 0 0.000487787 -0.0007378 0.000884469 39 -9 26 0 -0.000573224 0.000313359 0.000653284 37 -5 26 0 -0.000679411 2.11E-05 0.000679738 31 5 26 0 -0.00051563 -1.80E-05 0.000515945 33 7 26 0 -0.000293929 3.37E-05 0.00029585 35 9 26 0 -0.000603702 -0.00156045 0.001673159
53 -
20.5 30 0 -0.000137471 0.000276492 0.000308782
51 -
14.5 30 0 0.000595904 0.00E+00 0.000595904
49 -
10.5 30 0 -0.000378472 -0.000546754 0.000664967 43 11.5 30 0 -0.000282733 4.69E-05 0.000286593 45 15.5 30 0 0.00604625 0.00403746 0.007270366 47 19.5 30 0 -0.00095815 -0.000314298 0.001008382 61 -21 34 0 -0.00061384 0.000167746 0.000636348 59 -19 34 0 -0.000495811 -0.00036324 0.000614631 57 -10 34 0 -0.000723606 0.000291316 0.000780045 55 -8 34 0 -0.000130783 0.000273278 0.00030296 63 10 34 0 0.000577789 -0.00120348 0.001334992 65 14 34 0 -0.000109826 -6.15E-05 0.000125875 67 26 34 0 0.00147434 5.80E-06 0.001474351 69 12 43 0 -0.000343138 0.000316319 0.000466692 71 10 43 0 -0.000535447 0.000388657 0.000661633
Table 2 LDA measured in obstructed model
xxii
y=10mm y=20mm y=32mm y=60mm y=75mm Radii Velocity Radii Velocity Radii Velocity Radii Velocity Radii Velocity
-5 0.000463 -5 0.000447 -5 0.000418 -5 0.001543 -17 0.018009 -4 0.053382 -4 0.050152 -4 0.049199 -4 0.023506 -15.5 0.039071 -3 0.086575 -3 0.085546 -3 0.085013 -3 0.046672 -14 0.05087 -2 0.105447 -2 0.110145 -2 0.111286 -2 0.070364 -12.5 0.05683 -1 0.113623 -1 0.123493 -1 0.126552 -1 0.087567 -11 0.056974 0 0.115575 0 0.12726 0 0.131072 0 0.102868 -9.5 0.049796 1 0.112838 1 0.122511 1 0.125554 1 0.114841 -8 0.025123 2 0.102658 2 0.107363 2 0.10879 2 0.115179 8 0.062681 3 0.080845 3 0.080161 3 0.080005 3 0.097294 9.5 0.105131 4 0.046309 4 0.043387 4 0.042674 4 0.059312 11 0.123706 5 0.000359 5 0.00034 5 0.00034 5 0.00461 12.5 0.126592
14 0.113425 15.5 0.092408 17 0.036115
Table 3 Normal geometry CFD simulation points
20 21 22 26 32 37 Radii
Velocity
Radii
Velocity
Radii
Velocity
Radii
Velocity
Radii
Velocity
Radii
Velocity
-5 3.97E-05
-5 4.37E-05
-5 0.00026
-5 0.000998
-5 0.000374
-5 0.000752
-4 0.054544
-4 0.054679
-4 0.055173
-4 0.051803
-4 0.057485
-4 0.063589
-3 0.084409
-3 0.083001
-3 0.085158
-3 0.084283
-3 0.086618
-3 0.094256
-2 0.098176
-2 0.098395
-2 0.097738
-2 0.100241
-2 0.10149
-2 0.10698
-1 0.10412
-1 0.105478
-1 0.106293
-1 0.107055
-1 0.108115
-1 0.112791
0 0.106575
0 0.107422
0 0.107748
0 0.108902
0 0.109864
0 0.111843
1 0.10502
1 0.105692
1 0.105717
1 0.106664
1 0.1069 1 0.103872
2 0.098064
2 0.09843
2 0.098738
2 0.09922
2 0.098051
2 0.086476
3 0.084796
3 0.084192
3 0.082868
3 0.081547
3 0.080043
3 0.062689
4 0.054229
4 0.055257
4 0.055642
4 0.054799
4 0.050515
4 0.031791
5 0.000576
5 0.001163
5 0.000589
5 0.000504
5 6.79E-05
5 0.000212
Table 4a CFD points for obstructed geometry
xxiii
42.5 48 61 67 74 Radii Velocity Radii Velocity Radii Velocity Radii Velocity Radii Velocity
-5 0.00063 -5 0.00084 -7.5 0.0016 -6.5 0.044 -16.5 0.003 -4 0.083 -4 0.07307 -6 0.0153 -5.5 0.064 -15.5 0.0198 -3 0.110 -3 0.114 -4.5 0.042 -4.5 0.0689 -13 0.041 -2 0.126 -2 0.127 -3 0.069 2.5 0.0196 -11.5 0.053 -1 0.134 -1 0.13 0 0.0946 3.5 0.064 -10 0.062 0 0.127 0 0.124 1.5 0.084 8.5 0.090 -8.5 0.0577 1 0.018 1 0.0998 3 0.10 9.5 0.187 -7 0.030
2 0.061 4.5 0.10 10.5 0.157 7 0.00015 3 0.029 6 0.089 11.5 0.0268 8 0.037 4 0.007 7.5 0.048 10 0.098 5 0 11.5 0.143 12.5 0.15 13.5 0.146 14.5 0.11 15.5 0.068 16.5 0.011
Table 4b CFD points for obstructed geometry
xxiv
Publications resulting from this research 1. Petkova, S, Hossain, A, Naser, J & Palombo, E 2003, 'CFD modelling of blood flow in the portal vein hypertension with and without thrombosis' Third International Conference on CFD in the minerals and Process Industries, CSIRO, Melbourne, Australia, 10-12 Dec 2003, 527-530 2. Morsi, Y, Das, S & Petkova, S 2001, ‘Analysis of flow field in a T-bifurcation method’ In the proceeding of the first Asian-Pacific Congress on Computational Mechanics, Sydney, Australia, 20-23 Nov 2001
In process of reviewing 1. Petkova, S, Hossain, A, Naser, J & Palombo, E, ‘Particle tracking of blood cells using FLUENT’ 2. Petkova, S, Palombo, E & Robb, P, ‘Simple Versatile Easy Tissue-culture in the Laboratory Apparatus: Bioreactor for blood vessel growth in vitro’
xxv
![Antibiotic prophylaxis for cirrhotic patients with upper ...[Intervention Review] Antibiotic prophylaxis for cirrhotic patients with upper gastrointestinal bleeding Norberto C Chavez-Tapia1,](https://img.pdfslide.net/doc/110x75/5e937e618bf0364d7d5b6953/antibiotic-prophylaxis-for-cirrhotic-patients-with-upper-intervention-review.jpg)
