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A Peptide Modified Hydrogel Therapy for Acute Myocardial Infarction
by
Lewis A. Reis
A thesis submitted in conformity with the requirements for the degree of Doctor o Philosophy
The Institute for Biomaterials and Biomedical Engineering
University of Toronto
© Copyright by Lewis Reis 2015
ii
A Peptide Modified Hydrogel Therapy for Acute Myocardial Infarction
Lewis A. Reis
Doctor of Philosophy
The Institute of Biomaterials and Biomedical Engineering
University of Toronto
2015
Abstract
Myocardial infarction (MI) results in the death of cardiomyocytes (CM) followed by scar formation
and pathological remodeling of the heart. We postulate that immobilization of the pro-survival
angiopoietin-1-derived peptide, QHREDGS, to a chitosan-collagen hydrogel could produce a clinically
translatable thermo-responsive hydrogel to attenuate post-MI cardiac remodeling. Conjugation of
QHREDGS peptide to chitosan does not interfere with gelation, structure, or mechanical properties of
chitosan-collagen hydrogel blends. The storage modulus of 2.5mg/mL 1:1 mass:mass chitosan:collagen
was measured to be 54.9±9.1 Pa and loss modulus of 6.1±0.9 Pa. Dose response of the QHREDGS
peptide was assessed and it was found that CMs encapsulated in High peptide gel (651±8 nmol
peptide/mL-gel) showed improved morphology, viability, and metabolic activity in comparison to the
Low peptide (100±30 nmol peptide/mL-gel) and Control (No Peptide) groups. Construct (CMs in
hydrogel) functional properties were not significantly different between the groups, however success rate
of obtaining a beating construct was improved in the hydrogel with the High amount of QHREDGS
iii
peptide immobilized compared to the Low and Control groups. An adult Lewis rat left anterior
descending coronary artery ligation procedure to mimic acute MI model was used to assess in vivo
hydrogel performance. QHREDGS-conjugated hydrogel (QHG213H), Control gel, or PBS was injected
into 3 locations in the MI zone. By in vivo tracking and chitosan staining, the hydrogel was demonstrated
to remain in situ for 2 weeks and cleared in about 3 weeks. By echocardiography and pressure-volume
analysis, the QHG213H hydrogel significantly improved cardiac function compared to the controls. Scar
thickness and scar area fraction were also significantly improved with QHG213H gel injection compared
to the controls. Mechanistically, there were significantly more cardiomyocytes, determined by cardiac
troponin-T staining, in the MI zone of the QHG213H hydrogel group. No significant difference in
inflammatory response between groups was observed as determined by gene regulation and cytokine
analysis of excised heart sections 24 hours after treatment. The interaction of CMs with QHREDGS was
found to be mediated by β1-integrins and to increase expression of the pro-survival effector MAPK.
iv
Acknowledgements
First, I must thank my supervisor and mentor Dr. Milica Radisic for initially giving me the
opportunity to work in her lab, and for the years of support and guidance. I hope the successful
completion of this thesis and the project which we have worked on over the last number of years
is testament to that.
My committee members, Drs. Julie Audet and Molly Shoichet have scrutinized this work
many times and given invaluable advice on making it the best it can be, for which I am grateful.
The help, advice, and collaborative work from Dr. Ren-Ke Li and his research group,
specifically Dr. Jun Wu, have made a large part of this work possible and I must extend my
gratitude to them. Dr. Abdul Momen must also be acknowledged for his help in performing
much of the early in vivo studies.
All the past and present member of the Laboratory for Functional Tissue engineering have
given immeasurable support and help over the years and all have contributed to making this
project what it is. I must acknowledge Loraine Chiu who’s collaborative efforts undoubtedly
shortened my time here by many years, as well as the efforts of Nicole Feric, Carol Laschinger,
Jason Miklas, Kent Hyun, and Yan Liang for their contributions to this project and ensuring the
work was published.
Friends at U of T, Massey College, and those smart enough to not have gone into graduate
school, you all know who you are and I thank you for being there throughout.
To my family, none of this would have been possible without your love and support. Mom, if
it weren’t for you I never would have attempted this challenge nor seen it through to completion.
Emily, you came into my life at the beginning of this, changed my life forever, and have stood
by me throughout. Thank you and I love you.
Finally, I owe a debt of gratitude to all the animals who gave their lives in the name of science
and contributed to my work. Know that your sacrifice was not in vain.
v
Declaration of Co-Authorship
The original scientific content of this thesis is comprised of two published, peer reviewed
articles in internationally recognized journals. These articles were primarily the work of Lewis
Reis. Introduction and literature review are from three published review papers, the transcribed
writing being that of Lewis Reis. The contributions of co-authors are stated in the thesis, in
conformity with the requirements for the degree of Doctor of Philosophy.
vi
Permissions
Copyright © 2015 Wolters Kluwer Health. Contents of this thesis have been published in
Circulation: Heart Failure: Reis L.A., Chiu L.L.Y., Wu J., Feric N., Momen A., Li R.K., &
Radisic M. (2015). Hydrogels with integrin-binding angiopoietin-1-derived peptide, QHREDGS,
for treatment of acute myocardial infarction. Circulation: Heart Failure. DOI:
10.1161/CIRCHEARTFAILURE.114.001881. Reuse with permission from Wolters Kluwer
Health. A link to the published paper can be found at:
http://circheartfailure.ahajournals.org/content/early/2015/01/28/CIRCHEARTFAILURE.114.001
881.abstract
Copyright © 2014 John Wiley & Sons, Ltd. Contents of this thesis have been published in the
Journal of Tissue Engineering and Regenerative Medicine: Reis L.A., Chiu L.L.Y., Feric N., Fu
L., & Radisic M. (2014). Biomaterials in myocardial tissue engineering. Journal of Tissue
Engineering & Regenerative Medicine. DOI:10.1002/term.1944. Reuse with permission from
John Wiley & Sons, Ltd. A link to the published paper can be found at:
http://onlinelibrary.wiley.com/doi/10.1002/term.1944/abstract
Copyright © 2012 Frontiers in Bioscience. Contents of this thesis have been published in
Frontiers in Bioscience: Chiu LLY, Iyer RK, Reis LA, et al. (2012). Cardiac tissue engineering:
current state and perspectives. Frontiers in Bioscience. 17:1533-1550. Reuse with permission
from Frontiers in Bioscience. DOI: 10.2741/4002. A link to the published paper can be found at:
www.bioscience.org/2012/v17/af/4002/list.htm
Copyright © 2012 Elsevier. Contents of this thesis have been published in Acta Biomaterialia.
Reis LA et al. (2012). A peptide-modified chitosan-collagen hydrogel for cardiac cell culture
and delivery. Acta Biomaterilia. 8(3): 1022-36. Reuse with permission from Elsevier. DOI:
10.1016/j.actbio.2011.11.030. A link to the published paper can be found at:
www.sciencedirect.com/science/article/pii/S1742706111005332
Copyright © 2011 Elsevier. Contents of this thesis have been published in Current Opinion in
Biotechnology: Iyer RK, Chiu LLY, Reis LA, & Radisic M. (2011). Engineered cardiac tissues.
vii
Current Opinion in Biotechnology. 22(5):706-714. Reuse with permission from Elsevier. DOI:
10.1016/j.copbio.2011.04.004. A link to the published paper can be found at:
www.sciencedirect.com/science/article/pii/S0958166911000668
Copyright © 2014 Elsevier. Contents of this thesis have been published in Cardiac
Regeneration & Repair (Vol. II, Ch. 3). Reis LA et al. (January 2014). Injectable biomaterials
for cardiac repair. In Li R.K. & Weisel R.D. (EDs), Cardiac Regeneration & Repair (Vol. II, Ch.
3). Cambridge: Woodhead Publishing. Reuse with permission from Elsevier. DOI:
10.1533/9780857096715.1.49. A link to the book chapter can be found at:
http://www.sciencedirect.com/science/article/pii/B9780857096593500037
Copyright © 2011 Springer. Contents of this thesis have been published in Biomaterials for
Tissue Engineering Applications: A review of past & future trends: Odedra D, Chiu LLY, Reis
LA, et al. (2011). Cardiac Tissue Engineering. In J. A. Burdick & R. L. Mauck (EDs.),
Biomaterials for Tissue Engineering Applications: A review of past & future trends (1st ed., p.
562). Vienna: Springer Vienna. Reuse with permission from Springer. DOI: 10.1007/978-3-
7091-0385-2_15. A link to the book chapter can be found at:
http://link.springer.com/chapter/10.1007/978-3-7091-0385-2_15/fulltext.html
viii
Abstracts of Published Works Appearing in the Thesis
A peptide-modified chitosan-collagen hydrogel for cardiac cell culture and delivery
Reis, Lewis A; Chiu, Loraine L Y; Liang, Yan; Hyunh, Kent; Momen, Abdul; & Radisic, Milica.
Acta Biomaterialia. (2012). 8(6): 1022-1036.
Myocardial infarction (MI) results in the death of cardiomyocytes (CM) followed by scar
formation and pathological remodeling of the heart. We propose that chitosan conjugated with
the angiopoietin-1 derived peptide, QHREDGS, and mixed with collagen I forms a
thermoresponsive hydrogel better suited for the survival and maturation of transplanted
cardiomyocytes in vitro compared to collagen and chitosan-collagen hydrogels alone.
Conjugation of QHREDGS peptide to chitosan does not interfere with the gelation, structure or
mechanical properties of the hydrogel blends. The storage modulus of 2.5 mg ml(-1) 1:1
mass:mass (m:m) chitosan-collagen was measured to be 54.9 ± 9.1 Pa, and the loss modulus
6.1±0.9 Pa. The dose-response of the QHREDGS peptide was assessed and it was found that
CMs encapsulated in High-peptide gel (651 ± 8 nmol peptide ml-gel(-1)) showed improved
morphology, viability and metabolic activity in comparison to the Low-peptide (100 ± 30 nmol
peptide ml-gel(-1)) and Control (No Peptide) groups. Construct (CMs in hydrogel) functional
properties were not significantly different between the groups; however, the success rate of
obtaining a beating construct was improved in the hydrogel with the High amount of QHREDGS
peptide immobilized compared to the Low and Control groups. Subcutaneous injection of
hydrogel (Control, Low and High) with CMs in the back of Lewis rats illustrated its ability to
localize at the site of injection and retain cells, with CM contractile apparati identified after
seven days. The hydrogel was also able to successfully localize at the site of injection in a mouse
MI model.
Contributions: LAR-concept and design, performed all experiments and data analysis,
manuscript writing; LLYC- concept and design, assistance in surgical work, preparation of samples
for histology, data analysis; YL- image analysis for vasculature; KH- Pico green DNA analysis;
AM-performed surgical work; MR-concept, data interpretation, final approval of manuscript.
ix
Hydrogels with integrin-binding angiopoietin-1-derived peptide, QHREDGS, for treatment
of acute myocardial infarction
Reis, Lewis A; Chiu, Loraine L Y; Wu, Jun; Feric, Nicole; Laschinger, Carol; Momen, Abdul;
Li, Ren-Ke; & Radisic, Milica. Circulation Heart Failure. (2015). DOI:
10.1161/CIRCHEARTFAILURE.114.001881.
Background: Hydrogels are being actively investigated for direct delivery of cells or
bioactive molecules to the heart post-myocardial infarction (MI) to prevent cardiac functional
loss. We postulate that immobilization of the pro-survival angiopoietin-1-derived peptide,
QHREDGS, to a chitosan-collagen hydrogel could produce a clinically translatable thermo-
responsive hydrogel to attenuate post-MI cardiac remodeling.
Methods & Results: In a rat MI model, QHREDGS-conjugated hydrogel (QHG213H),
Control gel, or PBS was injected into the peri-infarct/MI zone. By in vivo tracking and chitosan
staining, the hydrogel was demonstrated to remain in situ for 2 weeks and was cleared in ~3
weeks. By echocardiography and pressure-volume analysis, the QHG213H hydrogel
significantly improved cardiac function compared to the controls. Scar thickness and scar area
fraction were also significantly improved with QHG213H gel injection compared to the controls.
There were significantly more cardiomyocytes (CMs), determined by cardiac troponin-T
staining, in the MI zone of the QHG213H hydrogel group; and hydrogel injection did not induce
a significant inflammatory response as assessed by PCR and an inflammatory cytokine assay.
The interaction of CMs and cardiac fibroblasts with QHREDGS was found to be mediated by β1-
integrins.
Conclusions: We demonstrated for the first time that the QHG213H hydrogel can be injected
in the beating heart where it remains localized for a clinically effective period. Moreover, the
QHG213H hydrogel induced significant cardiac functional and morphological improvements
post-MI relative to the controls.
Contributions: LAR-concept and design, performed all experiments and data analysis,
manuscript writing; LLYC- concept and design, assistance in surgical work, preparation of samples
for histology, image & data analysis; JW- design, performed surgical work, preparation of samples
for histology, data analysis and interpretation, manuscript writing; NF- Attachment assay,
x
manuscript writing; AM-performed surgical work; RKL & MR-concept, data interpretation, final
approval of manuscript.
xi
Table of Contents
Abstract .................................................................................................................................................... ii
Acknowledgements ................................................................................................................................. iv
Declaration of Co-Authorship ................................................................................................................. v
Permissions ............................................................................................................................................. vi
Abstracts of Published Works Appearing in the Thesis ....................................................................... viii
1. Introduction ................................................................................................................................... 1
1.1. Motivation ................................................................................................................................. 1
1.2. Hypothesis ................................................................................................................................. 2
1.3. Specific aims ............................................................................................................................. 2
2. Literature Review .......................................................................................................................... 3
2.1. Cardiovascular Disease, Outcomes, and Treatment Options .................................................... 3
2.2. Cell sources, stem cells, & differentiation ................................................................................ 5
2.3. Design criteria for biomaterials in cardiac tissue engineering .................................................. 7
2.3.1. Biocompatibility ................................................................................................................ 7
2.3.2. Biodegradability ................................................................................................................ 7
2.3.3. Mechanical Support .......................................................................................................... 9
2.3.4. Injectability ..................................................................................................................... 11
2.3.5. Clinically Relevant Thickness......................................................................................... 11
2.3.6. Application Time ............................................................................................................ 11
2.4. Tissue engineered cardiac grafts (in vitro engineering) .......................................................... 12
2.5. Injection (in vivo engineering) ................................................................................................ 14
2.5.1. Cell only based therapies ................................................................................................ 15
2.5.2. Hydrogels to promote endogenous repair ....................................................................... 16
2.5.3. Hydrogels for the delivery of cells for regeneration ....................................................... 21
2.5.4. Hydrogels for the artificial maintenance of ventricle geometry and repair .................... 26
2.5.5. Clinical application ......................................................................................................... 29
2.6. Chitosan & collagen ................................................................................................................ 30
2.7. QHREDGS peptide ................................................................................................................. 33
2.8. Peptide modified chitosan-collagen hydrogel for cardiac tissue engineering ......................... 34
3. Hydrogel development, in vitro characterization & preliminary in vivo models ........................ 36
xii
3.1. Materials & Methods .............................................................................................................. 36
3.1.1. Peptide modified chitosan (UP-G113-QHREDGS) ........................................................ 36
3.1.1.1. Conjugation .............................................................................................................. 36
3.1.1.2. Assessing conjugation efficiency ............................................................................. 36
3.1.2. Chitosan-collagen hydrogel ............................................................................................ 37
3.1.2.1. Hydrogel formulation ............................................................................................... 37
3.1.2.2. Scanning electron microscopy .................................................................................. 37
3.1.2.3. Hydrogel degradation ............................................................................................... 38
3.1.2.4. Rheological assessment of hydrogels ....................................................................... 38
3.1.3. In vitro cell culture .......................................................................................................... 38
3.1.3.1. CM isolation ............................................................................................................. 38
3.1.3.2. CM media ................................................................................................................. 39
3.1.3.3. CM encapsulation & culture ..................................................................................... 39
3.1.4. In vivo studies.................................................................................................................. 39
3.1.4.1. Subcutaneous injection ............................................................................................. 39
3.1.4.2. Mouse MI model ...................................................................................................... 40
3.1.5. Construct Characterization .............................................................................................. 40
3.1.5.1. Gel compaction ......................................................................................................... 40
3.1.5.2. Live/Dead staining .................................................................................................... 41
3.1.5.3. Functional testing ..................................................................................................... 41
3.1.5.4. XTT assay ................................................................................................................. 42
3.1.5.5. LDH assay ................................................................................................................ 42
3.1.5.6. PicoGreen DNA Assay ............................................................................................. 43
3.1.5.7. Histological & immunofluorescent staining ............................................................. 43
3.1.6. Statistical analysis ........................................................................................................... 45
3.2. Results & Discussion .............................................................................................................. 45
3.2.1. Hydrogel composition and characterization .................................................................... 45
3.2.1.1. Base collagen-chitosan hydrogel .............................................................................. 45
3.2.1.2. Collagen-chitosan-QHREDGS hydrogel .................................................................. 50
3.2.2. In vitro studies with cardiomyocytes in collagen-chitosan-QHREDGS hydrogels ........ 52
xiii
3.2.2.1. Distribution, viability, and metabolism of CMs encapsulated in the hydrogels ....... 52
3.2.2.2. Construct functional & morphological properties .................................................... 57
3.2.3. In vivo studies.................................................................................................................. 59
3.2.3.1. Subcutaneous injection ............................................................................................. 59
3.2.3.2. Mouse MI model study ............................................................................................. 64
3.3. Summary ................................................................................................................................. 69
4. Utility of QH-G213-H hydrogel as a treatment for acute MI ..................................................... 70
4.1. Materials & Methods .............................................................................................................. 70
4.1.1. Experimental design overview ........................................................................................ 70
4.1.2. Peptide modified chitosan-collagen hydrogel (QHG213H) ............................................ 71
4.1.3. Peptide modified polyethylene glycol (PEG-QHREDGS) ............................................. 71
4.1.4. In vivo studies.................................................................................................................. 72
4.1.4.1. Assessment of cardiac function (6 Week time point) ............................................... 72
4.1.4.2. 24 hr time point ........................................................................................................ 73
4.1.4.3. In vivo hydrogel lifespan .......................................................................................... 74
4.1.5. Immunohistochemistry .................................................................................................... 74
4.1.6. Quantitative PCR ............................................................................................................ 75
4.1.7. Western blotting .............................................................................................................. 75
4.1.8. Image analysis techniques ............................................................................................... 76
4.1.9. Statistical analysis ........................................................................................................... 76
4.2. Results ..................................................................................................................................... 76
4.2.1. In vivo degradation (lifespan).......................................................................................... 76
4.2.2. Functional Data ............................................................................................................... 77
4.2.3. Gross morphology and histology .................................................................................... 81
4.2.4. Cardiomyocyte survival mechanism ............................................................................... 84
4.3. Discussion ............................................................................................................................... 86
4.4. Summary ................................................................................................................................. 94
5. Recommendations for future work ............................................................................................. 96
5.1. Specific Aim SA-2. i. Show ability to localize in vivo and quantify duration of its in vivo
presence. 96
5.2. Specific Aim SA-2. iii. Determine mechanism of action. ....................................................... 96
xiv
6. Conclusions & Contributions to New Knowledge ...................................................................... 99
6.1. Publications & Contributions ................................................................................................ 101
7. References ................................................................................................................................. 105
Appendix A. Reaction schematics and chemical structures ........................................................... 133
Appendix B. In vitro construct CM distribution............................................................................. 136
Appendix C. Utility of QHG213H hydrogel for MI treatment published supplementary data ...... 139
Appendix D. Unpublished rat MI model supplementary material .................................................. 146
List of Tables
Table 3-1: Hydrogel rheological assessment ........................................................................... 48
Table 4-1: Comparative-analysis of reported MI studies ......................................................... 94
Table C-1: Genes and primers used in RT-qPCR .................................................................. 139
Table C-2: Description of studies considered for comparative-analysis .............................. 140
List of Figures
Figure 2-1: Histological stages of myocardial infarction ........................................................... 4
Figure 2-2: The structures of chitosan and chitin, and their proportions in 85% deacetylated
chitosan ......................................................................................................................................... 31
Figure 2-3: Molecular structure of Ang-1 derived peptide QHREDGS .................................. 34
Figure 2-4: QHREDGS peptide modified chitosan (QHG213H) ............................................ 35
Figure 3-1: Hydrogel Characterization .................................................................................... 49
Figure 3-2: Characterization of peptide conjugation and hydrogel morphology ..................... 52
Figure 3-3: Gel encapsulation results in uniform distribution of live cells.............................. 53
Figure 3-4: Metabolic activity and total number of cells encapsulated in 1:1 chitosan:collagen
hydrogels ....................................................................................................................................... 56
Figure 3-5: 1:1 Chitosan:collagen hydrogel with immobilized QHREDGS enables cultivation
of beating cardiac tissue ................................................................................................................ 58
Figure 3-6: Subcutaneous injection study ................................................................................ 61
Figure 3-7: Immunostaining for different cell populations in subcutaneously injected nodules
after 7 days in vivo ........................................................................................................................ 63
xv
Figure 3-8: Chitosan:collagen gel is suitable for injection into the infarcted heart ................. 65
Figure 3-9: Vascularization and wall thickness in the mouse MI model ................................. 67
Figure 4-1: In vivo experimental design................................................................................... 71
Figure 4-2: In vivo lifespan of Dy-light 800 labelled chitosan-collagen gel............................ 77
Figure 4-3: Cardiac function post MI measured by echocardiography .................................... 79
Figure 4-4: Load dependent cardiac function and LV volumes at 6 weeks post MI ............... 80
Figure 4-5: Load independent pressure-volume analysis at 6 weeks post MI ......................... 81
Figure 4-6: Gross heart morphology at 6 weeks post MI......................................................... 82
Figure 4-7: Vascularization at 6 weeks post MI ...................................................................... 83
Figure 4-8: Apoptosis at 6 weeks post MI ............................................................................... 85
Figure 4-9: Cardiac cell attachment to PEG and PEG-QHREDGS gels ................................. 86
Figure 4-10: Proposed QHREDGS mediated mechanism ....................................................... 92
Figure A-1: Polyelectrolytic complexation of chitosan and collagen .................................... 133
Figure A-2: Modified chitosans used ..................................................................................... 134
Figure A-3: EDC/Sulfo-NHS reaction chemistry for conjugation of biomolecules .............. 134
Figure A-4: QHREDGS conjugation to PEG......................................................................... 135
Figure B-1: In vitro cell leakage and migration ..................................................................... 136
Figure B-2: Quantification of in vitro construct viability ...................................................... 137
Figure B-3: High QHREDGS hydrogel constructs contain elongated cardiomyocytes with
visible cross-striations ................................................................................................................. 138
Figure C-1: Quantification of hydrogel lifespan .................................................................... 141
Figure C-2: Histological staining 6 weeks post MI ............................................................... 142
Figure C-3: RT-qPCR of MI and border zone tissues from hearts excised 24 hrs post MI ... 143
Figure C-4: Western blot analysis of border zone tissue from hearts excised 24 hrs post MI144
Figure C-5: Reported cardiac morphological and functional data from studies used in meta-
analysis ........................................................................................................................................ 145
Figure D-1: Heart morphology & histology 3Wk post MI .................................................... 146
Figure D-2: Further histological staining 3Wks post MI ....................................................... 147
Figure D-3: 3Wk MI quantification measurements ............................................................... 149
Figure D-4: 3Wk MI scar extent assessment ......................................................................... 150
Figure D-5: 6Wk MI quantification measurements (from histology pictures) ...................... 151
xvi
Figure D-6: Vascularization within the MI scar zone at 3Wks post infarct ........................... 152
Figure D-7: Apoptosis in the MI boundary zone at 3Wk ...................................................... 153
Figure D-8: 3Wk TUNEL/cTnT staining controls ................................................................. 154
Figure D-9: 6Wk TUNEL/cTnT staining raw counts ............................................................ 154
Figure D-10: 6Wk TUNEL/cTnT staining controls ............................................................... 155
1
1. Introduction
Parts of the text used in this chapter has been published in Circulation Heart Failure (Reis et
al. 2015), and is used with permission from Wolters Kluwer Health.
1.1. Motivation
Cardiomyocyte (CM) death, myocardial remodeling, and scar tissue formation following
myocardial infarction (MI) results in severe cardiac dysfunction and serious health problems. If
not treated, and often with treatment, heart failure progresses. Direct transplantation of various
cell types and/or bioactive molecules has shown promise, however both strategies have been
hampered by low injection site retention and even lower long-term survival of the cells/bioactive
molecules (Mayfield et al. 2014; Nelson et al. 2012). To mitigate these problems these
treatments are combined with hydrogels to confine the cells/bioactive molecules to the site of
injection and promote their long-term survival therein.
The acute phase post-MI might be the most appropriate time to utilize a hydrogel-based
therapy because therein hydrogel injections can prevent cardiac remodeling, deliver cells to
replace the damaged tissue, and/or recruit endogenous stem cells (Menasché 2008). Chitosan and
collagen are natural, biodegradable, biocompatible polymers that have been explored for their
potential use in the treatment of cardiac dysfunction (Lu et al. 2009; Rask, Dallabrida, et al.
2010; Garbern et al. 2011; Suuronen et al. 2006; Y. Zhang et al. 2008; Wu et al. 2011).
Typically, collagen and chitosan, alone or in combination, are cross-linked using exogenous,
often toxic, chemical cross-linkers to improve the hydrogel mechanical properties (Lu et al.
2009; Deng et al. 2010; Fujita et al. 2005). We propose that chitosan-collagen composites can gel
naturally at physiological temperatures and pH to form mechanically stable hydrogels that are
appropriate for in vivo application without the need for exogenous cross-linkers. Furthermore,
the collagen-chitosan interaction within the gels resembles the collagen-glycosaminoglycan
interaction found in vivo in the extracellular matrix (ECM) (Tan, Krishnaraj, and Desai 2001).
Thus, chitosan-collagen may mediate physiological cell-matrix interactions.
The functional success of hydrogel-based cardiac and cell therapies can be improved by
modifying biomaterials with bioactive molecules because they (cytokines, growth factors, etc.)
have the potential to increase transplanted cell survival, reduce resident cell apoptosis, recruit
desired regenerative cells, and promote stem/progenitor cell differentiation (Chiu, Iyer, et al.
2
2012; Reis et al. 2014a). One such bioactive molecule is the growth factor angiopoietin-1
(Ang1). In endothelial cells, Ang1 binds the Tie2 receptor (Hayes et al. 1999; I. Kim et al. 2000;
Papapetropoulos 2000) but in cells such as neonatal rat cardiomyocytes (NCMs) that lack the
Tie2 receptor, Ang1 binds to integrins (Carlson et al. 2001) and activates pro-survival pathways
(Dallabrida et al. 2005). We identified the short sequence QHREDGS as the integrin-binding
motif of Ang1, and the QHREDGS peptide was found to support CM attachment and survival
similar to full-length Ang1 (Rask, Dallabrida, et al. 2010; Rask, Mihic, et al. 2010). It is
therefore possible that the QHREDGS peptide could retain/restore cardiac contractile function
post-MI by promoting CM survival. Importantly, the QHREDGS peptide is water-soluble, stable,
fully-synthetic with a precisely defined composition, and does not require a specific orientation
to be functional. We therefore look to incorporate the pro-survival peptide QHREDGS into an
optimized chitosan-collagen hydrogel to produce a novel hydrogel-based cardiac regenerative
therapy. We chose to immobilize the peptide onto the hydrogel because the effects of
QHREDGS were found to be attachment-dependent, and immobilization would ensure sustained
localization at the injection site and requires a single low dose.
1.2. Hypothesis
Conjugation of Angiopoietin-1 derived peptide QHREDGS to chitosan, injected as a peptide-
modified chitosan-collagen gel, will enhance survival of cardiac muscle after myocardial
infarction in vivo and attenuate pathological remodeling.
1.3. Specific aims
SA-1. Develop a chitosan-collagen hydrogel that supports CM viability and phenotype.
Conjugate QHREDGS and demonstrate dose dependent improvement in CM phenotype and
function in vitro.
SA-2. Assess effect of the developed hydrogel in a rat MI model.
i. Show ability to localize in vivo and quantify duration of its in vivo presence.
ii. Demonstrate potential to attenuate cardiac remodeling post MI.
iii. Determine mechanism of action.
3
2. Literature Review
Contents of this chapter have been published in (Iyer et al. 2011; Odedra et al. 2011; Chiu,
Iyer, et al. 2012; Reis et al. 2012; Reis et al. 2014b; Reis et al. 2015) and are used with
permission from respective publishers.
2.1. Cardiovascular Disease, Outcomes, and Treatment Options
Cardiovascular disease (CVD) is currently the leading cause of death in the world and it is
projected that it will remain as such throughout the next decade (World Health Organization
2011a; World Health Organization 2011b; Mathers and Loncar 2006). In 2008 alone, CVD
accounted for 1 in 3 deaths in the US (811,940 of 2,471,984) and of those, half (approximately 1
in 6 American deaths) were attributable to coronary heart disease (Roger et al. 2012).
Furthermore it is estimated that each year 785,000 Americans have a new coronary attack, about
470,000 have a recurrent attack and a further 195,000 suffer their first silent myocardial
infarction (Roger et al. 2012). While the disorders categorized as CVDs have such divergent
causes as atherosclerosis, rheumatic fever, congenital malformations and thrombosis, they all
converge to cause damage to the heart muscle. Unfortunately, the damage is irreversible because
the heart muscles cells, cardiomyocytes, are thought to be terminally differentiated and non-
proliferative (Sutton and Sharpe, 2000; Laflamme and Murry, 2005), which necessarily limits the
regenerative potential of the heart.
Among CVDs, ischemic heart disease is the most prevalent (World Health Organization
2011b; Roger et al. 2012). It has therefore been the subject of intensive research. The partial or
complete blockage of a coronary artery reduces/prevents blood supply to the downstream heart
muscle and the affected tissue becomes severely nutrient and oxygen deprived, inducing
cardiomyocyte death: an event known as a myocardial infarction (MI). Myocardium undergoes
irreversible damage within 20 minutes of MI and a subsequent wave-front of cell death sweeps
over the area of ischemia over a three to six hour period finally resulting in the death of up to a
billion cells (M. A. Laflamme and Murry 2005). In the weeks that follow, cardiac function is
greatly reduced due to the invasion of leukocytes into the infarct area, removal of dead tissue,
and deposition of granulation tissue. Occurring in as little as two months the ventricular
remodeling results in the formation of tough, rigid, collagenous fibrotic scar tissue at the site of
(and surrounding the) infarct, and thinning of the ventricular wall; both leading to increased
4
ventricular wall stress and volume, and decreased ejection fraction and contraction force; see
Figure 2-1(Sutton and Sharpe 2000; M. A. Laflamme and Murry 2005). In the end, without
intervention and in most cases even with intervention, a myriad of health complications manifest,
ultimately leading to irreversible heart failure and premature death.
Figure 2-1: Histological stages of myocardial infarction (Reproduced with permission from Nature Publishing Group, Laflamme and Murry 2005)
Even with treatment of the underlying cause of heart disease, approximately 50% of patients
will experience heart failure within five years of an acute MI (Pantilat and Steimle 2004; Go et
al. 2013). At present, the gold standard treatment for those who have reached end-stage heart
failure is a heart transplant; but the insufficiency of donors combined with the need for patient-
donor matched organs severely limits the number of patients that can be treated (Zammaretti and
Jaconi 2004; Menasché 2008). In 2010, there were 2,333 heart transplants performed in the US
(Roger et al. 2012). In 2011, greater than 3,000 people were on the waiting list for a heart
transplant and thus living with end-stage heart failure (Roger et al. 2012).
Over the last few years left ventricular assist devices (LVADs) have shown great
improvement in safety and efficacy. As such, LVADs have been increasingly used as a bridge to
transplantation, giving those awaiting a transplant a credible life-saving therapy option by
supporting the failing heart, with the consequence of reducing transplant waitlist deaths (Wilson
et al. 2009; Kirklin et al. 2013). Additionally, after months to a year on an LVAD, many patients
experience improvement in the global contractile properties of the heart (e. g. ejection fraction)
allowing their removal from the device. More recently, some VADs have been approved as a
destination therapy, providing individuals ineligible for a heart transplant with a long-term
therapy option (Wilson et al. 2009; Kirklin et al. 2013). However, there remains a subset of
5
patients that do not experience a significant and/or long-lasting improvement on an LVAD.
Moreover, there is a great need to develop therapies that prevent end-stage heart failure.
Consequently, the field of cardiac tissue engineering emerged as a means of developing
alternative sources of cardiac tissue and methods for replacing tissue damaged by CVD.
Biomaterials have featured prominently in cardiac regenerative therapy and can be divided
among two main strategies for inducing functional repair to the heart muscle post-injury: (i) the
production of functional cardiac patches in vitro that can be implanted onto the damaged area
and thereby directly replace the non-viable portion of a damaged heart, and (ii) injectable
biomaterials (hydrogels) to help prevent cardiac remodeling, deliver cells to replace the damaged
tissue, and recruit endogenous cell types in an attempt to restore functionality (Menasché 2008).
In considering possible delivery of cells to the heart for tissue repair, the first issue that arises is
that of cell source.
2.2. Cell sources, stem cells, & differentiation
One of the biggest issues with cardiac repair is that adult cardiac cells do not proliferate, so
damaged tissue cannot repair itself. One cannot take a heart tissue biopsy, expand the cells to a
sufficient number, and re-introduce them back into the patient. This has led to the need to find a
cell source, ideally autologous, capable of differentiating into functional cardiomyocytes for use
in cardiac tissue engineering. Many cell types have been assessed or used as potential
cardiomyocyte replacements post infarction (Q. Z. Chen et al. 2008). The use of various adult
cell types including skeletal myoblasts and mesenchymal and hematopoietic stem cells have all
been assessed and although advantageous as they can come from autologous sources, results
have been varied, as will be discussed in coming sections. Neonatal cardiomyocytes are a
common cell source for they retain some proliferative capacity and have shown ability to couple
with adult myocytes, however their clinical use in humans is not feasible (Reinecke et al. 1999;
Reffelmann and Kloner 2003). They do however remain a good model cell source in developing
methods for regeneration as they provide a model of what can be achieved (Reffelmann and
Kloner 2003).
Embryonic and induced pluripotent stem cells (ESC and iPS cells, respectfully) are at the
moment the most promising cell sources for true myocardial repair. Studies from a number of
groups have shown that it is possible to generate CMs from mouse and human ESCs, and human
iPS cells (Kattman, Huber, and Keller 2006; Yang et al. 2008; M. Zhang et al. 2001).
6
Flk1+ cells derived from ESCs and cultivated as embryonic bodies (EBs) were identified as
multipotent cardiovascular progenitors, and it was further demonstrated that Flk1 expression can
be used to enrich for cardiac specific progenitors (Kouskoff et al. 2005). When isolated from the
differentiated EBs, on day four of culture, and cultured as a monolayer, these progenitors
generate cultures highly enriched for contracting CMs. As these progenitors differentiate they
progress through the developmental stages thought to be involved in the establishment of the
cardiovascular lineages in vivo and the presence of specific cytokines is required. A combination
of activin-A and bone morphogenetic protein-4 over the first four days of EB differentiation
induces the initial stages of specification and subsequent exposure to dickkopf homolog 1
(DKK1) and vascular endothelial growth factor (VEGF) significantly enhances the
differentiation of Flk1+ progenitors into CMs (Kattman, Huber, and Keller 2006; Yang et al.
2008). The method reported shows the potential of the indentified progenitor to successfully
differentiate into the three cell types found in healthy adult tissue and thus is of great interest as a
cell source. There does remain some concern in that ESCs are associated with teratoma
formation due to their proliferative capacity and as they are not an autologous source there is
immune response issues, as well as the ethical issues of using ESCs (Vunjak-Novakovic et al.
2010).
A method for producing iPS cells from mouse, and subsequently human, somatic cells was
originally reported by Yamanaka and colleagues, and is significant as iPS cells present a much
needed autologous cell source for transplantation into the heart and differentiation into CMs and
vasculature (Takahashi and Yamanaka 2006; Takahashi et al. 2007). Problems with the reported
method, however, are low success rate at producing iPS cells (1/10000) and the use of a viral
vector to deliver the four factors needed to induce pluripotency. The latter is disconcerting as it
entails modification of the cells genome and thus raises questions of long-term effects. A new
method recently reported by Warren et al. (2010) uses mRNA to deliver the same transcription
factors as Yamanaka but with much higher efficiency and success rate at producing iPS cells
very close in phenotype to true ESCs. Furthermore, addition of further mRNA with
differentiation factors showed success at directing iPS cell fate (Warren et al. 2010). Advantages
of the newly reported method are that cell reprogramming to a pluripotent state and subsequent
differentiation can be achieved without compromising genetic integrity (Warren et al. 2010). As
advances in iPS cell production continues they remain a clear favorite in terms of a bench mark
7
cell source for transplantation, however the concern of teratoma formation upon transplantation,
as with ESCs, remains.
2.3. Design criteria for biomaterials in cardiac tissue engineering
The complexity of the native heart environment and pathophysiology of post-MI remodeling
challenges the development of strategies for the treatment and design of biomaterials. An
important aspect of biomaterial design is the consideration of the objective of the biomaterial
application. Design criteria will be different whether the biomaterial is a vehicle for cell delivery,
a functionalized material aimed at artificially retaining normal ventricular geometry, or a
scaffold to generate tissue patches. However, there are some common design criteria that must
first be addressed: (i) biocompatibility, (ii) biodegradability, (iii) mechanical support, (iv)
injectability (in the case of hydrogels), (v) clinically relevant thickness (in the case cardiac
patches), and finally (vi) envisioned application time post-infarct (Leor, Landa, and Cohen 2006;
Q. Z. Chen et al. 2008; Vunjak-Novakovic et al. 2010; Bouten et al. 2011).
2.3.1. Biocompatibility Biocompatibility is generally defined as the “ability of a material to perform with an
appropriate host response in a specific application” (Williams 1987). In the context of cardiac
tissue engineering this encompasses the need for the material to function without initiating a
significant foreign body response in vivo, while retaining the ability to support both
cardiomyocyte survival in vitro and in vivo readily without cytotoxicity, as well as the contractile
function of the myocardium (Leor, Landa, and Cohen 2006; Q. Z. Chen et al. 2008; Vunjak-
Novakovic et al. 2010). This does not preclude the activation of the host inflammatory and
immune response but rather focuses on mitigating and controlling the type of response in order
to prevent further injury to the heart and not impede its function. Specifically, a biocompatible
material should be resistant to blood clotting and bacterial colonization, and if immunogenic
should not recruit cell types that can exacerbate the remodeling process. For example, activation
of the host immune response such that there is preferential recruitment of reparative M2
macrophages over cytotoxic M1 macrophages is generally considered to be a beneficial trait for
biomaterials.
2.3.2. Biodegradability
8
Biodegradability refers to the mechanism through which an implanted material breaks down
and the inherent life-span of the material. A thorough discussion of the various mechanisms and
subsequent definitions are covered in detail elsewhere (Treiser et al. 2010). However, in brief,
there are three modes to consider: (i) bioerosion is degradation of a material through hydrolytic
mechanisms (covers both surface and bulk erosion); (ii) bioresorption is degradation through
cellular activity; and (iii) biodegradation is degradation through enzymatic activity. In the
context of cardiac tissue engineering, a biomaterial is considered to be biodegradable if
degradation occurs through disintegration, a hydrolytic mechanism or by enzymatic activity that
the biomaterial will encounter in vivo; and that the degradation products similarly conform to the
requirements of both biocompatibility and biodegradability (Leor, Landa, and Cohen 2006; Q. Z.
Chen et al. 2008; Vunjak-Novakovic et al. 2010; Bouten et al. 2011). While biocompatibility and
biodegradability are distinct concepts, they are often considered in tandem during biomaterial
design as there is little use in designing a biocompatible material that degrades into toxic
components.
Many physiological extracellular matrix (ECM)-based biomaterials readily fit these dual
criteria (e. g. fibronectin and collagen) as they inherently contain the correct molecular
composition required for cell attachment and survival, and they are readily degraded in vitro and
in vivo within days to weeks by enzymes secreted by cells into biocompatible and biodegradable
degradation products (Q. Z. Chen et al. 2008). Notably, cells can turnover these ECM
biomaterials and replace them with their own ECM components, thereby remodeling their
environment as necessary (Z. Li and Guan 2011). However, sourcing these materials can be
complicated by the fact that they may retain many of their surface antigens and may elicit an
immune response if used in xeno-transplantation. Despite this limitation there are sources of the
ECM-biomaterials that have been approved for human use, including fibrin that can be isolated
from a patient’s own blood (Odedra et al. 2011).
Conversely, synthetic materials have been developed such as polyethylene glycol (PEG), poly
(glycerol-sebacate) (PGS), and poly (tetrafluoroethylene) (PTFE) with defined chemical
compositions and designed to have no foreign body response. Modification of the chemical
composition can permit the selection of degradation rates in the range of a few weeks to years.
However there is the concern as to whether the degradation products are truly being removed
from the body or rather accumulate, the long-term effects of which are unknown (Q. Z. Chen et
9
al. 2008; Z. Li and Guan 2011). An additional limitation is that these synthetic biomaterials often
do not support cell adhesion and survival, and therefore need to be modified with appropriate
bioactive molecules (Vunjak-Novakovic et al. 2010).
Another important consideration in terms of biodegradability is the issue of how quickly the
material should be removed in order to properly execute its desired function when applied for
cardiac ventricular repair. Biomaterials designed for cell delivery, recruitment, and survival
(anti-apoptotic, pro-angiogenic) should survive in vivo at least one week, based on the fact that
most cell death occurs within the first few days post-MI; and should be fully degraded in 6-8
weeks (in animal models) recognizing that pathological remodeling is complete by
approximately 6 weeks after MI (Patten et al. 1998; Krzemiński et al. 2008; Z. Li and Guan
2011). Thus, the biomaterial should remain long enough to have the desired effect but no longer
than necessary as it may become a hindrance to repair. For example, improved cardiac function
post treatment was demonstrated with fibrin glue used to transplant skeletal myoblasts into
ischemic myocardium wherein it degraded in 7-10 days (Christman, Vardanian, et al. 2004).
Scaffold degradation is likely to require a similar timeframe. Most scaffolds are designed to be
quickly replaced by new ECM secreted from the seeded cells, with the engineered tissues that are
implanted containing very little of the original scaffold material. For scaffolds that are present at
the time of implantation, in vivo degradation should not exceed weeks to months and should be
quickly replaced by functional tissues. Zimmermann et al. have spent considerable time
developing functional cardiac tissue constructs in vitro demonstrating that their collagen-based
scaffold can be remodeled and replaced by maturing cardiomyocytes resulting in a scaffold-free
transplantable engineered heart tissue construct (Eschenhagen et al. 1997; W.-H. Zimmermann
2001). In the case of biomaterials designed to provide support to the failing ventricle, they
should have relatively slow degradation rates on the order of months to years and while
controversy persists as to the timeframe required, it may be desirable for such materials to
remain for the very long-term (see Section 2.5.4 for examples and further discussion) (Nelson et
al. 2011). It is important to note however that there is still much debate as to the mechanical
properties and degradation requirements that are necessary for certain outcomes due to the large
disparity between studies (Nelson et al. 2011). These topics therefore continue to be active areas
of investigation.
2.3.3. Mechanical Support
10
The third criteria of mechanical support requires some forethought as to the envisioned
application with consideration of whether the biomaterial can withstand the mechanical demands
that will be placed upon it after ventricular application and also whether it is likely to interfere
with the normal mechanical functioning of the surrounding tissue. The in vivo model system for
the biomaterial should be considered with respect to species-specific mechanical demands. This
is because the mechanical forces placed upon a biomaterial by the human heart will vastly differ
from the forces exerted by a small rodent heart in an animal application. Specifically, the human
myocardium ranges in stiffness from 20kPa (end of diastole) to 500kPa (end of systole), whereas
rat myocardium ranges from 0. 1 to 140kPa (Q. Z. Chen et al. 2008; Vunjak-Novakovic et al.
2010; Bouten et al. 2011; Venugopal et al. 2012). A material envisioned to artificially thicken the
ventricle wall and maintain ventricular geometry during remodeling should have a stiffness in
the high end of the range characteristic for the native ventricle; whereas a material designed to be
injected, to act as a temporary matrix for transplanted cells and/or to recruit endogenous cells can
have a low-end stiffness so long as it is sufficiently stiff to withstand the contraction/dilation of
the heart. Biomaterials intended for in vitro applications can have a very low stiffness, as long as
the cells seeded into it (hydrogels or scaffold) are able to remodel it into a material that as a final
product is mechanically similar to the native myocardium. This is especially evident in the case
of cardiac tissue grafts for repair of full thickness defects, wherein mismatching the mechanical
properties can result in the grave consequences of inducing undue strain on the injured heart if
the graft is too stiff or graft failure due to the stresses experienced in vivo if the graft is
insufficiently stiff (Ozawa et al. 2002). Furthermore, issues such as burst pressure and suture
retention must be considered as the graft has no time to integrate with the host tissue since it
experiences full cardiac load immediately upon implantation. Results from tissue engineering of
arterial grafts provide useful benchmarks for these properties that are achievable using current
tissue engineering methods e.g. burst pressure of over 3000mmHg and suture retention strength
of over 160g (L’Heureux et al. 2006; Dahl et al. 2011). Motivated by these challenges, Lang et al
have been developing a surgical glue to improve cardiac grafts and achieved success in repairing
a full thickness defect (N. Lang et al. 2014).
In general, naturally-derived biomaterials have weak mechanical properties, with moduli in
the tens of Pa to tens of kPa range (Q. Z. Chen et al. 2008). Moreover, there is a batch-to-batch
and source-to-source variability in the physical properties of these biomaterials. As a
11
consequence, biomaterials made purely of naturally-derived components are limited as to their
mechanical support applications. Synthetic materials, on the other hand, are more consistent in
their composition between batches and have mechanical properties such as stiffness, elasticity,
and porosity that can be precisely controlled (Q. Z. Chen et al. 2008).
2.3.4. Injectability A hydrogel that can pass through a fine gauge needle (~27G) is described as injectable as it is
possible to safely administer it into the heart in a minimally invasive manner. Injectability can be
achieved by two approaches wherein gelation of the hydrogel (by temperature, chemical, light-
induced cross linking (Yeo et al. 2007; Habib et al. 2011), etc.) is: (i) initiated but not completed
prior to the hydrogel passing through the needle, and (ii) initiated after delivery to the desired
site. Importantly, the polymerization time should be in the order of minutes to tens of minutes to
ensure the hydrogel is delivered and successfully localized at the site of injection and not
completely washed out (Vunjak-Novakovic et al. 2010). This is because for polymerization that
requires tens of minutes to hours to complete there is enough time for the biomaterial to be
subjected to the contraction of the heart and to be carried away in the blood stream, rather than
gel properly in the ventricle wall.
2.3.5. Clinically Relevant Thickness Biomaterials used for tissue engineering strategies where cells are cultured with the
biomaterial in vitro then implanted in vivo have their own distinct requirement in that they must
be capable of supporting the cultivation of tissues of clinically relevant thickness: up to ~10 mm
for full thickness cardiac grafts, whereas tissue patches can be thinner (Chiu and Radisic 2010).
The limits of oxygen diffusion within a metabolically active tissue of high cell density (e.g.
108cells/cm3) restricts tissue thickness to approximately 200µm, thus scaffolds often require that
a primitive vascular network or a channel array for culture medium perfusion be incorporated
into the design to allow sufficient nutrient exchange to the centre of the tissues during in vitro
cultivation (Radisic et al. 2006). This is a major issue in producing in vitro cardiac patches that
will be discussed in greater detail in Section 2.4.
2.3.6. Application Time An additional consideration that can influence biomaterial design for cardiac regeneration
therapy is the time post-infarction at which the biomaterial is to be applied since new and old
infarcts present their own unique challenges. The rapid cell death that results from nutrient and
12
oxygen deprivation downstream of a coronary artery blockage suggests that a cell injection-
based strategy should be most effective if applied shortly after an MI. The application of a
biomaterial modified with both cells and growth factors within hours or days after an MI may
promote directed wound repair such that the scar tissue formed would be minimized, the
contractile function maintained in the border zone, the ischemic area reduced, and consequently
pathological remodeling attenuated. Recent work has shown that early application of a collagen
based hydrogel 3 hours after MI significantly improved pathological remodeling when measured
4 weeks after treatment in comparison to application of gel 7 or 14 days after infarct (Blackburn
et al. 2014). Injection 7 days after MI did show significantly improved cardiac function in
comparison to injection 14 days after, which showed no improvement over PBS injection,
however not to the same level injection 3 hours after infarct imparted, thus indicating the earlier
the application the better the outcomes (Blackburn et al. 2014). Notably, incorrectly timed
administration of the therapy could potentially exacerbate the problem. While early delivery of
cells might conceptually be more effective by initiating early re-vascularization and contributing
to the protection of the spared myocardium, it may at the same time expose the delivered cells to
a very hostile environment because of the significant immune response, the presence of cell
death-associated cytokines and the by-products of dead and dying cells in the infarcted area soon
after insult, which can compromise the viability of the cells. The correct time to deliver cells has
been investigated in the clinic with both the TIME and SWISS-AMI trials showing that injection
of cells from 3 days to 4 weeks after infarct did not improve LV cardiac functional or
pathological outcomes (Traverse 2011; Traverse 2012; Sürder et al. 2013). A scaffold-based
contractile tissue engineering strategy while applicable in the acute phase may have a more
significant effect if implanted after scar formation. Similarly, larger areas of damaged cardiac
muscle evident in chronic cases might benefit from a scaffold-based regenerative therapy
approach. Therefore, choosing the right time point post-MI for an intervention is a challenge and
no clear consensus has been reached.
2.4. Tissue engineered cardiac grafts (in vitro engineering)
In vitro tissue engineering approaches involve pre-organizing cells seeded into scaffolds or on
their own into a functional tissue (i.e. one capable of contracting and propagating electrical
signals). To be clinically relevant any engineered cardiac tissue must have properties (functional
and morphological) similar to that of native heart tissue and remain viable after implantation
13
(Wolfram-Hubertus Zimmermann, Melnychenko, and Eschenhagen 2004). Challenges in the
development of such in vitro tissues is that oxygen diffusion limits tissue sizes to 100-200 µm,
thus vascularization is required, and generating tissues that can electromechanically couple with
the native myocardium (and generate significant contractile force) (Vunjak-Novakovic et al.
2010). A number of different groups have tried to tackle these problems using different
approaches ranging from using decellularized tissue as a scaffold, using mechanical/electrical
stimulation of cells seeded in hydrogel scaffolds, to layering individual cell sheets.
In an effort to improve the properties of engineered heart tissue (EHT) Eschenhagen et al.
(1997) seeded a mix of collagen I, extracellular matrix proteins (Matrigel), and neonatal rat
cardiomyocytes into lattice or circular molds and saw spontaneous remodeling of the liquid
reconstitution mixture and a development of spontaneously and synchronously contracting solid
EHTs after 5-7 days (Eschenhagen et al. 1997). Subsequent culture of EHTs with cyclic
mechanical strain improved morphological, functional, and mechanical properties of the EHT,
and by stacking a number of these constructs together formed spontaneously contracting tissue 1-
4 mm thick that could be implanted at the site of infarct in a rat MI model. One month later it
was observed that the constructs showed un-delayed electrical coupling with the native tissue
and improved diastolic and systolic function of rats who received the patches in comparison to
sham-operated rats (W. H. Zimmermann et al. 2006).
Another approach to mechanical stimulation for achieving functional improvement of EHT
was noted by Radisic et al. (2004). It was demonstrated that neonatal cardiomyocytes seeded on
collagen sponges with Matrigel and exposed to physiologically relevant electric field stimulation
during culture induced the formation of mature myocardium with elongated, viable cells aligned
in parallel. Ultra-structural organization in stimulated cardiac constructs was remarkably similar
to that present in the native myocardium, with hallmarks of electromechanical cell coupling
including gap junctions, intercalated disks and sarcomeres all markedly more frequent than in
non-stimulated constructs (Radisic et al. 2004).
A pioneering study done by Ott et al. (2008) looked to tackle the larger problem of whole-
organ transplant, while at the same time illustrating an approach for answering the questions of
oxygen transport and tissue thickness in engineered heart tissues. By decellularizing whole,
adult, cadaveric Fisher rat hearts by coronary artery perfusion of detergents they were able to
obtain whole heart scaffolds. The process preserved the underlying extracellular matrix (ECM),
14
and produced an acellular scaffold with perfusable vascular architecture, competent acellular
valves and intact chamber geometry. Reseeding the scaffolds with neonatal rat CMs or
endothelial cells and culturing by perfusion in a bioreactor mimicking cardiac physiology they
observed macroscopic contractions by day four. At day eight they were able to show a small, but
significant pump function comparable to about 2% of the adult heart when the engineered hearts
were exposed to physiological loads and electrical stimulation (Ott et al. 2008).
A method for developing functional EHT was also accomplished using no scaffold at all, but
stacking sheets of neonatal rat cardiomyocytes cultured on the surface of Poly(N-
isopropylacrylamide)-grafted polystyrene dishes. The treated dishes allow for cell attachment at
37°C, but culturing at 20°C changes the properties of the surface causing cell layers to detach, at
which point individual layers can be stacked, and these stacked sheets looked like homogeneous
heart-like tissue (Miyagawa et al. 2005). When these scaffold free cell stacks were transplanted
onto infarcted rat hearts they became attached to the infarcted myocardium, showed
angiogenesis, expressed connexin-43 (a gap junction protein associated with functionally
coupled cells), appeared as a homogeneous tissue in the myocardium, and significantly improved
cardiac performance (determined by echocardiography) in comparison to those that did not
receive the treatment (Miyagawa et al. 2005). Furthermore, connexin-43 staining and
transmission electron microscopy (TEM) showed the existence of gap junctions and intercalated
disks (associated with coordinated cardiac muscle contraction) between implanted and host
tissue. One limitation of this technique is that it is still limited to the hundreds of µm in thickness
unless methods for oxygen transport or vascularization are introduced.
Many of the issues with engineering tissue in vitro are the same as those for direct in vivo
application (injection methods), including appropriate biomaterial (scaffold) choice, cell source,
and survival/integration once transplanted, thus the two methods are not mutually exclusive.
Injection of cells/hydrogels into heart tissue shortly after MI has the potential to minimize the
formation of scar tissue and attenuate the pathological remodeling process whereas tissue
engineered cardiac patches may be more useful in later stages of remodeling such as the
complete replacement of non-contractile areas. In vitro tissue engineering may also provide
living patches for repair of congenital malformations that cannot be addressed through injectable
methods (Soonpaa and Field 1998).
2.5. Injection (in vivo engineering)
15
The first type of injection that must be discussed is that of direct cell transplantation. Based on
the successes and limitations from such injections the use of biomaterials was developed to aid in
the successful transplantation of cells. In general, injectable hydrogels for cardiac regeneration
and repair can be divided into three categories according to their state of development and
current application. The first group of hydrogels is designed to promote endogenous repair
through maintaining cell survival, recruiting endogenous cells or inducing neovascularization.
The second group acts a temporary matrix for cell transplantation and exogenous repair. The
final group consists of injectable hydrogels that act as bulking material to support the failing left
ventricle, in turn maintaining or restoring normal geometry of the heart, and improving cardiac
function. It should be noted that there is a significant overlap between the categories, and most
hydrogels fit into more than one category.
2.5.1. Cell only based therapies The first type used in clinical studies for cell based cardiac repair was skeletal myoblasts and
results were varied. Skeletal myoblasts were initially chosen due to their relative ease of isolation
and culture (expansion) from small muscle biopsies, pre-clinical efficacy, non-immunogenicity,
and ability to cope well with ischemic environments (Menasché et al. 2003). Injection showed
reduced infarct size, survival of transplanted cells, and slight improvement of cardiac function,
however it was found that most cells did not trans-differentiate into cardiomyocytes (as hoped,
becoming skeletal muscle) and therefore lack the proteins needed to electromechanically couple
with the native myocardium and it was reported that the grafts did not beat in synchrony with the
surrounding tissue (Reinecke, Poppa, and Murry 2002; Reinecke et al. 2000). Even so, Phase I
trials were deemed successful and Phase II trials are currently underway.
Neonatal cardiomyocytes retain the ability to proliferate and have received much attention as
a potential cell source. Injection of CMs isolated from neonatal rats and transplanted to the site
of infarct in adults have shown attenuation of pathological remodeling leading to improved
function, and even integration of cells into the host myocardium through gap junctions and
intercalated disks (Reinecke et al. 1999; Muller-Ehmsen 2002; R. R. K. Li et al. 1996). Further
studies, however, show that direct injection of these cells results in survival of only ~50% of
cells immediately after injection, loss of up to ~90% of cells due to extrusion, and only ~10%
survival of the remaining cells after as little as one week (Müller-Ehmsen et al. 2002; M. Zhang
et al. 2001).
16
Mesenchymal and hematopoietic stem cells have also been looked at as potential cell sources
for direct injection and have had some promising results. They are advantageous for the two cell
lines have multi-lineage differentiation potential, and in vitro studies showed promise in the two
cell type’s ability to become cardiomyocytes. Direct injection of these cell types into MI areas
showed incorporation into host myocardium, reduced infarct size, and even showed functional
improvements, however it has been noted that fusion events are rare and no significant muscle
mass has been generated in vitro (Bittira et al. 2002; Makino et al. 1999; Shake et al. 2002; Orlic,
Kajstura, Chimenti, Jakoniuk, et al. 2001; Murry et al. 2004).
Perhaps the most interesting observation about cell transplantation is that regardless of cell
type, developmental stage, or origin most studies showed some improvement in cardiac function
post injection; yet all seem to have the same issues with incorporation into the native tissue,
survival, and retention (W. H. Zimmermann et al. 2006). Meta-analysis of clinical trial results
demonstrated a significant, albeit low 3%, increase in left ventricular ejection fraction (LVEF) as
well as a significant reduction in infarct size (-5.6%) and end systolic volume (-7.4ml) in patients
treated by intracoronary cell injection after acute MI (Lipinski et al. 2007). Dose-response
between injected cell volume and LVEF change was reported (Lipinski et al. 2007). Although
these studies are encouraging, modest improvements motivate investigation of new cell sources
and methods that increase survival and retention of injected cells.
It has also been suggested that the improvements seen with cell transplantation may be due to
release of cardio-protective and angiogenic cytokines from transplanted cells that modulate
remodeling rather than integration of the cells into the host myocardium (Menasché 2008). As
such, groups have also looked at identifying specific cytokines and directly injecting them in
place of cells. While the delivery of cells to the injury site is a major aspect of cardiac
regeneration therapy and the cell types used are many and varied, the topic is discussed in further
detail elsewhere (Shiba et al., 2009; Hilfiker et al., 2011; Martinez and Kofidis, 2011; Liau et al.,
2012).
2.5.2. Hydrogels to promote endogenous repair The group of injectable hydrogels that promote endogenous repair works based on either the
bioactivity of the base biomaterial or the incorporation of bioactive molecules that are anti-
apoptotic or cardioprotective (i.e. QHREDGS, Tβ4, insulin-like growth factor-1), angiogenic
(i.e. Tβ4, bFGF, VEGF, PDGF, hepatocyte growth factor, pleiotrophin plasmid), or chemotactic
17
(i.e. stromal cell-derived factor-1). The incorporation of bioactive molecules into injectable
hydrogels is important for the successful delivery of the biomolecules and in turn cardiac
regeneration, since biomolecules in the soluble form are rapidly diffused from the injection site
and proteolytically degraded in the in vivo environment.
The existence of different cardiac progenitor cell types has been reported in native heart
tissue, and it was demonstrated that these progenitors can be induced to migrate to areas of
infarct with the addition of specific peptides or cytokines (Beltrami et al. 2003; Oh et al. 2003;
Martin et al. 2004; Cai et al. 2003). One such peptide is thymosin β-4 (Tβ4) which has been
shown to promote migration and survival of CMs, endothelial cells, smooth muscles cells, and
vascular progenitors from cardiac explants in vitro and has garnered attention for clinical trials as
a drug for patients with acute MI (Smart et al. 2007; Bock-Marquette et al. 2004; Crockford
2007).
A temperature-responsive hydrogel made of chitosan and glycerol phosphate was used to
encapsulate bFGF, which was released through hydrogel degradation (H. Wang et al. 2010).
After 4 weeks of injection in a rat MI model, the bFGF-hydrogel improved cardiac function (i.e.
decreased LV end-diastolic diameter and LV end-systolic diameter, as well as increased LV
ejection fraction and LV fractional shortening), as compared to PBS control, bFGF in PBS, and
chitosan hydrogel alone (H. Wang et al. 2010). This may be due to the increase in arteriole
density and in turn decreased infarct size in the bFGF-hydrogel group compared to all other
groups. The chitosan hydrogel alone also increased arteriole density and decreased infarct size
compared to PBS control and bFGF in PBS (H. Wang et al. 2010). This shows that the presence
of the injectable hydrogel was necessary for mechanical support and sustaining the release of
encapsulated bFGF, while the incorporation of bioactive molecules such as bFGF further
improved endogenous cardiac repair through promoting arteriogenesis. By contrast, treatment
with soluble bFGF showed no improvement compared to PBS control due to the short half life of
the growth factor and its high diffusibility from the injection site (H. Wang et al. 2010). Other
groups also investigated the incorporation and release of bFGF from various hydrogel systems,
including encapsulation into biodegradable gelatin hydrogel microspheres (Iwakura et al. 2003;
Yamamoto et al. 2001; Y. Liu et al. 2006), UV crosslinkable chitosan hydrogels (Fujita et al.
2005), and synthetic random copolymer poly(N-isopropylacrylamide-co-propylacrylic acid-co-
butyl acrylate) (Garbern et al. 2011). These studies all showed increased vascular density with
18
the injection of bFGF hydrogels, and in turn improved cardiac function after myocardium
infarction in rats, rabbits, or dogs.
Conjugation of VEGF to a temperature-sensitive, aliphatic polyester hydrogel, poly (δ-
valerolactone)-block-poly (ethylene glycol)-block-poly (δ-valerolactone) (PVL-b-PEG-b-PVL),
showed that the hydrogel provided localized and sustained VEGF activity, which increased
blood vessel density (Wu et al. 2011). While the hydrogel alone maintained thicker ventricular
wall and improved cardiac function (i.e. fractional shortening, ventricular volumes, preload
recruitable stroke work, end-systolic elastance) after MI in rats, the conjugation of VEGF to the
hydrogel further enhanced cardiac repair (Wu et al. 2011).
While alginate, a bio-inert natural material, has been shown to prevent cardiac remodeling and
dysfunction in the rat MI model, it has also been modified with peptides and growth factors to
create a bioactive injectable hydrogel for endogenous cardiac repair (Landa et al. 2008). Alginate
that was modified with RGD conjugation led to increased angiogenic response compared to non-
modified alginate when injected into the infarct area in rats 5 weeks post-MI (Yu, Gu, et al.
2009).
In addition, alginate hydrogel was found to be an effective injectable delivery system for
cardiac repair post-MI, since it could sequentially deliver VEGF and PDGF to induce formation
of mature blood vessels and in turn improve cardiac function (Hao et al. 2007). In this alginate
system, PDGF was released more slowly than VEGF likely due to the difference in the affinities
of VEGF and PDGF to alginate. As such, VEGF induced angiogenesis, while PDGF was present
in the later time to mature the newly formed capillaries. This sequential delivery approach
improved vascularization and cardiac repair compared to the delivery of VEGF or PDGF
individually (Hao et al. 2007).
Injectable affinity-binding alginate hydrogel microbeads were also used to sequentially
release cytoprotective IGF-1 and pro-angiogenic HGF to increase angiogenesis and formation of
mature blood vessels, as well as to prevent infarct expansion and scar fibrosis after 4 weeks in a
rat acute MI model (Ruvinov, Leor, and Cohen 2011). The injection of these hydrogel
microbeads also reduced cell apoptosis, induced cell cycle re-entry of CMs and enhanced
presence of GATA-4 positive cells, indicating an improved endogenous cardiac regeneration
(Ruvinov, Leor, and Cohen 2011).
19
Hydrogels have also been used to enhance gene therapy. The human VEGF plasmid was
injected in a rat MI model with an amphiphilic, thermo-responsive polymer synthesized by
alternatively cross-linking poloxamer and di-(ethylene glycol) divinyl ether (Kwon et al. 2009).
The sustained human VEGF expression within the infarct area increased capillary density and
formation of larger blood vessels. In a separate study, the pleiotrophin plasmid was shown to
induce neovasculature formation in the ischemic rat myocardium without angioma formation, but
the injection of the plasmid in saline resulted in low transfection efficiency due to limited
exposure of cells to the plasmid (Christman et al. 2005). The incorporation of pleiotrophin
plasmid in fibrin glue created a gene-activated matrix that increased neovasculature formation
compared to injection of the plasmid in saline.
To deliver stromal cell-derived factor-1 (SDF-1) as a chemokine to attract stem cells for
cardiac regeneration, a modified chemokine S4V was designed to retain chemotactic bioactivity
while being resistant to cleavage by matrix-metalloproteinase-2 and exopeptidase (Segers et al.
2007). S4V was then tethered to self-assembling peptides that formed nanofibers suitable for
local delivery. The intra-myocardial delivery of S4V-nanofibers post-MI in rats led to the
recruitment of c-kit positive stem cells and increased capillary density, as well as improved
cardiac function (i.e. increase in ejection fraction) (Segers et al. 2007).
Hydrogels derived from the native cardiac extracellular matrix (ECM) retain the complexity
of the extracellular environment normally found in vivo and were demonstrated to improve
maturation of cardiomyocytes derived from human embryonic stem cells (DeQuach et al. 2010),
thus they constitute a promising approach for myocardial repair (Duan et al. 2011). ECM
components have been isolated from the healthy myocardium and used as injectable hydrogels
for endogenous cardiac repair. Cardiac ECM-derived hydrogels are prepared by decellularizing
heart tissues with sodium dodecyl sulfate, followed by lyophilization and milling into a fine
powder that is then solubilized by enzymatic digestion using pepsin and HCl and neutralized
using NaOH (Singelyn et al. 2009; Seif-Naraghi et al. 2010; Singelyn et al. 2012). Myocardial
matrix yielded from the decellularization of porcine myocardial tissue gelled at 37°C and showed
collagen and glycosaminoglycan content (Singelyn et al. 2009). In addition, the matrix induced
the homing of endothelial and smooth muscle cells, as shown by cell migration towards the
matrix in vitro, as well as cell infiltration into the gel and increased arteriole formation in vivo 11
days post-injection within the rat myocardium. While the solubilized form of the decellularized
20
porcine pericardium was a stronger chemo-attractant for endothelial and smooth muscle cells,
both decellularized porcine and human pericardium were shown to retain the native protein and
glycosaminoglycan and induce neovascularization when injected into the left ventricular free
wall of rats (Seif-Naraghi et al. 2010). Moreover, a low number of c-kit+ cells were found within
the injection site, indicating an increase in endogenous CMs with the injection of the material
(Seif-Naraghi et al. 2010; Singelyn et al. 2012). As such, the use of decellularized cardiac matrix
as injectable hydrogels can promote endogenous homing and in turn cardiac-specific tissue
formation, as mediated through the presence of cardiac-specific cues. In addition, cardiac
function was maintained without arrhythmias in the rat MI model, when treated with the ECM
hydrogel (Singelyn et al. 2012). Besides demonstrating the use of decellularized ECM in small
animal models, it has also been shown that hydrogels made from decellularized porcine
ventricular ECM can be delivered via percutaneous, transendocardial catheter injection in both
healthy and infarcted porcine myocardium (Singelyn et al. 2012). The injection was successful,
with no catheter clogging. While there was some leakage of the injectate into the ventricle, it was
confirmed that a hydrogel was formed within the porcine myocardium without any hydrogel
found in other organs, since any leaked matrix would be rapidly diluted in the blood and
prevented from gelation. This study shows the feasibility of delivering in situ gelling materials
via transendocardial injection in a large animal model and motivates the translation of injectable
materials for MI treatment in humans (Singelyn et al. 2012).
Keeping with ECM derived materials, gels derived from the small intestinal submucosa
extracellular matrix were found to be a suitable injectable material for cardiac repair, due to the
presence of bFGF (Okada et al. 2010). The gel induced angiogenesis and reduced infarct size
compared to saline control in the murine MI model (Okada et al. 2010). In another study, the
injection of small intestine ECM emulsion into the myocardium recruited c-kit positive cells,
myofibroblasts and macrophages after rat MI (Zhao et al. 2010). It also increased VEGF levels,
enhanced angiogenesis and improved cardiac function (i.e. fractional shortening, ejection
fraction, stroke volume) compared to saline control, likely mediated by the cell recruitment
(Zhao et al. 2010).
Similarly, fibrinogen, the precursor to fibrin, has also been used in biomaterial strategies.
Rufaihah et al have demonstrated the applicability of a photocrosslinkable semi-synthetic PEG-
fibrinogen hydrogel loaded with VEGF in rat models of MI. They showed that a PEG-fibrinogen
21
hydrogel was able to store and release VEGF in a sustained and controlled fashion, as well as
significantly improve arteriogenesis and cardiac performance after infarct (Rufaihah et al. 2013).
With tunable mechanical properties they have also shown the ability of the PEG-fibrinogen
hydrogel on its own to mechanically support the failing LV and to improve cell transplantation,
giving the developed hydrogel multiple courses of action as well as promise as a clinically
relevant biomaterial (Habib et al. 2011; Rufaihah et al. 2013).
G-CSF has also been studied in depth but has seen varied results in experiments with various
animal species (Orlic, Kajstura, Chimenti, Limana, et al. 2001; Iwanaga et al. 2004; Deten et al.
2005). While it undoubtedly promotes migration of progenitors to the site of infarct the
population is most likely non-cardiac lineage. It has been suggested, and to some extent shown,
by a number of groups that G-CSF may not induce cardiac regeneration but mitigate the damage
done during MI and influence cardiac repair (Nygren et al. 2004; Minatoguchi et al. 2004). One
group showed in mice that G-CSF treatment reduced final infarct size and resident cell apoptosis,
and increased vascular density and ventricular function in comparison to mice who did not
receive the treatment (Harada et al. 2005). Despite some conflicting studies and varied results G-
CSF has already entered clinical trials being administered to patients with recent MI, and has
shown some promising results (H.-J. Kang et al. 2004; H.-J. Kang et al. 2007; J. Kang et al.
2014).
2.5.3. Hydrogels for the delivery of cells for regeneration Early studies in cell transplantation for cardiac repair lacked control over cell retention,
survival, and function after injection of the cells into the heart. Later studies used natural and
synthetic injectable hydrogels that polymerized in situ as a cell delivery vehicle to overcome
such limitations. Some hydrogels used in cell delivery for cardiac regeneration include fibrin,
chitosan, collagen, Matrigel™, and various synthetic polymers.
Fibrin was one of the earliest biomaterials studied for use in the heart and for cell
transplantation. It exhibits several advantages: it is naturally occurring and could be patient-
specific; it is biocompatible, biodegradable, and pro-angiogenic. Furthermore, fibrin glue has
ideal properties for cell delivery in cardiac repair due to its binding domains for growth factors
and receptors. A three-dimensional fibrin hydrogel is formed upon injection with a dual-barreled
syringe, which contains fibrinogen and fibrinolysis inhibitor aprotinin in the first barrel and
thrombin and CaCl2 in the second barrel (Christman, Fok, et al. 2004). The cross-linking
22
mechanism of the fibrin hydrogel is similar to how clotting occurs in vivo. Skeletal myoblasts
that were injected with fibrin glue into the myocardium post-MI were found in both the border
zone and within the infarct, while myoblasts were found only in the border zone when injected in
bovine serum albumin (Christman, Fok, et al. 2004). Fibrin gel provided a temporary scaffold for
transplanted cells and contained RGD motifs that bind to cell receptors, thus entrapping the cells
upon injection and improving their survival (Christman, Vardanian, et al. 2004). The surviving
myoblasts at five weeks post-injection into the myocardium were located in clumps around
arterioles (Christman, Vardanian, et al. 2004). The injection of fibrin glue with or without
skeletal myoblasts led to smaller infarcts compared to injection of bovine serum albumin or
myoblasts in bovine serum albumin (Christman, Vardanian, et al. 2004). There was also
increased arteriole density in the infarct area when skeletal myoblasts were delivered in the fibrin
glue compared to delivery in bovine serum albumin (Christman, Vardanian, et al. 2004).
Although fibrin seemed to significantly decrease infarct size and increase arteriole formation
without the presence of transplanted cells, it could be seen that transplantation of cells without an
appropriate injectable hydrogel was not capable of appreciable cardiac repair (Christman, Fok, et
al. 2004; Christman, Vardanian, et al. 2004).
In a separate study, rat adipose-derived stem cells isolated from subcutaneous adipose tissues
were injected in fibrin glue into the rat left ventricular wall post-MI (X. Zhang et al. 2010). At 4
weeks after injection, the co-injection of cells and hydrogel showed greater graft size, better
cardiac function (decreased LV end-diastolic diameter and LV end-systolic diameter, as well as
increased fractional shortening and ejection fraction), and increased arteriole density compared
to injection of cells alone (X. Zhang et al. 2010). In addition, another study was conducted to
deliver human bone marrow-derived mesenchymal stem cells in fibrin glue in a nude rat MI
model using percutaneous injection catheters (Martens et al. 2009). Fibrin glue increased cell
retention and survival. More importantly, this study determined the viscosity limits for the
delivery of the cell-hydrogel suspension using several commercially available catheters. By
controlling the composition of fibrinogen and thrombin, the fibrin hydrogel could be tailored to
be compatible with the catheters. This also shows the potential of using catheter delivery for cell
transplantation therapies involving other cell types and hydrogels (Martens et al. 2009). While
fibrin-based biomaterials show promise and have the additional benefit of fibrin being FDA
approved for human use, these biomaterials may only provide short-term cardiac functional
23
benefits that decrease with time (Yu, Christman, et al. 2009; Nelson et al. 2011; Segers and Lee
2011).
Another natural hydrogel used for cell transplantation is a temperature responsive hydrogel
made of chitosan and glycerol phosphate similar to that described by Wang et al above (Lu et al.
2009). This hydrogel gels in approximately 10 to 15 minutes at physiological temperatures.
Embryonic stem cells were injected into the infarcted heart wall of rats in the chitosan hydrogel,
showing better cell retention at 24 hours and increased graft size, improved cardiac function,
thicker ventricular wall and greater micro-vessel density at 4 weeks as compared to injection of
the cells in phosphate buffered saline (PBS) (Lu et al. 2009). Histological staining of excised
heart sections showed a significant presence of chitosan at 24 hours after injection, while sparse
positive chitosan staining was found at 4 weeks. No trace of the hydrogel was found at 6 weeks
post-injection, showing complete degradation of the hydrogel in vivo (Lu et al. 2009). Chitosan
has also been explored for the delivery of adipose-derived mesenchymal stem cells into the
infarcted heart (Z. Liu et al. 2012). The chitosan hydrogel improved MI microenvironment by
recruiting SDF-1, which is a key chemokine for homing of stem cells, and scavenging reactive
oxygen species generated by ischemia that would otherwise impair adhesion molecules of the
delivered stem cells (Z. Liu et al. 2012).
Pre-treatment of animals with gelatin hydrogel microspheres incorporating bFGF was used to
improve transplantation of fetal rat CMs in rat MI models (Sakakibara et al. 2002). It has been
shown that bFGF can enhance viability of CMs through neovascularization and also directly
affect CMs by stimulating DNA synthesis, myocyte proliferation and differentiation during
development. In the experimental groups with bFGF microspheres alone and with bFGF
microspheres followed by CM transplantation, there was neovascularization in the scar tissue
after 1 week. Importantly, more transplanted cells survived in the scar area with the pretreatment
with bFGF microspheres compared to no pretreatment, in which transplanted cells only survived
in the peri-infarct regions (Sakakibara et al. 2002).
In another study, human embryonic stem cell (hESC)-derived CMs were delivered to the heart
within Matrigel™ using a cocktail of pro-survival factors (M. a Laflamme et al. 2007). The pro-
survival cocktail resulted in a 4-fold improvement in myocardial graft size relative to hESC-CMs
injected with Matrigel™ alone after 1 week. After 4 weeks, increased LV wall thickening and
reduced ventricular dilation, improved global function (fractional shortening and ejection
24
fraction), and attenuated changes in LV chamber dimensions were observed in hearts injected
with hESC-CMs along with the pro-survival cocktail relative to pro-survival cocktail-only,
buffer-only, or non-cardiac cell controls (M. a Laflamme et al. 2007).
An example of a synthetic hydrogel aimed at improving cell transplantation therapies to
reduce LV remodeling post-MI is an injectable α-cyclodextrin/poly(ethylene glycol)–b-
polycaprolactone-(dodecanedioic acid)-polycaprolactone–poly(ethylene glycol) (MPEG–PCL–
MPEG) hydrogel (T. Wang, Jiang, et al. 2009). Culturing bone marrow-derived stem cells
(BMSCs) in the hydrogel showed it was non-toxic and allowed for the maintenance of cell
morphology in vitro. In a rabbit LV MI model co-injection of α-cyclodextrin solution with
BMSCs and MPEG–PCL–MPEG solution in the infarct area 7 days post-MI showed immediate
gelation and localization (T. Wang, Jiang, et al. 2009). Histological analysis of excised heart
cross sections 4 weeks after injection showed the hydrogel was absorbed (degraded), and that the
hydrogel significantly increased BMSC retention and vessel density around the infarct compared
to injecting BMSCs alone. Functional echocardiography studies demonstrated an increased LV
ejection fraction (~77% increase) and attenuated left ventricular dilatation with hydrogel-BMSC
co-injection compared to PBS and BMSC only controls (T. Wang, Jiang, et al. 2009).
A synthetic injectable hydrogel made of oligo[poly(ethylene glycol) fumarate] (OPF) was
developed to deliver mouse embryonic stem cells into the LV wall of a rat MI model (H. Wang
et al. 2012). The hydrogel improved cell retention compared to delivery in PBS, and was
biodegradable with complete degradation by 6 weeks after injection in the infarcted heart. The
hydrogel alone significantly reduced the infarct size and improved cardiac function at 4 weeks
after injection (H. Wang et al. 2012). However, the combination of cells and hydrogel further
increased revascularization, improved cardiac function, decreased infarct size and fibrotic area,
as compared to hydrogel alone and cells injected in PBS. Importantly, the transplanted
embryonic stem cells expressed cardiovascular markers, such as cardiac troponin-T, von
Willebrand factor and α-smooth muscle actin. As such, the OPF hydrogel not only supported cell
transplantation by improving cell retention and survival, but the stem cells also underwent
differentiation into the cardiac linage, thus showing potential of this delivery method for heart
regeneration (H. Wang et al. 2012).
Synthetic hydrogels have also been modified by incorporation of bioactive peptides and
growth factors to improve their use for cell delivery. A synthetic PEG-based hydrogel modified
25
with matrix metalloproteinase (MMP) cleavable peptide domains was developed for the co-
delivery of Tβ4 and vascular cells (Kraehenbuehl et al. 2009). Tβ4 was physically entrapped in
the hydrogel, which gels in approximately 30 minutes at physiological conditions. Due to the
MMP cleavable peptide domains, the hydrogel was prone to cell-mediated proteolytic
degradation and the encapsulated Tβ4 could be released due to the MMP secretion of
encapsulated human umbilical vein endothelial cells in vitro. At the same time, the presence of
Tβ4 led to improved adhesion, survival, migration and organization of the cells (Kraehenbuehl et
al. 2009). In vivo, the MMP-responsive PEG hydrogel was used to co-deliver Tβ4 and hESC-
derived vascular cells into the infarcted myocardium of rats (Kraehenbuehl et al. 2011). By 6
weeks after MI, there was improved cardiac function (i.e. decreased end-systolic volume and
increased ejection fraction) in the group with vascular cells injected in Tβ4-encapsulated
hydrogel, as compared to treatment with PBS control, Tβ4-free hydrogel and Tβ4-encapsulated
hydrogel. This improvement was due to the better organization of native CMs and endothelial
cells in the combined Tβ4 and cell delivery, as supported by the presence of Tβ4 and the
secretion of human vascular cytokines (i.e. VEGF, EGF, HGF) and pro-survival factors (i.e.
survivin) by the transplanted vascular cells (Kraehenbuehl et al. 2011). They hypothesized that
the gel provided temporary support and pro-survival factors as it substituted for the degrading
endogenous matrix, and hESC-derived vascular cells contributed to formation of capillary-like
vessels, stabilization of host vessels, secretion of paracrine factors, or induction of paracrine
factor secretion from native rat cells (Kraehenbuehl et al. 2011). The injection of cell-free PEG-
hydrogel with Tβ4 improved cardiac function compared to PBS control, showing that the
injection of the bioactive hydrogel alone could also be explored for endogenous cardiac repair
(Kraehenbuehl et al. 2011).
Injection of a synthetic self-assembling hydrogel RAD16-II (initially developed by Zhang et
al. 1993) formed nanofiber microenvironments in the host myocardium and promoted
recruitment of endogenous ECs and vascular smooth muscle cells, which appeared to form
functional vascular structures, necessary for improving blood flow and improving the general
environment for CM survival. Furthermore, co-injection of the self-assembling peptide with
CMs showed survival (albeit low) of the transplanted cells, but significant increase in
recruitment of endogenous progenitors (M. E. Davis et al. 2005). The addition of insulin-like
growth factor-1 (IGF-1, a CM growth and differentiation factor) proved to be anti-apoptotic and
26
increased cell growth in transplanted (injected) CMs. It showed sustained, controlled, and
targeted release of IGF-1 from the nanofibers for 28 days. Improved systolic function was
observed when CMs were injected with the new functionalized hydrogel in a rat MI study (M. E.
Davis et al. 2006).
2.5.4. Hydrogels for the artificial maintenance of ventricle geometry and repair The third approach for preservation of cardiac function is the injection of a bulking material
into the ventricular wall to prevent progressive adverse remodeling due to high wall stresses that
develops after MI. Theoretical modeling suggests that hydrogels change the geometry and
dilation mechanics of the ventricle in such a way as to reduce the elevated local wall stresses,
and further that the mechanical support provided by the hydrogels improves the ejection fraction
and the stroke volume/end-diastolic volume relationship (Z. Li and Guan 2011; Wall et al. 2006).
To this end Fujimoto et al developed an N-isopropylacrylamide (NIPAAm), acrylic acid (AAc),
and hydroxyethyl methacrylate-poly(trimethylene carbonate) (HEMAPTMC) based synthetic
hydrogel (Fujimoto et al. 2009). It was found that a feed ratio of 86/4/10 poly(NIPAAm-co-AAc-
co-HEMAPTMC) monomers formed a hydrogel at 37°C, which gradually became soluble over a
5 month period, and furthermore that no degradation product cytotoxicity was observed in vitro
(Fujimoto et al. 2009). The hydrogel or PBS was injected into the MI zone two weeks after MI in
a rat model, and analysis was done 8 weeks after injection. In the PBS group, LV cavity area
increased and contractility decreased at 8wk post-MI, while in the hydrogel group both
parameters were preserved during this period, and furthermore the hydrogel was still present at 8
weeks (Fujimoto et al. 2009). Echocardiography was used to measure heart function and it was
reported that the hydrogel maintained fractional area change (contractility) at the same level
measured just prior to injection, and was ~55% better than the PBS group at 8 weeks post
injection (10 weeks post MI). Tissue in-growth was observed in the hydrogel injected area and a
thicker LV wall (~100% thicker) and higher capillary densities were found for the hydrogel
versus PBS group (Fujimoto et al. 2009).
Thermo-responsive materials that transform from solution to gel near body temperature are
frequently employed for tissue engineering because their transitional properties make them easy
to deliver and manipulate. T. Wang et al developed a series of hydrogels by introducing
hydrophobic PCL (polycaprolactam)-grafted polysaccharide chains into the thermo-responsive
poly(N-isopropylacrylamide) (PNIPAAm) network (T. Wang, Wu, et al. 2009). The resulting
27
hydrogel (Dex-PCL-HEMA/PNIPAAm) transitions from solution to gel within 30s at 37°C and
is reversible within the same time frame. To test the reparative effect of this hydrogel to
damaged cardiac function, 200µL was injected four days after MI in a rabbit model. One month
after the treatment, histological staining of heart sections showed a significant increase in scar
thickness and reduction in infarct size, compared with the control group (T. Wang, Wu, et al.
2009). Moreover, LV end systolic pressure was increased and end diastolic pressure reduced in
comparison with the MI control group. It was also observed that hydrogel implantation partially
reduced the dilatation of the LV, as shown by a significant increase in ejection fraction and a
significant decrease in LV end diastolic and systolic diameters (T. Wang, Wu, et al. 2009). These
results suggest that this hydrogel could serve as an injectable biomaterial that prevents LV
remodeling and dilation for the treatment of MI.
Similarly, Wu et al developed a temperature-sensitive, aliphatic polyester hydrogel (HG)
conjugated with VEGF and looked at its effects on cardiac recovery after MI (Wu et al. 2011). It
was shown that the hydrogel gelled in ~10 minutes at physiological conditions, and that in vitro
it was stable for up to 5 weeks, but completely degraded by 6 weeks when injected
subcutaneously. Injection around the infarct site was done seven days after coronary artery
ligation and rats were monitored up to five weeks post-injection (six weeks post-MI). Compared
with outcomes in the PBS group, ventricular volumes, preload recruitable stroke work, and end-
systolic elastance were all significantly preserved with HG-VEGF, and furthermore the VEGF
conjugated hydrogel led to improved blood vessel density in the infarct area (Wu et al. 2011).
Functionally, ejection fraction was ~60% higher, fractional shortening ~37% better with the HG-
VEGF group over the PBS injected group, and while the infarct thinned and dilated after PBS
injection it was ~30% smaller and 50% thicker in hearts treated with HG-VEGF (all measured at
5 weeks post injection) (Wu et al. 2011). They did note that most of these effects could be seen
with injection of hydrogel (no VEGF), or hydrogel with VEGF mixed in (not conjugated).
However, none of the effects were as significant as those observed with their VEGF conjugated
polyester hydrogel leading to the conclusion that conjugation provided sustained, localized
VEGF function promoting regeneration (Wu et al. 2011).
Furthermore, an important relationship on how hydrogel mechanical properties correlate their
therapeutic outcomes has been brought up and studied by Ifkovits et al (Ifkovits et al. 2010).
They compared two injectable methacrylated hyaluronic acid (MeHA) formulations that exhibit
28
similar degradation and tissue distribution upon injection but have differential moduli, using
infarct only as a control group. The MeHA or control hydrogels were injected 30 min after the
infarction and the result showed both treatments significantly increased the wall thickness in the
apex (>200% increase) and basal (>40% increase) infarct regions compared with the control
infarct (Ifkovits et al. 2010). Specifically, the treatment with the higher modulus (MeHA high)
group had significantly smaller infarct area (approximately 16% smaller) compared with the
control infarct group. Moreover, the normalized end-diastolic and end-systolic volumes were
reduced (though not significantly) for the MeHA High group (1.7 and 1.9, respectively)
compared to control (2.1 and 2.5, respectively) (Ifkovits et al. 2010). Functionally, real time 3D
echocardiography was used to assess the LV dimension and cardiac function for each animal
before and immediately after infarction, suggesting that MeHA group also tended to have a
better cardiac output and higher ejection fraction (again however, not significant) than the low-
modulus (MeHA Low) and control infarct groups (Ifkovits et al. 2010).
Some care must be taken in interpreting the results of bulking material strategies as recent
work suggest passive wall support on its own does not prevent LV remodeling and preservation
of cardiac function. Rane et al looked to decouple biomaterial effects from mechanical effects of
such treatments by injecting bio-inert, non-degradable PEG hydrogel into a rat MI model nine
days after infarction (Rane et al. 2011). While infarct wall thickness was significantly increased
in the injected group verses saline injected control animals, there was no difference in cardiac
function (ejection fraction, end diastolic and systolic volumes) between groups. Confirmation of
the decoupling between material bioactivity and mechanical effects was done by comparing
cellular response between groups and showing no difference, thus leading to the conclusion that
benefits reported from other studies are likely due to differences in inflammatory and cellular
response to the material and not due to the bulking effect of the material itself (Rane et al. 2011).
For an injected hydrogel to act as a long term mechanical support, it would be assumed its
presence long-term (slow or no degradation) is required. To investigate this, a non-degradable, in
situ polymerizable PEG hydrogel was assessed for its effect on short and long term cardiac repair
after MI. Chosen for its degradation properties (i.e. lack thereof) as well as its very high
mechanical stiffness and inertness, gel components were injected into the infarcted region and
polymerized in situ following permanent ligation of the left anterior descending artery in male
Wistar rats (Dobner et al. 2009). Benefits were seen after 4 weeks with increased wall thickness
29
(100%) and modestly attenuated dilation (43%) compared to a saline treated control group.
However, morphological improvements were only seen during the first 4 weeks as by 3 months
the hydrogel injection group had regressed to a cardiac geometry and function similar to the
control group (Dobner et al. 2009). The results suggest that while maintaining ventricle geometry
using stiff biomaterials is beneficial in the short term, successful long term application will most
probably require substantial further research toward developing a material with the necessary
characteristics and a greater understanding of the degree of stress relief needed to halt
pathological remodeling (Dobner et al. 2009).
As is evident, complex synthetic polymer hydrogels are generally the choice for ventricle
geometry preservation strategies due the high tunability of mechanical properties to match the
needs of the failing LV. This does not exclude groups from using natural polymers as well,
however. Dai et al investigated whether commercially available collagen injection could thicken
the infarcted LV wall and improve function by preventing systolic bulging in Fischer rats (Dai et
al. 2005). 100μL of collagen or saline was injected into the scar area of rats with week old MIs,
and then cardiac morphology and function was monitored a further six weeks. The results
showed significantly increased scar thickness in the collagen injected group (719±26μm)
compared with the saline-treated group (440±34μm), and also that injected collagen remained
present in the scar at six weeks. Furthermore, stoke volume was significantly larger in the
collagen-treated group (163±8μL) than in the saline-treated group (129±6μL), and LV ejection
fraction was greater in the collagen-treated group (48.4±1.8%) than in the saline-treated group
(40.7±1.0%) (Dai et al. 2005).
2.5.5. Clinical application The true implication of any of these injectable biomaterial strategies can only really be
assessed in a clinical setting. If the hydrogels are inert like alginate, they could be regulated as
devices, leading to a shorter time to reach the clinic. Two groups have developed alginate based
hydrogels currently in Phase II clinical trials (NCT01226563, & NCT01311791). The calcium-
cross-linked alginate hydrogel developed by Leor et al. was first tested in a swine anterior MI
model (Leor et al. 2009). Animals were monitored up to 60 days (56 days post-injection) and
control animals (intracoronary saline injection) showed an increase in LV diastolic area by 44%,
LV systolic area by 45%, and LV mass by 35%, whereas intracoronary hydrogel injection (2mL)
prevented and even significantly reversed LV enlargement by greater than 100% compared to
30
control for all measures. Furthermore scar thickness was 53% greater with 2mL hydrogel
injection compared to the control, and further proved that their developed biomaterial was
feasible, safe, and effective (Leor et al. 2009). Dubbed IK-5001, the hydrogel has been approved
for Phase II clinical trials and they are currently recruiting patients to test the safety and
effectiveness of the device. Intracoronary injection of 4mL of IK-5001 into the blocked artery
(after successful stent placement) will be done with the goal of prevention of ventricular
remodeling and congestive heart failure when administered after a recent acute MI
(NCT01226563).
By contrast, the hydrogel developed by Lee et al is designed for direct injection into the
infarcted wall for restoration of LV geometry. Called Algisyl-LVR™, it is another calcium-
cross-linked alginate biomaterial that gels in 3-4 minutes and achieves a material strength of 3-
5kPa (Lee et al. 2012). An initial pilot study in 6 patients suffering from dilated cardiomyopathy
demonstrated sustained improvements in LV size and function that were accompanied by
statistically significant improvements in clinical status and quality of life 3 months post-
injection, with no implant related complications (Lee et al. 2012). The approved Phase II clinical
trial will evaluate the concept that direct mid LV intramyocardial injections of the biomaterial
into the free wall of the failing LV of patients with dilated cardiomyopathy will reduce LV size,
restore LV shape, lower LV wall stress and improve global LV function (NCT01311791). The
goal is to show significantly improved peak maximum oxygen uptake 6 months after treatment
between Algisyl-LVR™ treated patients versus those receiving medical management.
2.6. Chitosan & collagen
Chitosan is a naturally occurring linear polysaccharide, composed of N-acetyl-D-glucosamine
and D-glucosamine monomers linked by β(1-4) glycosidic bonds, as seen in Figure 2-2. It is
derived from the deacetylation of chitin, the most abundant organic molecule and main
component of the exoskeleton of crustaceans, molluscs, and insects. It is biocompatible, and its
degradation products are nontoxic and non-immunogenic (Prasitsilp et al. 2000; Muzzarelli
1993). Its biodegradation products are lower molecular weight chitosans, chito-oligomers, and
the monomers (N-acetyl-D-glucosamine, itself an anti-inflammatory drug, and D-glucosamine).
Additionally, the degradation of chitosan through enzymatic hydrolysis by lysozyme can be
controlled by modifying the degree of deacetylation (I. Y. Kim et al. 2008).
31
Figure 2-2: The structures of chitosan and chitin, and their proportions in 85% deacetylated chitosan
The presence of amino groups on the polysaccharide backbone of chitosan allows for
conjugation with various bioactive molecules (Figure 2-2) (Shi et al. 2006; I. Y. Kim et al.
2008). Furthermore, the amino groups make chitosan poly-cationic, meaning it can adhere more
easily to negatively-charged tissue (Chenite et al. 2000), and can be used for the controlled
release of negatively-charged molecules (Fujita et al. 2004). However, chitosan suffers from
mechanical weakness and instability, as well as an inability to maintain predefined shapes (Puppi
et al. 2010). Chitosan is not a novel material but thanks to the diversity of potential uses it has
been well studied and utilized in many biomedical applications including wound repair, tissue
engineering, and drug delivery (as discussed above and reviewed extensively in Q Li et al. 1992;
Majeti and Kumar 2000; IY Kim et al. 2008).
Collagen I is a poly-anionic natural protein polymer and is a major constituent of the cardiac
extracellular matrix. Currently, collagen based materials are one of the most common ECM
materials used for culturing cells in vitro, and form the backbone of many hydrogels used in vivo
(as described earlier) as it is the most characterized and abundant component of the ECM
(Johnson, Lin, and Christman 2011). Native fibrous collagen can be processed into an injectable
solubilized form, which can self-assemble under specified conditions into a hydrogel with
nanoscale fibrous structure (Johnson, Lin, and Christman 2011). Collagen molecules assemble to
form microfibrils, which, in turn, form fibrils to create collagen fibers. Collagen can be
reconstituted to form hydrated gels that are similar to loose connective tissue in vivo, with cell
behavior in three dimensional hydrated collagen gels being more typical of in vivo behavior than
those grown on two-dimensional plastic culture surfaces. The material properties—material and
gelation kinetics, stiffness, nanoscale fibrous architecture—and cell-matrix interactions of
32
collagen gels have been well characterized and can be manipulated by changes in temperature,
ion concentration, phosphate content, and pH to specify the properties of the hydrogel. It is
desirable as a biomaterial since it is biodegradable, has attachment sites for many cell types, and
has a mild inflammatory response when implanted (Pulapura and Kohn 1992).
Limitations in its use for cardiac tissue engineering, however, are its mechanical properties
and suitability for culturing cardiomyocytes, as well as its rapid biodegradation (Puppi et al.
2010). Numerous methods for improving the mechanical properties of collagen gels have been
explored including addition of chemical cross-linkers, but the effect of the by-products of these
chemical reactions can face scrutiny from regulatory and clinical bodies (Nicodemus and Bryant
2008). In contrast, in the natural ECM the structural properties of collagen are known to be
modified by glucosaminoglycans. The role they (glucosaminoglycans) play in modifying
collagen fibril formation is by intertwining the microfibrils, forming thicker fibrils, and thus
improving structural properties and increasing pore size. The structural characteristics of
chitosan make it very similar to glucosaminoglycans, thus it was proposed and shown that
addition of poly-cationic chitosan improves the mechanical stability and compressive strength of
the highly poly-anionic collagen based gels through ionic interactions between the two
components (Tan, Krishnaraj, and Desai 2001). Furthermore, the addition of chitosan to collagen
gels has been shown to reduce the rate of degradation by collagenase (H. Chen et al. 2008). The
interactions between collagen and chitosan in polyelectrolyte complexed hydrogels is illustrated
in Figure A-1 (Appendix A).
Collagen-chitosan hydrogel mixes have shown positive results by other groups as well. The
work of Lu et al suggested a thermo-responsive chitosan based hydrogel increased
vasculogenesis in rat infarcts when co-injected with ESCs in comparison to phosphate buffered
saline (PBS) control (Lu et al. 2009). Based on this study, chitosan was added to an already
established collagen hydrogel from another group, forming a chemically cross-linked collagen-
chitosan hydrogel aimed at improving recruitment, differentiation and survival of vascular
progenitors (Suuronen et al. 2006; Deng et al. 2010). While Lu et al were focused primarily on
CM regeneration (they were successful in improving ESC survival and differentiation), the work
by Deng et al showed that addition of chitosan stabilized the collagen hydrogel, was better suited
for the maturation of endothelial cells (in vitro), and promoted greater vascular growth and
recruited more endothelial/angiogenic cells than collagen alone (Lu et al. 2009; Deng et al.
33
2010). Addition of the oligosaccharide sialyl LewisX (sLeX, which binds to the adhesion
molecule L-selectin) to the hydrogel improved mobilization and recruitment of endogenous and
transplanted circulating progenitor cells. Furthermore it enhanced neovascularization, and
improved tissue perfusion in a rat hind-limb ischemia model (Suuronen et al. 2009).
2.7. QHREDGS peptide
The angiopoietins are a family of growth factors that are known to bind to the receptor
tyrosine kinase Tie2 acting as both agonists and context-dependent antagonists (S. Davis et al.
1996; Maisonpierre 1997). In endothelial cells, angiopoietin 1 (Ang1) was found to act as a
survival factor by activating the phosphoinositide-3-kinase (PI3K)/Akt anti-apoptotic pathway
(Hayes et al. 1999; I. Kim et al. 2000; Papapetropoulos 2000). It was later found that fibroblasts
lacking the Tie2 receptor could still bind to Ang1 and that the resultant adhesion could be
inhibited by blocking integrin receptors (Carlson et al. 2001). This led to the discovery that the
αV integrin subunit, and more specifically the integrin receptor αVβ3, was responsible for
angiopoietin-mediated attachment (independent of Tie2), which induced endothelial cell
adhesion, migration, and activated characteristic integrin signal transduction pathways
(Camenisch et al. 2002). It has also been reported that while neonatal rat cardiomyocytes (NCM)
do not express the Tie2 receptor, they could still adhere to Ang1; adhesion to Ang1 conferred a
significant survival advantage to NCMs; Akt and mitogen-activated protein kinase (MAPK)
pathways were involved in the Ang1-mediated survival of CMs; and finally adhesion to Ang1
was integrin-dependent (Dallabrida et al. 2005). Further study identified the short sequence
QHREDGS, conserved among mouse, rat, human, and other species, as the integrin-binding
motif in the fibrinogen-like domain of Ang1, and it was demonstrated that the peptide
QHREDGS was capable of supporting CM attachment and survival similar to full-length Ang1
(Figure 2-3) (Rask, Mihic, et al. 2010; Rask, Dallabrida, et al. 2010). While chitosan on its own
does not promote cell attachment previous work has shown the ability of the peptide QHREDGS
to promote CM attachment and growth, as well as survival and maturation when covalently
attached to a photo-crosslinkable Az-chitosan hydrogel (Rask 2009; Rask, Mihic, et al. 2010;
Rask, Dallabrida, et al. 2010). Although a well known integrin ligand, RGD, also promotes
attachment of CMs, it prevents caspase-3 activation, it decreases the force of contraction of
papillary muscle, and leads to pathological hypertrophy of CMs (Boateng et al. 2005;
34
Balasubramanian and Kuppuswamy 2003). The work previously described with Az-chitosan also
showed the superiority of QHREDGS to an RGD based peptide, specifically RGDS.
Figure 2-3: Molecular structure of Ang-1 derived peptide QHREDGS
2.8. Peptide modified chitosan-collagen hydrogel for cardiac tissue engineering
As it is known chitosan can interact with collagen to improve the mechanical stiffness of
collagen based gels it is proposed that a mixture of the two components can form a hydrogel
suitable for the culture of CMs in vitro as well as be mechanically stable enough for injection
into the heart for treatment of acute MI. The QHREDGS peptide sequence has been
demonstrated to promote the attachment and survival of CMs, suggesting it might have the
therapeutic potential to help restore cardiac contractile function post MI. Moreover, since the
beneficial effects of the QHREDGS peptide were found to be attachment dependent, we
postulated that immobilizing the peptide onto a hydrogel would promote its pro-survival activity
in addition to ensuring its sustained localization at the site of injection. Importantly, as a peptide,
QHREDGS has an advantage over the protein-based therapies in that it is water-soluble, very
stable, fully-synthetic with a precisely defined composition, it can be produced in a cost-effective
and facile manner, and does not require a specific orientation to be functional.
As discussed earlier, the presence of amino groups on the backbone of chitosan allows for its
conjugation with the carboxyls of bioactive molecules. Conjugation can be achieved through
many routes; however one of the most common methods for stable cross-linking of molecules is
35
the use of carbodiimide compounds as they provide the most popular and versatile method for
labeling or cross-linking to carboxylic acids. The most readily available and commonly used
water soluble carbodiimide is 1-ethyl-3-(-3-dimethylaminopropyl) carbodiimide hydrochloride
(EDC) (Hayworth 2014). EDC reacts with carboxylic acid groups to form an active, unstable O-
acylisourea intermediate that is easily displaced by nucleophilic attack from primary amino
groups in the reaction mixture. The primary amine forms an amide bond with the original
carboxyl group, and a soluble urea derivative is released as the EDC by-product. Failure of the
intermediate to quickly react with an amine results in hydrolysis of the intermediate, regeneration
of the carboxyls, and the release of an N-unsubstituted urea (Hayworth 2014). The efficiency of
the reaction can be greatly improved with the addition of sulfo-N-hydroxysuccinimide (Sulfo-
NHS). Sulfo-NHS replaces EDC during the reaction to form a Sulfo-NHS ester that is
considerably more stable than the O-acylisourea intermediate, yet is still highly reactive with
primary amines at physiologic pH. The Sulfo-NHS is regenerated upon successful conjugation of
the carboxyl and amine, and it along with other reaction byproducts can subsequently be
removed and the products purified through methods such as dialysis or chromatography. The
modified chitosan-QHREDGS product is illustrated below (Figure 2-4), with further
modifications used seen in Figure A-2 (Appendix A). A full schematic of the reaction and
involved components can be seen in Figure A-3 (Appendix A).
Figure 2-4: QHREDGS peptide modified chitosan (QHG213H)
36
3. Hydrogel development, in vitro characterization & preliminary in vivo
models
The original work described in this chapter was published in (Reis et al. 2012), and is used
with permission from Elsevier.
3.1. Materials & Methods
All animal experimental procedures were approved by the Animal Care Committee of the
Toronto General Research Institute and the University of Toronto Committee on Animal Care,
according to the Guide for the Care and Use of Laboratory Animals.
3.1.1. Peptide modified chitosan (UP-G113-QHREDGS)
3.1.1.1. Conjugation
QHREDGS peptide was conjugated to chitosan using EDC chemistry in a manner similar to
that previously described for conjugation to Az-chitosan (Rask, Dallabrida, et al. 2010). To
assess the dose response of the peptide two levels were used, a Low and a High concentration of
the peptide in proportion to the amount of chitosan present. Briefly, chitosan (UP-G113,
Novamatrix) was dissolved at 20 mg/mL in 0.9% normal saline and peptide at 10 mg/mL in
phosphate buffered saline (PBS, Lonza). These were then mixed with EDC and N-
hydroxysulfosuccinimide (S-NHS) dissolved in PBS to obtain final reaction solutions of 5
mg/mL chitosan and 0.5 mg/mL (Low) or 3 mg/mL (High) peptide with the ratio of
[EDC]/[peptide] and [S-NHS]/[EDC] kept constant in the two reaction mixtures at 0.8 and 2.75,
respectively.
The reaction solution (1.5 mL) was then left on a vortex mixer (VWR) at 650 RPM for 3
hours. The solutions were diluted 4X with PBS and dialysed (using a dialysis membrane,
Spectra/POR MWCO 3500, Spectrum Labs) against two 2 L changes of distilled water for ~24
hours.
Sterilization of the dialyzed reaction mixture was done by passing the solution through a 0.2
µm syringe filter (Progene) and then the material was recovered through lyophilisation for
approximately 48 hours. The recovered material was stored at -20°C until use.
3.1.1.2. Assessing conjugation efficiency
37
Fluorescently labelled peptide, Fitc-Ahx-QHREDGS was used to assess the conjugation
efficiency and true final concentration of peptide attached to chitosan post dialysis. Fitc-labelled
QHREDGS (Biomatik) was substituted for regular peptide in the protocol above (3.1.1.1) and all
steps were protected from light. As the molecular weight cut-off of the dialysis membrane is at
most a tenth that of chitosan it can be safely assumed that all of the chitosan is retained and
recovered, and the peptide present is only that which was successfully attached to the chitosan.
Standards of the FITC-Ahx-QHREDGS in PBS was made ranging from 0.0005 to 0.01
mg/mL. For both the standards and the reaction solutions recovered post dialysis, the pH was
adjusted to 7, as fluorescence is greatly affected by pH (Graber et al. 1986). The samples (Low &
High) were diluted a further 1:10 and 1:100 and then run, with the standards, through a
fluorometer (Spectra Max Geminin EM, Molecular Devices) at an excitation wavelength of 490
nm and emission of 520 nm, with all samples run in triplicate. The true final concentrations of
peptide and conjugation efficiency were calculated by comparing the fluorescence of the samples
to the standards, correcting for the dilution factor and for the volume recovered post dialysis.
3.1.2. Chitosan-collagen hydrogel
3.1.2.1. Hydrogel formulation
Chitosan-collagen hydrogels were made through modification of the standard protocol for
collagen gelation from BD Biosciences. Briefly, pure or peptide-modified chitosan was dissolved
in sterile 0.9% normal saline and mixed with 10X PBS, collagen (BD Biosciences), and 1N
NaOH on ice, in that order. The final solution consisted of 2.5 mg/mL chitosan, 2.5 mg/mL
collagen, 10% of the final (solution) volume of 10X PBS (V10XPBS=0.1*VTotal), and 2.3% the
volume of collagen added of 1N NaOH (VNaOH=0.023*VCollagen). The final hydrogel solution was
mixed thoroughly and kept on ice until needed. An appropriate amount of hydrogel solution was
pipetted onto a tissue culture plate bottom, or mixed with cells first, and allowed to gel for 30
minutes in a humidified 5% CO2 37°C incubator. To make 0:1 or 3:1 mass:mass (m:m)
chitosan:collagen hydrogels the final solution was the same as above, but with 0 mg/mL or 7.5
mg/mL chitosan, respectively.
3.1.2.2. Scanning electron microscopy
Samples were imaged using environmental scanning electron microscopy (Hitachi S-3400 N).
Hydrogel samples (N=2/group) were allowed to gel for 30 min and individually placed into the
38
specimen chamber. A filter paper was used to gently remove the excess water from the sample.
The chamber was closed and the temperature of the chamber was decreased to -20°C. The
samples were imaged under variable pressure mode at 70 Pa and 15 kV.
3.1.2.3. Hydrogel degradation
To assess the in vitro degradation of the hydrogel 2.0 mL Eppendorf tubes were weighed after
having 5 holes punched in the caps with an 18G1/2 needle. 500 μL of 2.5 mg/mL 1:1 m:m
chitosan-collagen hydrogel was then added to each tube and allowed to gel as previously
described. Culture media (1 mL) was added to each tube and incubated for different amounts of
time, with media changes (1 mL removed and replaced) every 48 hours. Gels were left for 1, 72,
120, or 240 hours (N=3/time point) when 1 mL of media was removed and the tubes and gel
frozen at -80°C. All groups were lyophilized together for 72 hours, and then weighed again.
Degradation was assessed by comparison of the lyophilized material weight to that of the 1 hour
group.
3.1.2.4. Rheological assessment of hydrogels
To assess differences in the mechanical stability of hydrogels with varying ratios of chitosan
present, as well as the effect of conjugating QHREDGS peptide to chitosan, a rheological
assessment measuring the loss and storage moduli of varying hydrogel compositions was
performed. Rheology was performed using a TA Instruments AR1000 rheometer with a 6 cm
acrylic cone and plate geometry and calibrated as per the manufacturer’s instructions prior to
seeding approximately 1 mL of appropriate hydrogel (0:1, 1:1, 3:1, or 1:1 with conjugated
QHREDGS peptide). Temperature was controlled using an integrated Pelletier plate and
equilibrated to 4°C and prior to hydrogel seeding. To determine the Linear Visco-elastic Range
(LVR), and show gel transition, a strain sweep was done at 4°C from 0-5% strain and frequency
of 1Hz, and after allowing the hydrogel to equilibrate at 37°C for 30 minutes. Second, from the
data collected from the strain sweeps a frequency sweep was performed on samples from all
formulations at 37°C. The frequency sweeps were performed by seeding hydrogel onto the
rheometer at 4°C, raising the temperature of the plate to 37°C and gelling for 30 minutes. A total
of N=3 samples per group per analysis (strain or frequency sweep) were performed.
3.1.3. In vitro cell culture
3.1.3.1. CM isolation
39
Cardiomyocytes used for in vitro work were isolated from neonatal (1-2 day old) Sprague-
Dawley rat hearts as previously described (Radisic et al. 2003; Naito et al. 2006; Rask,
Dallabrida, et al. 2010). Briefly, rats were euthanized; hearts were removed, quartered and then
digested overnight with trypsin followed by five rounds of collagenase digests. Cells were then
pre-plated for 1hr to enrich for CMs, and pre-plate supernatant was collected. CM number in the
pre-plate supernatant was determined using a haemocytometer, and then cells were used for
experiments. Quantification of the cell population before and after pre-plating to enhance the CM
fraction has been performed previously, and it is reported after pre-plating the cell fractions are
63±2% CM, 33±3% cardiac fibroblasts, 3-4% smooth muscle cells, and 2-3% endothelial cells
(Naito et al. 2006). A second pre-plate can be performed to improve the CM fraction to ~80% if
desired (Iyer, Chui, and Radisic 2009).
3.1.3.2. CM media
The cardiomyocyte culture media was comprised of 10% (v/v) FBS, 1% Hepes, 100 U/mL
penicillin-streptomycin, 0.02 U/mL Insulin, 5 μg/mL vitamin C, and the remainder Dulbeco’s
Modified Eagle medium (10% F.I.V. media). Cells were always re-suspended prior to
encapsulation in a serum rich media, specifically 30% F.I.V (30% FBS).
3.1.3.3. CM encapsulation & culture
Unless otherwise specified each sample prepared for in vitro experiments consisted of 5x105
CMs re-suspended in 2.5 μL 30% F.I.V. media and encapsulated in 12.5 μL of appropriate gel
prepared as described in 3.1.2.1. The gel, media, and CM mix was pipetted onto the surface of a
24-well tissue culture plate and put in an incubator for 30 minutes to allow for gelation, after
which samples were supplemented with 500 μL warm 10% F.I.V. media and cultured for 120
hours (five days). Media was aspirated, or collected, and 100% changed every 48hrs (two days).
3.1.4. In vivo studies
3.1.4.1. Subcutaneous injection
CMs from neonatal Lewis rat pups were collected in the same manner as described in 3.1.3.1,
spun down and 2 million CMs were re-suspended in 10μL 30% F.I.V. media first, followed by
the addition of 100μL appropriate gel (Control (No Peptide), Low, and High) in Eppendorf tubes.
The cell/gel suspension was then randomly injected subcutaneously (100 μL) with a 23G1/2
needle into the back of adult female Lewis rats (four injections per rat). The area where samples
40
were injected had been shaved and rough injection sites were marked with permanent marker in
order to locate the injected material later. Samples (N=6/group) were left for seven days and then
the animals were sacrificed, the skin surrounding the area where injections were made was
removed and the material (hydrogel with cells) recovered and fixed with 10% formalin at 4°C for
approximately 12hrs. Recovered samples were then transferred to PBS and sent to the Pathology
Research Program (PRP) histology lab at University Health Network for paraffin-embedding and
sectioning.
3.1.4.2. Mouse MI model
To assess the potential of the developed hydrogel to be injected into the infracted heart, a
mouse MI model was used. Adult male C57 Black-6 mice were subjected to a left ventricular
anterior descending coronary artery ligation procedure (LAD procedure) to mimic severe MI
with the help of Dr. A. Momen from the University Health Network. The mouse MI model has
become important for studying myocardial regeneration following interventions such as cell
based therapies, and studies have shown the LAD procedure gives rise to MIs of reproducible
size and severity as well as an immediate surgical survival of 60%, and 2 week survival of ~83%
(or 50% overall survival) (Kumar et al. 2005). Animals were divided into three groups: Sham,
MI Only, and MI+Control gel. Briefly, animals were initially anaesthetized with 5% isoflurane
and maintained with 2-2.5% isoflurane during surgery. Animals were intubated and ventilated
using a Minivent ventilator (Harvard Apparatus, March-Hugstetten, Germany) at ~200
breaths/min. A left thoracotomy was performed, and the left coronary artery was ligated using a
7-0 silk suture (Ethicon) passed with a reverse cutting needle. Sham mice did not receive the
LAD ligation and the chest was closed immediately. MI+Control gel mice had 50µL of
2.5mg/mL 1:1 m:m chitosan:collagen Control (No Peptide) gel injected with a 27G1/2 needle
into the area immediately below the ligation suture. The hydrogel was warmed to ~37°C for
~10min prior to injection to allow gelling to initiate. The animals were removed from the
ventilator and allowed to recover below a heat lamp for approximately 30 minutes. (N=3 animals
for MI, N=3 for MI+Control Gel, N=1 for Sham).
3.1.5. Construct Characterization
3.1.5.1. Gel compaction
41
To assess the chitosan-collagen ratio that was most appropriate for formation of beating
cardiac constructs, CMs were encapsulated in the same manner as described in 3.1.3.3, with
groups consisting of CMs in 0.5:1, 1:1, or 3:1 m:m chitosan-collagen hydrogel, with the
concentration of collagen kept at 2.5 mg/mL. The cell/hydrogel mix was pipetted onto 12mm
circular glass cover-slips located in the wells of a 12-well tissue culture plate. This was done to
obtain uniform initial construct sizes and prevent cell/gel attachment to the bottom of the plates,
which would confound results. After 144 hours (6 days) of incubation (with media changes every
48 hours) the media was removed and pictures of the hydrogel constructs taken (N=4/group).
Adobe Photoshop CS3 was then used to measure the area of each hydrogel sample, and mean
values compared between groups.
3.1.5.2. Live/Dead staining
Live/dead staining was performed according to the manufacturer’s protocol (Invitrogen) using
5-carboxyfluorescein diacetate acetoxymethyl ester (CFDA, green for live cells) and propidium
iodide (PI, red for dead cells) on N=2 constructs/group. CFDA is a non-fluorescent compound
that can penetrate intact cell membranes wherein it is cleaved by non-specific esterases to release
polar fluorescent fluorescein which is retained inside viable cells, whereas PI is a nucleic acid
probe that is unable to cross intact cell membranes but penetrates the membranes of dead cells.
Quantification was performed using ImageJ as previously described (Iyer, Chiu, and Radisic
2009). To show live cell distribution confocal images (z-stacks) were taken at the edge and
center of cell/gel constructs.
3.1.5.3. Functional testing
Two measurements were taken to assess the electrical function of encapsulated
cardiomyocytes in chitosan-collagen hydrogels (No peptide, Low, and High peptide) cultured for
five days as described in 3.1.3.3. Excitation threshold (ET) is the minimum voltage required to
pace most of the cells in a construct and the maximum capture rate (MCR) is the maximum
frequency at which the encapsulated CMs could be induced to beat simultaneously (Radisic et al.
2004). N=11 constructs/group were cultured as described above and then ET/MCR were
measured using bi-phasic electrical pulses as previously reported (Rask, Mihic, et al. 2010).
Briefly, media was removed, a pair of parallel carbon electrodes were placed surrounding the
gels, and the system was immersed in warm Tyrode’s solution (Sigma, pH≈7.3). All
42
measurements were done in an environmental chamber at 37°C and ET was measured by
increasing the voltage until ~90% of cells in the field of view were seen to be beating
synchronously with square biphasic pulses of 2 ms and frequency of 1 Hz. MCR was measured
by setting the voltage at 12 V increasing the frequency until most of the cells were no longer
beating synchronously with the driving signal.
Success rate was measured as the percentage of constructs for which ET/MCR measurements
were able to be taken (i.e. constructs with significant beating portions) to the total number of
constructs seeded for that group. Chi-square and Fisher’s Exact tests were performed to analyze
the association between groups.
3.1.5.4. XTT assay
CMs were cultured for five days as per 3.1.3.3 at which time media was removed and
replaced with fresh 10% F.I.V. media and XTT assay solution as per the manufacturer’s protocol
(XTT cell viability assay kit, Biotium, Inc.). Samples were incubated for a further three hours
with the mix and then three 100 μL samples of mix were taken from each hydrogel sample and
put in corresponding wells of a 96 well plate. The absorbance of wells was read using an
absorbance meter (Apollo LB911, Berthold Technologies) at a wavelength of 450 nm and a
reference of 620 nm. The triplicate readings per hydrogel were averaged to get a value for the
sample, and then all sample values were averaged to get a value for the group. A total of N=6 per
group were tested. In addition, N=2 per group of no-cell control gels were used. The mean
absorbance reading of the no-cell controls was subtracted from all test groups. A standard curve
was also made to show that XTT is indeed a good indicator of cell metabolism and indirectly cell
number by encapsulating varying numbers of CMs in the same volumes of serum rich media and
No-QHREDGS Control hydrogel and incubating for 1 or 120 hours. Relative absorbance is
reported, with mean absorbance of samples relative to that of constructs with no CMs. Higher
relative absorbance indicates a higher level of viability/metabolic activity.
3.1.5.5. LDH assay
The LDH cytotoxicity assay (Cayman Chemical Company) was used to assess cell death at
varying time points. It was performed as per the manufacturer’s protocol on media samples taken
every 24 hours from samples prepared as per 3.1.3.3. Briefly, 50 μL media samples (in triplicate)
were put into the wells of a 96 well plate and an equivalent volume of LDH reagent was added to
43
each well. The plate was placed on an orbital shaker for 30 minutes at 600 rpm protected from
light and then the absorbance of each well read using an absorbance meter (Apollo LB911,
Berthold Technologies) at a wavelength of 492 nm and reference of 620 nm. A total of N=6
gels/group plus 2 gels/group of no-cell controls were assessed. The absorbance of the no-cell
controls was subtracted from the readings of CM samples to get a more accurate value. A cell
standard was made in the same manner as for the XTT assay, and the LDH assay performed on
media samples from the standard (N=4/group) as well as the LDH standard.
3.1.5.6. PicoGreen DNA Assay
Constructs cultured as per 3.1.3.3 for four days were subsequently immersed in 1 mL of lysis
buffer (0.2% Triton X-100, 200 mM Tris-HCl, 20 mM EDTA, pH 7.5) and transferred to 2 mL
Eppendorf tubes. Several autoclaved 1 mm diameter silica beads were added to each vial, and the
constructs were homogenized using a Mini-Beadbeater-16 (Biospec), via six cycles of ten
seconds each. 500 μL of each homogenate was then transferred to separate wells of a 24-well
plate, and the Picogreen assay was performed as directed by Invitrogen. Briefly, 500 μL of
Picogreen reagent was added to each well, and incubated in the dark for 3-5 minutes, after which
fluorescence was read using a fluorescence micro-plate reader (Spectra Max Gemini EM,
Molecular Devices; excitation 480 nm, emission 520 nm). A separate assay, which was done on
gels lysed immediately after encapsulation with a known quantity of cells, served as a control to
determine the cell number in gels at the end of the experiment.
3.1.5.7. Histological & immunofluorescent staining
Samples for immunohistochemical and immunofluorescent staining were fixed at 4°C with
10% neutral buffered formalin (Sigma, HT501129) for approximately 12 hrs, and then
transferred to PBS. For paraffin embedding and sectioning, as well as hematoxylin and eosin
(H&E), Mason’s trichrome, smooth muscle actin (SMA), Factor VIII (F8), and CD31 staining,
the fixed samples were sent to the Pathology Research Program (PRP) at the University Health
Network. Chitosan staining was performed using Cibacron Brilliant Red-3BA (CBR-3BA,
Sigma Aldrich) and Weigert’s Iron Hematoxylin as described (Rossomacha, Hoemanni, and
Shive 2004). Paraffin sections were also stained for cardiac troponin T (cTnT, mouse
monoclonal antibody, Thermo Scientific MS-295-P). Deparaffinised slides were blocked using
Normal Horse Serum (Gibco 16050-122) for 40 minutes and then primary antibody (cTnT) was
44
applied at a dilution factor of 1:200 for at least 12 hrs at 4°C. Secondary antibody solution
(Alexafluor 488 donkey anti-mouse IgG (H+L), Invitrogen A21202) was applied at a dilution of
1:200 with DAPI (Sigma) nuclear stain at 1:200 dilution for ~40 min at room temperature. The
slides were mounted using Fluoromount aqueous mounting medium (Sigma-Aldrich, F4680),
covered with cover-slips (VWR, 22x50 mm) and imaged.
Sections stained for CD3 and vimentin were deparaffinised and rehydrated with subsequent
baths (3 of each) of 100% xylene, 100% EtOH, 95% EtOH, 75% EtOH, and distilled water for 3
minutes per change. Antigen retrieval was performed by microwaving slides immersed in TRIS-
EDTA buffer for 5 min, and blocked using Dako Serum Free Protein Block (Dako X0909) for 30
min at room temperature. Primary antibody (poly rabbit anti-human CD3 diluted 1:200, Dako
A0452 or mouse monoclonal anti-vimentin diluted 1:400, Sigma V6630) was applied to sections
for 2 hrs at room temperature followed by secondary antibodies (goat anti-rabbit FITC IgG
(H+L), Jackson ImmunoResearch 111-095-144 or Alexafluor 488 donkey anti-mouse IgG
(H+L), Invitrogen A21202) at 1:400 dilution for 1 hr at room temperature. DAPI (Sigma) nuclear
stain was applied at 1:1000 dilution for 10 minutes and finally slides were mounted using Dako
Fluorescence Mounting Medium (Dako, S3023) and imaged.
Quantification of area covered by positively stained cells in immunostained sections was
performed using Adobe Photoshop CS3 and ImageJ updated from a previously reported method
(Chiu and Radisic 2010). Up to five distinct images per section per sample were gathered. The
colour profiles were saved such that the same colour selection profile could be applied
automatically to every image to select for positive staining (removing subjective human staining
selection). Unselected image areas were deleted and the image was converted to black and white
and saved as a JPEG. Adjusted images were then opened in ImageJ, thresholded to remove
potential background, and then analyzed to count particles. The resultant output was Area
Fraction of black particles (positive stain) in relation to the whole image which gives a relative
measure of positive immunostaining. The algorithm was tested on images from every group to
ensure consistency, colour profile refined as necessary, and then applied to all images. The five
images per section were averaged to obtain a % Positive Area measure for the sample, and then
sample values averaged to obtain Mean±SD for each group.
The number of F8+ vessels and their diameters were determined using ImageJ according to a
previously published method (Chiu and Radisic 2010). Five randomly selected 300x300pixels
45
(131x131mm2) areas were cropped and magnified from each original F8 image. The numbers of
F8+ vessels were manually counted from the cropped images and normalized to the area of the
image to determine the vessel density. The diameters of the vessels were evaluated by drawing a
line across the short axis of the vessel and then measuring the length of the line in ImageJ.
3.1.6. Statistical analysis Statistical analysis was performed using SPSS Statistics 17.0 and GraphPad Prism 5.0.
Differences between experimental groups were analyzed by using one-way ANOVA with post-
hoc Tukey tests or two-way ANOVA with Bonferroni post-tests. Categorical data was compared
using Chi-square or Fisher’s Exact tests. P<0.05 was considered significant for all statistical
tests. Results were plotted with GraphPad Prism 5.0, with all data being reported as mean ±
standard deviation (Mean±SD).
3.2. Results & Discussion
3.2.1. Hydrogel composition and characterization
3.2.1.1. Base collagen-chitosan hydrogel
The first obstacle in designing a hydrogel as a vehicle for cell transplantation to the heart is
demonstrating its suitability for cardiomyocyte culture as well as an appropriate gelation time to
retain transplanted cells/material in the site of injection. Using previous work conducted in
optimizing a chitosan-collagen hydrogel for a hematopoietic cell line, similar concentrations and
ratios of the two components were chosen as starting points (Tan, Krishnaraj, and Desai 2001).
Various initial concentrations of collagen varying from 1.5-3 mg/mL were qualitatively
assessed for structural stability once gelled as per the manufacturer’s protocol (BD Biosciences).
It was determined that concentrations below 2 mg/mL were not “gelled” enough, whereas high
concentrations of around 3 mg/mL provided a much more stable gel it severely limited the stock
concentration of chitosan required to get the necessary ratios (needed to be on the order of 100
mg/mL, which is unfeasible). As such, a collagen concentration of 2.5 mg/mL was chosen and
ratios of 0:1, 1:1, and 3:1 chitosan:collagen m:m gels were made to assess differences in gelling
time and remodelling when cardiomyocytes were encapsulated (see 3.1.2.1 & 3.1.3.3).
No apparent differences in gelling time between the three groups (0:1, 1:1, & 3:1) were
observed, all forming nice, stable gels in 30 min at 37°C, however the differences in remodelling
that took place during cell culture were marked, as seen in Figure 3-1 A. The initial size of all
46
gels used for gel compaction studies was 9 mm. Assuming gels formed a uniform cylinder with a
diameter of 9 mm and volume of 12.5 μL then the theoretical average initial height of the
constructs was 197 μm. The 0:1 group compacted almost six times more than the 3:1 group after
6 days of culture (which remained significantly unchanged from its initial size, data not shown),
and almost four times more than the 1:1 gel group (P<0.001 between all groups). The 3:1 group
gels, however where very brittle and broke apart easily with any outside disturbance, while both
the 0:1 and 1:1 group maintained their form after handling.
Cardiomyocytes require a matrix soft enough that they can remodel, yet stiff enough to resist
compaction from the remodelling as well as contraction of the cells. Our previous studies
demonstrated that substrates of stiffness comparable to the native heart maintains the
differentiated phenotype of cardiomyocytes and prevents fibroblast overgrowth in monolayer
cultures of heart cells (Bhana et al. 2010). If a matrix is not stiff enough to maintain its shape and
fully compacts, as seen in the 0:1 group, then proper elongation and maturation of the CMs is not
possible. Consistently, beating constructs were observed in the 1:1 group, not at all in the 0:1
group, and only in small portions of the 3:1 group (Supplemental Movie 1&2, from Reis et al.
2012).
The elongation and ability of the construct to contract in the 1:1 group but not in 0:1 or 3:1
groups after 6 days of culture (both spontaneous and induced) point toward the health and
performance of the cells (images and videos not shown).
Confirmation of the results seen in the in vitro cell culture work performed was obtained
through rheological assessment of the varying hydrogel formulations (seen in Figure 3-1 C-E).
At 4 °C all hydrogels are in a liquid state, as seen in Figure 3-1 C with the loss modulus (G11) of
all groups being greater than that of the storage modulus (G1) for all groups throughout a strain
sweep from 0-100%. All formulations show gel-like behavior after 30 minutes at 37°C (Figure
3-1 D), with a good LVR up to strains of at least 5%. The differences in the magnitude of loss
and storage moduli can be seen between groups, with the 1:1 group showing the highest storage
modulus and 3:1 the lowest. This observation is even more apparent in the results from the
frequency sweeps performed on samples of all formulations with the same trend of 1:1 having
the highest storage modulus and the 3:1 group the lowest, and similarly seen with the loss
modulus although the differences are not as pronounced (Figure 3-1 E). It should be noted that in
many instances at 4°C moduli values could not be obtained at low strains due to the sensitivity of
47
the rheometer, that is why there are clusters of readings some points. Likewise, due to the brittle
nature of the 3:1 gel the rheometer had difficulty measuring the loss and storage moduli at 37°C
for some samples. It should also be noted that time varying viscosity rheology measurements
were performed, however it was found that in initiating gelation through increasing temperature
from 4-37°C caused an immediate change over from liquid (G11> G1) to gel (G1>G11)which is
not a practical measure for our application. Furthermore, during the elastic storage modulus (G1)
never equilibrated during these experiments to give a meaningful measure of time to gelation.
Mechanical properties of collagen-chitosan hydrogels were studied previously (L. Wang and
Stegemann 2011; Tan, Krishnaraj, and Desai 2001). The initial increase in storage modulus and
subsequent decrease with further addition of chitosan is likely due to the marked changes in the
gel microstructure (Figure 3-1 D&E). The resultant structure of collagen only gels is a highly
interconnected fibrous network with fibers of fairly uniform size. Addition of chitosan does not
seem to affect fiber formation, however they interfere with the resultant network by sequestering
fibers together and thus forming thicker “struts”, yet retaining a high level of interconnectivity in
the structure (i.e. 1:1 gels). By bringing together many collagen fibers along a chitosan
backbone, formation of these thicker struts results in an increase in mechanical integrity.
Increasing the fraction of chitosan (i.e. 3:1 chitosan:collagen gels) causes a further increase in
collagen fiber sequestration (thus thicker struts), however the degree of connectivity between
resultant fibers is greatly diminished (see Appendix A, Figure A-1). Furthermore, Tan et al
(2001) found previously that in the case of a human hematopoietic cell line, changes in cell
growth were determined by chitosan:collagen ratio only, and that increasing or decreasing the
total protein concentration (while keeping the ratio constant) had no effect on cell proliferation
(Tan, Krishnaraj, and Desai 2001).
Protein based gels are commonly tested at strain and frequency values on the order of 1% and
1 Hz respectively (L. Wang and Stegemann 2011; Ikeda and Foegeding 2003; Whittingstall
2003). As these values fall in the LVR for all chitosan:collagen formulations tested here they
were chosen as a basis to compare the loss and storage moduli between gel formulations. The
values are summarized in Table 3-1, below, and as expected confirm that a 1:1 collagen:chitosan
gel formulation provides the most mechanically stable gel in terms of storage modulus, but also
shows the greatest difference (separation) between loss and storage moduli (P<0.0001).
48
Based on these findings it was determined that a 2.5 mg/mL 1:1 m:m chitosan:collagen gel
formulation provided the most suitable environment of the three tested for cardiomyocyte
culture, and all further work done used this formulation.
Table 3-1: Hydrogel rheological assessment
Hydrogel Storage
Modulus±SD (Pa) Loss
Modulus±SD (Pa)
ANOVA Multiple Comparisons (Sig.)
0:1 1:1 3:1 1:1-P
0:1 (Collagen Only)
32.32±7.55 4.20±1.16 <0.001 <0.001 0.014
1:1 (Control (No Peptide))
54.87±9.05 6.05±0.90 0.006 <0.001 0.119
3:1 11.63±3.46 2.45±0.66 0.010 <0.001 <0.001
1:1-HP (High Peptide)
45.37±4.29 5.53±0.45 0.073 1.000 <0.001
49
Figure 3-1: Hydrogel Characterization (A) Chitosan:collagen blends resist gel compaction by the heart cells significantly better than collagen alone
after 6 days in culture. (2.5 mg/mL collagen concentration, 1:1 or 3:1 mass ratio of chitosan:collagen). Images show cell/gel constructs after 6 days in culture. Gel sizes in the graph are normalized to the 0:1 group. N=4/group. (B) 2.5 mg/mL 1:1 chitosan-collagen m:m ratio shows no significant degradation over 10 days in vitro. N=3/group. (C-E) Rheological analysis of various ratios confirms 1:1 has mechanical properties suitable for
50
CM culture. (C) All gel formulations were liquid at 4°C. (D) All gel formulations exhibit linear viscoelastic behaviour gel properties between 0.05 and 5% strain after 30 minutes at 37°C. (E) 1:1 chitosan-collagen mixture has the highest storage modulus (G’) at all frequencies (with constant strain at 1%) after gelling for 30 minutes at 37°C. (C-E) Conjugation of QHREDGS peptide to chitosan does not alter the rheological properties of the 1:1 gel formulation. N=3/group C-E. All data presented as Mean±SD; ***P<0.001.
A study of the degradation of the hydrogel over the time course of in vitro experiments was
done to assess how much the material degraded on its own (hydrolytic and in response to the
components of culture media) as described in 3.1.2.3. The results are seen in Figure 3-1 B and
show no significant difference in dry material weight at all time points up to 10 days of exposure
to culture medium. A control of just Eppendorf tubes with culture media and no hydrogel was
used to demonstrate there was no inherent changes in dry tube mass over the course of the
experiment, as they could have an effect on the very small material masses measured (only 6-8
mg). It should be noted that the 1 hr group shows a dry weight close to the expected 5 mg total
for 1 mL of hydrogel but this increases slightly, not significantly, over the course of the
experiment and may be attributed to absorption of media components by the hydrogel.
3.2.1.2. Collagen-chitosan-QHREDGS hydrogel
The amount of QHREDGS peptide successfully bound to chitosan and the efficiency of the
process was evaluated using fluorescently labelled peptide, Fitc-Ahx-QHREDGS, as described in
3.1.1.2. The efficiency of peptide conjugation is affected by the ratios of EDC and S-NHS to
peptide in the reaction mixture, and as this study has been performed on Az-chitosan the ratios
used for the protocol described in 3.1.1.1 were kept the same as those previously reported (Rask,
Mihic, et al. 2010; Rask, Dallabrida, et al. 2010). Furthermore, the minimum level of peptide
needed to elicit a cellular response was characterized in the same study, thus target amounts of
QHREDGS peptide to be conjugated were kept in the same range. The results of the current
study are seen in Figure 3-2 A. The differences in conjugation efficiency of both the Low and
High amounts of peptide were not statistically significant and average to be 55%, the same as
that found in Rask, Mihic, et al. (2010). It follows that there was statistically significant
(P=0.002) differences in the concentration of peptide successfully conjugated, with the Low
group being 0.053±0.016 mg/mL and the High 0.346±0.004 mg/mL in post dialysis solution (or
100±30 nmol/mL-gel and 651±8 nmol/mL-gel for the Low and High final gel formulations
respectively), as seen in Figure 3-2 A.
SEM images of 2.5 mg/mL 1:1 hydrogels with varying levels of QHREDGS peptide are seen
in Figure 3-2 B and were done to identify differences in hydrogel morphology with the addition
51
of peptide. The images show that the gels have a porous structure, which is desirable for
encapsulation systems to allow for nutrient exchange and that the addition of peptide causes no
apparent differences in structure. This indicates, qualitatively at least, that addition of
QHREDGS to chitosan does not affect its structure, or ability to interact with collagen and form
a stable hydrogel.
Quantitative assessment of the effect peptide conjugation has on the hydrogel was done using
rheology as in 3.1.2.4. The results of strain sweeps at 4°C and 37°C and frequency sweeps at
37°C, seen in Figure 3-1 , illustrate transition from liquid to gel after 30 minutes at 37°C and no
significant difference in the loss or storage moduli of peptide conjugated hydrogel to Control (no
peptide) 1:1 collagen:chitosan hydrogel at constant 1% strain and frequency of 1Hz (see Table
3-1). The 1:1 chitosan:collagen blend with and without the peptide had a storage and loss
modulus similar to the blend of 25% cardiac ECM and collagen I used in previous studies to
increase maturation of cardiomyocytes derived from human embryonic stem cells (Duan et al.
2011).
Therefore it can be concluded that peptide conjugation does not affect the morphology nor
mechanical properties of the hydrogel and thus any differences in cell survival, maturation, and
performance within the gel should be due entirely to the varying amounts of QHREDGS peptide
present.
It should be noted that the current hydrogel preparation does not allow us to decouple
mechanical properties of the hydrogel from the composition. There are other more elegant
chemistries that allow for independent control of mechanical properties and concentration of
presented ligands (Deforest, Sims, and Anseth 2010). In addition, using identical EDC chemistry
and covalent immobilization of VEGF to collagen, we previously demonstrated that the
biomolecule was stably immobilized to the substrate with negligible biomolecule release in PBS
over 28 days (Chiu and Radisic 2010; Miyagi et al. 2011). Thus it can be expected that the
QHREDGS peptide will also be stably retained within the hydrogel during culture.
52
Figure 3-2: Characterization of peptide conjugation and hydrogel morphology (A) Fluorescent analysis using Fitc-QHREDGS shows no difference in conjugation efficiency of peptide to
chitosan and significant difference in the amount of peptide between Low and High groups. (B) Representative SEM images show no visible differences in landscape between groups. Results are shown for 2.5mg/mL 1:1 chitosan:collagen gel (Control) and 1:1 chitosan-QHREDGS:collagen with Low and High amount of the peptide immobilized.
3.2.2. In vitro studies with cardiomyocytes in collagen-chitosan-QHREDGS hydrogels
3.2.2.1. Distribution, viability, and metabolism of CMs encapsulated in the hydrogels
For in vitro studies the same amount of cell/gel mixture as above (12.5l) was pipetted onto
the surface of 12 well TC plates and spread to form uniformly sized constructs of 6.2mm±0.3mm
in diameter (N=4/group, p=0.07 between groups). These gels had a theoretical height of
413μm±43μm. The encapsulated cells were distributed evenly in three dimensions in the cross-
section immediately after gelation and at 5 days (120 hrs) of culture in the hydrogels (Figure
53
3-3). There was very minimal leakage of cells during gelation, and very little migration of cells
out of the gels after 5 days (Figure B-1, found in Appendix B). It was therefore assumed that
effects of cells outside of the gels could be ignored and all assays were performed in the wells.
Figure 3-3: Gel encapsulation results in uniform distribution of live cells Confocal cross-sectional images of CFDA stained (green) cell/gel constructs (A) 1 hr post seeding (30
minutes post gelation) and (B) 120 hrs post seeding. All scale bars=50 µm.
54
The XTT assay is a measure of cell viability based on metabolic activity and was assessed for
cells encapsulated in the various gels (Control (no peptide), Low, and High peptide). The results
are seen in Figure 3-4 A&B. Figure 3-4 A shows the applicability of XTT in measuring cell
viability/activity when encapsulated in the hydrogel (Control gel), indicating a low amount of
metabolic activity 1 hr after encapsulation and increased, significantly different activities
between varying groups (numbers of cells encapsulated) after 5 days in culture. When comparing
the three test groups (Control, Low, & High) directly as seen in Figure 3-4 B cell metabolism is
significantly higher (P<0.01) in the High QHREDGS group in comparison to the Control, but not
the Low group. There is an overall trend however pointing toward increased metabolic activity
with increased peptide concentration in the gel. A question arises in interpreting these results as
directly pointing toward increased viability in the High group as the cell population used to seed
the hydrogels is not a pure (non-proliferative) CM population, but contains fibroblasts and
endothelial cells as well. While there is no doubt that metabolic activity is indeed higher in the
High group it is possible that this is due to an increased level of fibroblast and endothelial cell
proliferation, and not CM activity.
Cell death at varying time points during culture of encapsulated CMs was assessed using an
LDH assay which measures cytotoxicity based on the activity of lactate dehydrogenase (LDH)
released from damaged cells. The assay was performed, as described in 3.1.5.5, to determine if
there was any difference in the survival/apoptosis of cells in the three gel test groups, and also to
give an indication of whether the results of the XTT assay could be due to differences in cell
number. Figure 3-4 C shows that discernable differences in LDH activity could be measured
from varying numbers of cells encapsulated in the hydrogels, and therefore the LDH assay could
be used to compare the Control, Low and High gel groups. The results of directly comparing
groups (Figure 3-4 D) show a general trend toward lower cell death (LDH activity, Control to
Low to High) at all time points measured, and there is a statistically significant difference in cell
death between Control and High groups over the course of the experiment (two-way ANOVA,
P<0.05). It should be noted that data was collected up to 120 hrs (5 days) however there was
negligible cell death in all groups after 72 hrs. This is consistent with previous reports on the
cultivation of heart tissues in well plates showing the majority of cell death occurs early on in
culture, within the first 24 hr (Radisic et al. 2003). Collectively, these data show the effect of
55
QHREDGS peptide in improving cell survival, and correlate with the results of the XTT assay as
lower cell death leads to higher cell metabolic activity and viability.
Finally, total cell number at the end of five days in culture was assessed using the Picogreen
DNA assay which uses an ultra-sensitive fluorescent nucleic acid stain to quantify double-
stranded DNA. Performed as described in 3.1.5.6, the strong correlation between initial cell
number encapsulated and fluorescence (total DNA) shows the hydrogel does not affect the assay
(Figure 3-4 E) and that the assay can be used to assess the differences in cell number between
test groups. The results of comparing the three groups (Figure 3-4 F) show no significant
differences in total cell number initially (data not shown) and after 120 hrs in culture. These data
confirm that the results seen in both the XTT and LDH assays were not due to differences in cell
number encapsulated (higher or lower respectively), that cell proliferation does not play a major
role, and thus point toward the improved health of CMs encapsulated in gels with QHREDGS
and specifically with the High concentration.
When the three assays, LDH, DNA and XTT are taken as a whole it is clear that there is no
appreciable proliferation of non-myocytes in this system. Higher XTT activity in the high
QHREDGS group indicates improved metabolic activity of all cells.
56
Figure 3-4: Metabolic activity and total number of cells encapsulated in 1:1 chitosan:collagen hydrogels
(A) Cell metabolic activity is significantly lower at 1Hr post encapsulation compared to 120Hrs after. (N=4/group for 1hr, N=3/group for 120hrs). (B) Cell metabolism in the Control (No QHREDGS) gels is significantly lower than in the High (3mg/mL QHREDGS) hydrogel group but not in the Low (0.5mg/mL QHREDGS) hydrogel group based on XTT assay (N=6/group). Data is relative to control gels with no cells. (C) LDH assay shows correlation between final cell number and cell death at all time points (N=3/group for ½ hr group, N=4/group for the remaining time points). (D) LDH assay shows significantly lower cell death in the High (3 mg/mL QHREDGS) compared to Control (No QHREDGS), but not in the Low (0.5 mg/ml QHREDGS)
57
hydrogel group (N=6/group/time point). (E) Picogreen DNA assay shows strong correlation between initial total cell number encapsulated and fluorescence (N=3/group). (F) Picogreen DNA assay shows no difference in total cell number between all groups after 120 hrs in culture (N=4 Low & High, N=3 Control; initial cell number=5·105 CMs). All data presented as Mean±SD, * P<0.05, ** P<0.01, *** P<0.001.
3.2.2.2. Construct functional & morphological properties
Native heart tissue responds and paces to electrical stimulation and this ability is usually
associated with healthy tissue. Therefore CMs encapsulated in the three hydrogel groups were
tested to determine their excitation threshold (ET) and maximum capture rate (MCR), as
described in 3.1.5.3. The results of measuring the ET/MCR of cells encapsulated for 120 hrs are
illustrated in Figure 3-5 A and it is seen that there is no significant difference between the
functional properties between groups. There does appear to be a trend toward lower ET and
higher MCR when comparing the Control and High peptide gels, however. The ET and MCR
values of healthy neonatal rat heart tissue were previously measured and are 3.2±1.5 V and
5.1±1.1 Hz respectively (Dengler 2009). While ET of all groups was about double that for native
tissue, the MCR was very close to what is seen in the native heart.
There was a significant difference in the success rate for obtaining beating constructs Figure
3-5 B. Spontaneous beating was consistently observed throughout the entire gel in the High
QHREDGS peptide group, while whole gel synchronous beating was only observed in the
Control group ~50% of time. These observations are consistent with the ~100% success rate for
inducing synchronous contractions in the High QHREDGS group compared to ~55% in the
Control (Figure 3-5 B, P<0.05). When whole gel synchronous beating was not observed there
were no significant local colonies of beating cells (please see L. A. Reis et al. (2011)
Supplementary Movie 1 & Supplementary Movie 2).
It should be noted that there are tissue engineering methods (e.g. electrical/mechanical
stimulation during cultivation) that could potentially improve the functionality of the cultivated
constructs, however since the end goal of this work is to design a hydrogel for cell delivery and
not creating a cardiac patch only these basic assessments of functionality were performed here
(Radisic et al. 2004; Wolfram-Hubertus Zimmermann, Melnychenko, and Eschenhagen 2004).
58
Figure 3-5: 1:1 Chitosan:collagen hydrogel with immobilized QHREDGS enables cultivation of beating cardiac tissue
(A) ET and MCR show no significant difference between groups after 120 hrs of cultivation (N=11/group, data presented as Mean±SD). (B) There is a statistically significant association between success rate and group, with the High group generating a significantly higher percentage of beating constructs than the Control (*P<0.05). (C) Representative images of paraffin-sectioned constructs with CMs encapsulated in gels and grown in vitro (N=4/group). Live/Dead (PI, red/CFDA, green); Hematoxylin & Eosin staining; Immunostaining for CD31 (brown), Vimentin (green) and cTnT (green) with DAPI counterstained nuclei shown (blue). Scale bars: 20µm H&E, CD31 and vimentin; 0.1mm cTnT.
Samples of constructs grown in vitro for five days were fixed in 10% formalin and sent to be
paraffin-embedded, sectioned, and stained for either Hematoxylin & Eosin (H&E) or cardiac
troponin T (cTnT) (see 3.1.5.7). Live/dead staining was also performed on other samples of
healthy constructs and imaged, as per 3.1.5.2. The results of all the staining are seen in Figure
3-5 C. There were no apparent differences in the Live/Dead (Figure 3-5 C & Figure B-2 A,
Appendix B) or H&E stained sections amongst the groups (Figure 3-5 C), and quantitative image
analysis confirmed no differences in cell viability as assessed by CFDA/PI staining (Figure B-2,
Appendix B). However, the viability was significantly lower in the middle of the hydrogel
compared to the edges (Figure B-2, Appendix B) when hydrogels were cultivated in the tissue
59
culture plates. This observation is consistent with oxygen diffusional limitations (Radisic et al.
2003). Definitive quantitative conclusions about viability and cell number were made from the
XTT, LDH, and Picogreen assays as due to the 3D nature of the constructs there was
considerable fluorescent background in the Live/Dead images, thereby possibly skewing results
(3.2.2.1).
The presence of endothelial cells was confirmed by CD31+ immunostaining in all groups
(Figure 3-5 C). While ECs were rounded in the Control gel and the Low peptide group, they
exhibited clearly elongated morphology in the High peptide group (Figure 5C). Immunostaining
for vimentin, which indentifies all non-myocytes (i.e. endothelial cells and fibroblasts together)
did not indicate appreciable differences in the presence of these non-myocytes in different
groups (Figure 3-5 C). Cardiac troponin T is a thin filament protein which takes part in muscle
contraction and is expressed and forms myofibrils in healthy cardiomyocytes. In sample images
taken for cTnT staining (Figure 3-5 C) there were larger areas of cTnT positive, elongated cells
forming contractile units in the High and Low groups, and much less in the Control group. High
magnification images indicated the presence of cross-striated cardiomyocytes in the High
QHREDGS group (Figure B-3, Appendix B).
3.2.3. In vivo studies
3.2.3.1. Subcutaneous injection
Subcutaneous injection of CMs and hydrogel was performed as an initial experiment to gauge
the immune response towards the injected biomaterial as well as determine how well the
hydrogel would localize at the site of injection. Prior to in vivo injection, the collagen:chitosan
mixture was pre-gelled for 10 min. In this state the gel was still injectable through a syringe and
needle thus it could be delivered either subcutaneously or into the heart and remain capable of
gelling and localizing at the delivery site. While photocrosslinkable gelation is faster, the low
penetration depth of UV light (~1.5mm) limits applicability of this gelation mechanism in larger
animals (Elisseeff et al. 1999). The recovery rate of the material injected under the skin in the
backs of the adult female Lewis rats was 50% after one week in vivo; five samples were
recovered from the Control group, three from the Low group, and four from the High group. The
unrecovered samples may have moved as the translucent nature of the material matched the
subcutaneous environment into which it was injected. We have studied the degradation rate of a
similar base gel (collagen:chitosan) in vivo using biotin labeling and 1 week subcutaneous
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implantation (Chiu and Radisic 2011). It was determined that there was no appreciable
degradation of the gel after 7 days in vivo which is the time-frame consistent with the in vivo
subcutaneous study described here (Chiu and Radisic 2011). The recovered samples exhibited
localized areas of concentrated collagen deposits in contrast to the surrounding tissue, as
identified by the darker blue areas of Mason’s trichrome stained sections (Figure 3-6 A).
Confirmation that these areas where indeed injected hydrogel was done through chitosan
staining, and H&E staining of nodule sections showed no apparent differences in tissue
morphology between samples, seen in Figure 3-6 B & Figure 3-7 A respectively. There was clear
infiltration of smooth muscle actin (SMA) and Factor-VIII (F8) positive cells into the injected
hydrogels (Figure 3-6 C&D respectively), however significant vascular formations could not be
identified. Quantification of the area of positive F8 and SMA staining was done and significant
differences could only be identified in the area of positive SMA staining between the Control
and High groups (Figure 3-6 E). The infiltrating SMA+ cells were most likely reparative
myofibroblasts.
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Figure 3-6: Subcutaneous injection study
(A-D) Histological staining of subcutaneously injected hydrogel samples with encapsulated Lewis rat neonatal CMs, recovered 7 days post injection. Full recovered nodules imaged with (A) Mason’s trichrome stain
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with areas containing the recovered sample highlighted with black box and confirmed with (B) chitosan staining (positive=dark pink/red), indicated with black arrows. (C) SMA staining indicates a higher expression in the High gel group compared to the Low and Control gel groups while (D) Factor VIII staining shows no major differences between groups. Scale bars=200µm A, 250µm B-D. (E) SMA and Factor VIII positive staining displayed as percent positive pixels. Significant differences seen between Control and Low vs High in the percentage of SMA positive area. Statistical analysis performed by One-way ANOVA with Bonferroni post-tests. *P<0.05, N=5 Control, N=3 Low, N=4 High. Data expressed as Mean±SD.
In the subcutaneous model, after 1 week, there was no appreciable angiogenic effect of the
peptide as assessed by F8 and CD31 immunostaining coupled with identification of tubular
capillary structures, although there were more CD31+ cells in the High group compared to the
Low and Control peptide group (Figure 3-7 B). Vimentin staining, identifying all non-myocytes,
was similar in all groups (Figure 3-7 C). cTnT staining of recovered samples proved the ability
of the hydrogel to retain cells at the site of injection, and further that the injected CMs survived
and matured to a point where they were forming contractile apparati (Figure 3-7 D&E). T-cells
play an important role in the immune and inflammatory response to biomaterials through
infiltration, recruitment of other immune cell types, resorption, fibrosis and ultimate rejection of
foreign bodies motivating the investigation of their presence in the injected hydrogel (van Luyn
et al. 1998). Rare CD3+ T-lymphocytes were identified in all gel groups upon sub-cutaneous
implantation indicating a minimal immune response (Figure 3-7 F).
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Figure 3-7: Immunostaining for different cell populations in subcutaneously injected nodules after 7
days in vivo (A) H&E staining shows no apparent differences between groups. Scale bars=250µm. (B) CD31
immunostaining identifies endothelial cells and (C) Vimentin immunostaining identifies non-myocytes:
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fibroblasts and endothelial cells. Scale bars=20µm B, 10�m C. (D, E) Expression of cTnT (green) is seen only in the center of nodules correlating with the positive Chitosan staining in Figure 7B, and CMs expressing cTnT are seen in all groups (E). (F) CD3 immunostaining identifies rare recruited T-cells (lymphocytes). Scale bars=10µm. Images correspond to the sections presented in Figure 3-6.
3.2.3.2. Mouse MI model study
Demonstrating the hydrogel’s ability to support CM viability, localize after injection
subcutaneously, retain cells at the site of injection, and promote survival upon transplantation
were all key aspects in its development. However, it was necessary to also demonstrate that the
hydrogel was capable of gelling and localizing at the site of infarct in the beating heart with its
current formulation, and to do this a mouse LAD model was used, as described in 3.1.4.2. Two
weeks post MI surgery the mouse hearts were collected and the site of infarct identified as the
area directly below the ligation suture, seen in Figure 3-8 A. Immediately noted was that both the
MI Only and MI+Control gel hearts appeared very much dilated in comparison to a control Sham
operated heart. Tissue cross-sections encompassing the MI area showed the extreme remodelling
that occurs post MI , with both MI groups showing significant tissue loss and collagenous scar
tissue deposition (Figure 3-8 B&C). Chitosan staining done on paraffin embedded cross-sections
confirmed the gels ability to localize at the site of injection in the infarct area (Figure 3-8 D).
Rare CD3+ T-cells were also identified in both the MI Only and MI+Control gel groups,
indicating injection of Control gel does not seem to significantly alter lymphocyte immune
response (Figure 3-8 E). CD31 positive endothelial cells were seen in the infarct area in both MI
Only and in the injection location in the MI+Control group (Figure 3-9 A). Immature vasculature
and infiltration of SMA expressing cells is seen in Figure 3-9 B&C, identified by F8 and SMA
staining respectively. The most pronounced difference was identified in the amount of SMA
expressing cells present, most likely reparative myofibroblasts, in comparison to the Sham heart.
No difference was found in LV wall thickness and quantification of LV blood vessel
diameter and density showed no significant difference between groups (MI Only and
MI+Control, Figure 3-9 D&E).
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Figure 3-8: Chitosan:collagen gel is suitable for injection into the infarcted heart
A left anterior descending coronary artery ligation mouse MI model was used. (A-D) Heart gross morphology and histology 2 weeks post surgery. (A) White arrows indicate ligation site identified by a suture and
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(B) infarct area. Injection was done in the area just below the suture immediately following ligation and confirmed with histological staining and immunostaining (C-E). (C) Trichrome images show tissue remodeling and extent of infarct; Scale bars=500µm. Black outlined sections are area of focus for (D and E). (D) Positive chitosan staining is seen in the MI+Control gel group only (black arrow). Scale bars=250µm. (E) Rare CD3+ T-cells are observed in the MI Only and MI+Control Gel Groups. Scale bars=10µm.
It should be noted at this point that there were many limitations to this study, especially in the
number of animals, and further quantification of heart performance pre- and post-MI and with
and without intervention must be done. The purpose of this feasibility study was to evaluate the
applicability of the chitosan:collagen blend for injection into the heart and it did not include the
injection of the QHREDGS peptide modified hydrogels. Detailed studies evaluating cell
injection with the hydrogel and functional improvements at six weeks, the time when
pathological remodeling post MI is completed in both rat and mouse models (Krzemiński et al.
2008; Patten et al. 1998), must be pursued in future studies.
Injection of biomaterials to the infarcted heart may exert several beneficial effects that
attenuate the extent of pathological remodeling. These include mechanical stabilization of the
left ventricle, delivery of bioactive molecules and delivery of reparative cells. To assess the
mechanical effects of biomaterial injection and determine where exactly to make such injections,
Wall et al (2006) performed finite element modeling studies with ovine left ventricle (Wall et al.
2006). They concluded that multiple injections in the border zone surrounding an infarct reduced
end systolic fibre stress proportional to the fractional volume of material added. Thus, hydrogel
injection therapies may have important beneficial effects on cardiac mechanics and can reduce
pathological remodeling post-MI irrespective of the addition of cells or cytokines (Wall et al.
2006). On the same note, Fujimoto et al (2009) set out to design a hydrogel to reduce wall stress
and bulk up the left ventricle (LV) post MI in order to reduce the extent of pathological
remodeling (Fujimoto et al. 2009). They designed a synthetic, thermo-gelling, biodegradable
hydrogel co-polymer from N-isopropylacrylamide, acrylic acid and hydroxyethyl methacrylate-
poly(trimethylene carbonate) (Fujimoto et al. 2009). Showing no in vitro toxic degradation
products they injected the gel into multiple sites along the MI border zone in a rat model. Results
showed preservation of LV cavity area and contractility, and tissue in-growth into the injected
hydrogel, none of which was seen in a PBS injection control (Fujimoto et al. 2009).
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Figure 3-9: Vascularization and wall thickness in the mouse MI model (A) CD31,(B) Factor VIII, and (C) SMA staining show appreciable differences between MI and MI+Control
gel groups in comparison to the Sham group. Scale bars=20µm A, 250 µm B&C. (D) LV wall thickness was measured from Trichrome images to assess attenuation of pathological remodeling with injection of Control gel. (E) Factor VIII positive staining was used to identify vasculature and neovascularization. No significant differences were identified between the groups at 2 week time point in this feasibility study. N=3 per group for MI Only and MI+Control Gel, N=1 Sham. Data expressed as Mean±SD.
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In our previous work we used photo-cross-linkable chitosan, Az-chitosan, modified with the
novel cell-protective peptide QHREDGS (Rask, Mihic, et al. 2010; Rask, Dallabrida, et al.
2010). While the photo-cross-linkable biomaterial may be applicable for treatment of small
animals such as mice and rats, the UV penetration depth may limit the applicability of photo-
cross-linkable hydrogels in delivery to porcine or human hearts. We thus focused on a hydrogel
blend that cross-links according to a combination of thermal gelling and ionic interactions
between positively charged chitosan and negatively charged collagen. As such this hydrogel has
a larger scope of applicability and it is suitable in catheter delivery as well. The gelation is
performed at mild conditions, cross-linker free, thus is suitable for encapsulation and delivery of
many cell types. In comparison to the photo-cross-linkable gelation which occurs in 1-2 min, the
blend explored here required appreciably longer time (30 min) to gel at 37°C. However, the
delivery and localization to the heart was possible by pre-gelling the blend for ~10 min prior to
injection.
The availability of amine groups on the chitosan backbone enables covalent immobilization of
different biomolecules. Covalent immobilization of ligands is desired over controlled release if
the effects of the ligand are to be observed in the biomaterial rather than the surrounding tissue.
Additionally, covalent immobilization was reported to prolong receptor/ligand signaling (Ito
2008; Fan et al. 2007). In addition to QHREDGS, angiogenic growth factors could also be
covalently immobilized to enhance revascularization of the infarct area. We have successfully
used covalently immobilized VEGF in collagen scaffolds to enhance angiogenesis in a right
ventricular free wall repair (Miyagi et al. 2011). Wu and colleagues have developed a
temperature-sensitive aliphatic polyester hydrogel conjugated with vascular endothelial growth
factor (VEGF) showing that it provided localized and sustained VEGF function, attenuated
adverse cardiac remodeling and improved cardiac function after injection in a rat MI model study
(Wu et al. 2011).
Even though the current feasibility study was not conclusive a number of conclusions can be
made from it. First, the developed hydrogel is capable of localizing at the site of injection in a
beating heart, a result that is undoubtedly the most significant as it means no further components
must be added to the formulation to improve mechanical properties or reduce gelling time.
Subsequent to this observation is that the hydrogel does not fully degrade in vivo at two weeks,
therefore future work must include studies of the in vivo degradation rate of the hydrogel.
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Secondly, while the base hydrogel formulation (1:1 Control- No Peptide) does not appear to
improve heart function post MI it at least does not cause any further damage to the infarcted
mouse heart (in animals that survived the procedure) and is biocompatible. This is significant as
any hydrogel, scaffold, etc. destined to be implanted into the body must not cause (further)
damage, or alternatively diminish improvement, or else a lack of intervention would be the best
option. Finally, this model provides a basis for assessing the performance of hydrogel with
conjugated QHREDGS peptide, which should hopefully show significant improvement in
comparison to both no treatment and Control gel, and therefore puts into context the lack of
improvement seen in the current mouse study.
3.3. Summary
SA-1 Develop a chitosan-collagen hydrogel that supports CM viability and phenotype.
Conjugate QHREDGS and demonstrate dose dependent improvement in CM phenotype and
function in vitro.
These data indicate that a 2.5 mg/mL 1:1 m:m chitosan:collagen blend gels without addition
of exogenous cross-linkers in ~30 minutes at 37°C. Furthermore, this formulation provides the
appropriate stiffness and mechanical properties to support cardiomyocyte survival both in vitro
and in vivo. Addition of the peptide did not significantly affect the rheological properties or
porous structure of the hydrogel. XTT assay and success of obtaining beating constructs
demonstrated clear effects of the incorporation of the QHREDGS peptide on cell viability and
functional properties in vitro. In vivo, High QHREDGS gel attracted more reparative
myofibroblasts and improved the presence of cardiomyocytes in the subcutaneous model. The
developed hydrogel has been shown to successfully localize at the site of injection both
subcutaneously and at the site of infarct in a mouse MI model. Further studies are required to
determine hydrogel degradation rate in the infarcted heart as well as functional improvements
upon injection of the QHREDGS modified hydrogels with and without cells.
Nonetheless, the work done represents that toward completion of Specific Aim SA-1, and the
further animal models described above is the focus of Specific Aim SA-2 and therefore the focus
work described in Chapter 4.
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4. Utility of QH-G213-H hydrogel as a treatment for acute MI
The original work described in this chapter has been published in Circulation Heart Failure
(Reis et al. 2015), and is used with permission from Wolters Kluwer Health.
Having developed a hydrogel formulation that performed very well in vitro and was showing
promise as a delivery vehicle for both cells and cytokines/growth factors in our subcutaneous in
vivo work, it was both logical and necessary to move toward an in vivo animal MI model. As
described in 3.2.3.2, in order for the developed hydrogel to be determined sufficient in its current
formulation it would have to prove capable of being injected and localizing at the site of infarct
within the beating heart. As the mouse LAD procedure is widely accepted as a good MI model
and the surgeon (Dr. Abdul Momen, University Health Network) had vast experience performing
the MI surgeries in mice we proceeded with a study to determine the effect of the developed
hydrogel in vivo. Our preliminary mouse MI model study design was for a total of 25 animals.
However, the severe trauma induced by the LAD surgery, and in many instances followed by
injection of a large volume of gel (in comparison to overall heart volume), proved too much and
there was only ~30% survival at two weeks compared to reported ~50% survival for similar
studies (Kumar et al. 2005; Sam et al. 2000). To address this issue it was decided to switch from
mice to rats for MI model studies as the injection post MI should cause much less trauma to the
animals in comparison to the trauma from the initial MI, and the rat LAD MI model is also
widely accepted and studied (Krzemiński et al. 2008). Initial results from a trial study using a
Lewis rat LAD MI model, done in the exact same manner as that with mice proved to be very
promising. We achieved 100% survival at three weeks, and it seemed as though injection of both
Control (No Peptide) and peptide modified (QH-G213-H) 2.5mg/mL 1:1 chitosan-collagen
hydrogel were having a marked positive effect on LV remodeling post MI. These results
motivated us to pursue further studies into the effect the developed hydrogel (alone) had on LV
remodeling, and the mechanism by which it acts.
4.1. Materials & Methods
All animal experimental procedures were approved by the Animal Care Committee of the
Toronto General Research Institute and the University of Toronto Committee on Animal Care,
according to the Guide for the Care and Use of Laboratory Animals.
4.1.1. Experimental design overview
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A brief overview of the experimental design for this chapter, along with all relevant time
points and quantification done at each step are illustrated in Figure 4-1.
Figure 4-1: In vivo experimental design A brief overview of the design used for the large scale Lewis rat LAD MI model with major time-pints and
procedures performed.
4.1.2. Peptide modified chitosan-collagen hydrogel (QHG213H) QHREDGS peptide (Biomatik) was conjugated to chitosan (UP-G213, Novamatrix) using 1-
ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) chemistry as described previously (Reis et
al. 2012). Similarly, 2.5 mg/mL 1:1 mass:mass chitosan-collagen hydrogel was prepared as
described to make Control hydrogel (no peptide) or QHG213H hydrogel (651±8 nmol
peptide/mL-gel) and kept on ice for up to 3 hours until use (Reis et al. 2012).
4.1.3. Peptide modified polyethylene glycol (PEG-QHREDGS) QHREDGS peptide was conjugated to polyethylene glycol (PEG) by incubating 8.0 mg
acrylate-PEG-NHS (MW 3500; Jenkem Technologies, Allen, TX) in a solution of 50 mM Tris
pH 8.5 containing 0 mM or 29 mM (24 mg/mL) QHREDGS solution for a final volume of 200
µL, for 3h at 700rpm. After 3 h, the reaction solutions were dialyzed extensively against ddH2O;
lyophilized and resuspended in ddH2O, to which 80 mg polyethylene glycol diacrylate (PEGDA
Mn 700; Sigma Aldrich, Oakville, ON) and 0.8 mg 2-Hydroxy-2-methyl-propiophenone was
added; 150 µL of this solution was added to a circular silicon mold (11-mm diameter, 1-mm
thick) placed atop a glass microscope slide and a polyvinyl coverslip was placed on top of the
mold. The hydrogel was placed under UV light for 10 min to polymerize, washed for 10 min in
phosphate buffered saline (PBS) then incubated at 4°C for 1 hr. After 1 hr the cover-slip was
gently peeled off the hydrogel, the mold was removed and the hydrogel was placed in 70% EtOH
under UV for 1 hr to sterilize, then placed in sterile PBS for 1 hr at 37°C.
Neonatal rat CMs or cardiac fibroblasts, isolated as previously described (Rask, Dallabrida, et
al. 2010; Radisic et al. 2003), were washed and resuspended in equal volumes of PBS or 50
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µg/mL anti-β1-integrin antibody (α-CD29; BD Biosciences, Mississauga, ON) at 37°C for 20
min. The CMs or cardiac fibfroblasts were then spun for 5 min at 200g, resuspended in CM
media, seeded onto the PEG-hydrogels (22,000 cells per hydrogel) in a volume of 10 µL and
incubated at 37°C for 50 min. After 50 min, the hydrogels were washed once with PBS, fixed
with 4% paraformaldehyde solution for 20 min, blocked for 1 hr in 5% FBS, 0.3% Triton X-100
in PBS, washed thrice in PBS, stained with DAPI for 15 min, then washed thrice with PBS.
Washed hydrogels were imaged in PBS using an Olympus IX-81 fluorescent microscope with
CellSens Dimension Imaging software (Olympus America Inc., Center Valley, PA). The surface
of each hydrogel was imaged in its entirety and the total number of cells was determined using
the maxima plug-in filter of ImageJ software (NIH, Bathesda, MD) and summing the values
determined from each individual image for a given hydrogel. The total number of adherent cells
was normalized by dividing by the total number of cell that adhered to the PEG hydrogels that
did not contain any QHREDGS peptide for a given experiment.
4.1.4. In vivo studies Lewis rats (weighing 200–250 g) were obtained from Charles River Laboratories (Saint-
Constant, QC, Canada), and myocardial infarction (MI) was generated under general anesthesia
by occluding the left anterior descending (LAD) coronary artery as previously described (Kan et
al. 2007). The Animal Care Committee of the University Health Network approved all animal
procedures. Experiments were performed according to the Guide to the Care and Use of
Experimental Animals from the Canadian Council on Animal Care. Injection of PBS (MI Only),
Control gel (no peptide), or QHG213H (peptide modified) hydrogel was done immediately
following LAD ligation and prior to closure of the thorax. Injection (3 locations, 50 µL total)
was performed using 28-guage syringes (BD Biosciences) by inserting the needle into the peri-
infarct LV wall and directing into the developing scar (MI zone). Syringes with PBS or hydrogel
were prepared prior to surgeries and kept on ice, then warmed at ~37°C for 10 min prior to
injection (Reis et al. 2012).
4.1.4.1. Assessment of cardiac function (6 Week time point)
N=8 animals per group were used for a long term 6 week study. Cardiac function was
evaluated by echocardiography (Sequoia C256 System, Siemens Medical; 15 MHz linear array
transducer) before MI (pre-ligation baseline), 3 and 6 weeks post MI. M-mode images were
obtained in the parasternal short-axis view at the level of the papillary muscles. The
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measurements were performed by a single, blinded examiner. LV internal diastolic dimension
(LVIDd) and internal systolic dimension (LVIDs) were measured. Three consecutive cardiac
cycles were recorded and averaged. Percent fractional shortening (%FS) was calculated as
follows: %FS = (LVIDd − LVIDs) / LVIDd 100 (Kan et al. 2007). The 6 week time point was
selected as it is assumed the heart has completed remodeling by this time. PBS, Control gel (no
peptide), or QHG213H (peptide modified) hydrogel was injected into 3 locations in the LV wall
(50 µL total) surrounding the infarct using a 28-gauge insulin syringe. At 3 weeks post injection
rats with no infarct or with a very large infarct were excluded from the study; only those
exhibiting a 20-40% fractional shortening as determined by echocardiography were included.
One animal from each group did not meet the criteria and were excluded, reducing numbers to
N=7/group. Cardiac function was also assessed at the end of the study (6 weeks after treatment)
with a pressure-volume catheter as previously described (Z. Sun et al. 2008). Briefly, the rats
were anesthetized and intubated with mixed oxygen and room air by a rodent ventilator (Harvard
apparatus, Canada). A pressure-volume catheter-tipped pressure transducer (2F, Millar
Instruments, USA) was inserted into LV cavity via right carotid artery. Volumes were measured
by conductance, and end-systolic and end-diastolic values were recorded. The ejection fraction,
dP/dt Max, dP/dt Min, end systolic pressure-volume relationship (ESPV) and preload recruitable
stroke work (PRSW) were calculated.
After the pressure-volume analysis was complete, hearts were perfusion fixed with 10%
formaldehyde, excised, and sectioned into five 2-mm thick slices as previously described (Kan et
al. 2007; Z. Sun et al. 2008).
4.1.4.2. 24 hr time point
N=6 animals per group were used in a short 24 hr study to elucidate the mechanism by which
the injected hydrogel preserved cardiac function. Surgery and injections were performed in the
same manner as described above and animals were monitored overnight. The animals were
anaesthetized 24 hr after injection with isoflurane, intubated, ventilated, and maintained with 2%
isoflurane. The heart was accessed in the same manner as performed for LAD surgery and
arrested using 10% KCl injected directly into the LV. Hearts were then immediately excised and
the right ventricle and atria were removed. The LV was then sectioned into a MI region, border
region, and remote region. These sections were placed in separate cryo-tubes and snap frozen in
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liquid N2. Frozen tissues were subsequently ground using standard mortar and pestle with liquid
N2 and kept at -80°C prior to RNA and protein extraction.
4.1.4.3. In vivo hydrogel lifespan
To assess the in vivo lifespan of our hydrogel, chitosan was conjugated with the near infrared
dye Dy-Light 800 (Thermo Scientific, Waltham, MA) and injected into the heart of live animals.
Briefly, 1 vial of Dy-Light 800 label (suitable for 1 mg of protein) was dissolved in 57.7 μL PBS
and mixed immediately with 92.3 μL of chitosan at 20 mg/mL in 0.9% normal saline (final
reaction volume of 150 μL and chitosan concentration of 12.31 mg/mL). The mixture was
protected from light, placed on a microtube shaker at 650 rpm for 1 hr, and then stored at 4°C
until use. Upon use, the Dy-Light labelled chitosan was used to make hydrogels as per the
standard protocol (Reis et al. 2012), kept on ice, and then 50 μL of the labelled gel was injected
directly into 3 locations of the healthy LV wall of N=7 Lewis rats. Rats (anesthetized under
isoflurane) were then imaged using a Kodak In-Vivo FX Pro Imager (Kodak Molecular Imaging
Systems), with a 5 sec exposure X-ray image overlaid with a 40 sec exposure fluorescent image
(760 nm excitation, 830 nm emission). Full-sized images were processed to illustrate fluorescent
decrease over time by applying the same threshold and gamma adjustment to all images using
Kodak MI 4.0 software (Kodak Molecular Imaging Systems). Inset images were processed to
highlight fluorescent localization by applying the optimal brightness/contrast and threshold to
each individual image using ImageJ (NIH). N=3 rats were sacrificed 1 hr after injection, and
N=4 were monitored up to 14 days post injection. Excised hearts were stored at -80°C until use,
at which time they were homogenized and fluorescently imaged to quantify degradation.
4.1.5. Immunohistochemistry Samples for immunohistochemical and immunofluorescent staining were prepared as
described above. For paraffin embedding and sectioning, as well as hematoxylin and eosin
(H&E), Mason’s trichrome, smooth muscle actin (SMA), Factor VIII (FVIII), and CD31
staining, the fixed samples were sent to the Pathology Research Program (PRP) at the University
Health Network. Chitosan staining was performed using Cibacron Brilliant Red-3BA (CBR-
3BA, Sigma Aldrich) and Weigert’s Iron Hematoxylin as described (Rossomacha, Hoemanni,
and Shive 2004; Reis et al. 2012).
Sections stained for cardiac troponin-T (cTnT) and TUNEL were deparaffinised and
rehydrated with subsequent baths (3 each) of 100% xylene, 100% EtOH, 95% EtOH, 75% EtOH,
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and distilled water for 3 min per change. Antigen retrieval was performed by microwaving slides
immersed in Tris-EDTA buffer for 5 min, permeabilized with 0.1% Triton-X, 0.1% Sodium
Citrate in Tris-buffered saline (TBS) solution (10 min), and blocked using Dako Serum Free
Protein Block (Dako Canada, Inc., Burlington, ON X0909) for 30 min at room temperature, with
3x3 min TBS washes between steps. Primary antibody (mouse monoclonal cTnT antibody,
Thermo Scientific MS-295-P, diluted 1:200) was applied to sections for 2 hr at room temperature
followed by secondary antibody (Alexafluor 594 goat anti-mouse IgG (H + L), Invitrogen) at
1:400 dilution for 1 hr at room temperature. TUNEL label and enzyme (Roche Applied Science,
Indianapolis, IN, 11767291910) were applied as per manufacturer’s instructions. DAPI (Sigma
Aldrich) nuclear stain was applied at a 1:1000 dilution for 10 min, and finally slides were
mounted using Dako Fluorescence Mounting Medium (Dako Canada, Inc., S3023), and imaged.
4.1.6. Quantitative PCR Whole heart or heart sections frozen at -80°C were ground using a mortar and pestle and
liquid N2. RNA was extracted with TRIZOL Reagent following the manufacturer’s protocol.
Isolated RNA was resuspended in 200 μL RNAse/DNAse free water and quantified using a
NanoDrop 1000. cDNA synthesis and RT-qPCR were performed as previously described (Nunes
et al. 2013). All oligonucleotide sequences used are listed in Table C-1 (Appendix C). Data was
standardized between samples to reflect gene expression changes of the progressing scar,
quantifying gene expression in the border zone compared to the established MI zone. Fold
changes in gene expression between groups were analyzed and genes showing a >2-fold or <0.5-
fold change in the QHG213H gel group over the MI Only (PBS) injection group were considered
significant.
4.1.7. Western blotting Whole heart or heart sections frozen at -80°C were ground using a mortar and pestle and
liquid N2. Protein was isolated using Pro-Prep Protein Extraction Solution (iNtRON
Biotechnology, Inc., Seongnam, KR) according to the manufacturer’s instructions. Samples were
run on Novex® Tris-Glycine 1.5-mm, 10-well, 10% Gels (Life Technologies) for 1.5 hr at 125 V
then transferred to nitrocellulose membrane by wet transfer at 125 V for 2.5 hr. Membranes were
probed with: α-phospho-MAPK (1/1000 in 5% BSA w/v in PBS with 0.1% Tween20), MAPK
(1/1000), α-ILK (1/1000), or GAPDH (Millipore, MAB374, 1:10,000) antibodies. Secondary
antibodies used were peroxidase-conjugated (DAKO, P0448 or P0447, 1:2000). Membranes
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were developed with ECL reagent Luminata Classico Substrate (Millipore) and exposed to film.
The films were scanned and densitometry was performed using ImageJ software.
4.1.8. Image analysis techniques Morphometric analysis was performed as described previously (Kan et al. 2007; Z. Sun et al.
2008). TUNEL/DAPI cell counts were made using a macro developed in Olympus CellSens
Dimension software and tested on control images. TUNEL/cTnT MI border zone quantification
was based on 2 independent images on either side of the MI zone in each animal, and MI zone
quantification performed by averaging at least 4 independent images within the scar. All counts
and normalizations were performed for each image and averaged to get a single value per
sample. Quantifying differences in vasculature in the MI zone and border zone of hearts excised
3 and 6 weeks post MI was done using SMA and FVIII/CD31 stained sections as described
previously (Chiu and Radisic 2010).
Automation of the image quantification process eliminates human errors and bias between
images, ensuring consistent quantification among all images and more accurate results.
4.1.9. Statistical analysis Statistical analysis was performed using SPSS Statistics 17.0 and GraphPad Prism 5.0.
Differences between experimental groups were analyzed using one or two-way ANOVA with
Bonferroni post-hoc tests unless otherwise specified. P<0.05 was considered significant for all
statistical tests. Results were plotted with GraphPad Prism 5.0, with all data being reported as
mean ± standard deviation (Mean±SD), unless otherwise noted.
4.2. Results
4.2.1. In vivo degradation (lifespan) Animals imaged 1 hr after injection of the Dy-light 800 labelled hydrogel showed localization
of labelled hydrogel within the heart; with unconjugated dye and gel washout in what appeared
to be the chest of the animals (Figure 4-2). Images taken 24 hr after injection similarly showed
localization of labelled hydrogel within the heart while the fluorescence in the chest that was
visible in the 1 hr images was no longer evident, rather waste gel/dye appeared to be present in
the stool of the animals. Subsequent images at 3, 5, 7, and 14 days post injection show continued
localization of labelled hydrogel in the heart; however the extent of fluorescence decreased with
time. By 14 days post injection only a minimal amount of fluorescent label was visible (Figure
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4-2). Staining for the presence of chitosan in heart sections derived from hearts excised 3 weeks
post injection confirmed the presence of hydrogel. However, positive chitosan staining was
evident in only ~20% of animal hearts indicating the in vivo lifespan of the hydrogel in the heart
was approximately 3 weeks (Figure 4-2).
The remaining (N=4) animals were sacrificed after imaging on day 14, the excised hearts
were homogenized, and the fluorescence intensity was measured and compared to N=3 hearts
excised 1 hr post injection. The results indicated that approximately 60% of the gel present at 1
hr is cleared by 14 days (~40% remained on day 14, Figure C-1, Appendix C).
Figure 4-2: In vivo lifespan of Dy-light 800 labelled chitosan-collagen gel Dy-light 800 labelled gel was detectable in the heart of live animals using a Kodak In-Vivo FX Pro imager
from 1 hour (hr) up to 14 days (d) post injection (visualized on day 14 with increased exposure, ↑exp.). Insets. Enlarged and enhanced view of the region indicated by the white square. Chitosan staining of heart sections extracted 3 weeks (Wk) post injection showed continued presence of chitosan in small islands in the infarct area (purple dots); however this was only observed in a minority (~20%) of samples indicating almost complete clearance of the hydrogel from the heart.
4.2.2. Functional Data All animals exhibited nearly identical echocardiographic parameters prior to surgery
irrespective of their grouping. The LAD coronary artery ligation procedure resulted in significant
ventricular dysfunction and LV dilation. These were seen as a decrease in fractional shortening
and an increase in both LV internal dimension at systole and diastole (LVIDs and LVIDd,
respectively) at 3 weeks, which was exacerbated by 6 weeks (Figure 4-3 A-C). Interestingly
there was a significant amelioration of cardiac remodeling, measured by cardiac fractional
shortening and LVIDs, in the Control gel (no peptide) group as compared to the MI Only (PBS)
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group at both 3 and 6 weeks, however the QHG213H gel group outperformed both the MI Only
(PBS) and Control gel groups (Figure 4-3 A). Most of the decrease in fractional shortening for
the MI Only and Control gel groups occurred in the first three weeks after MI, as there was no
significant reduction in fractional shortening between 3 and 6 weeks for either group, while the
decrease was lesser and more gradual with QHG213H gel injection (Figure 4-3 A&D).
Measurement of LVIDs showed that gel treatments with and without the QHREDGS peptide
significantly reduced LV remodeling during systole (LVIDs) over the monitoring period (Figure
4-3 B), but did not have a marked effect on the diastolic dimension (LVIDd, Figure 4-3 C).
There was a statistically significant level of interaction between factors Group and Time for %FS
and LVIDs (P<0.001) but not for LVIDd (Figure 4-3 A-C).
At 6 weeks post MI, hearts were further evaluated using load dependent and load independent
measurements to assess cardiac function and LV volumes. Steady-state (load dependent) ejection
fraction, dP/dt Max, and dP/dt Min were all significantly better in the QHG213H gel group as
compared to either the Control gel or MI Only (PBS) groups (Figure 4-4 A-C). Notably,
injection of the Control gel (no peptide) itself showed a significant improvement in cardiac
function as compared with the MI Only (PBS) group (Figure 4-4 A-C). Furthermore, the smallest
end systolic LV volumes were seen in the QHG213H group, followed by the Control gel group,
which were both significantly lower than the MI Only (PBS) group (Figure 4-4 D). However, no
difference was observed in the end-diastolic LV volumes between groups (Figure 4-4 E).
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Figure 4-3: Cardiac function post MI measured by echocardiography QHG213H gel injection significantly maintained cardiac function (measured by %fractional shortening,
%FS) at 21 days (3 weeks) and 42 days (6 weeks) post LAD surgery as compared to the Control gel (no peptide) and MI Only (PBS) injection groups; and the Control gel injection induced a significant maintenance in %FS relative to the PBS injection (MI Only) (A). Injection of gel with or without peptide significantly reduced LV dilation during systole at 3 and 6 weeks relative to the PBS injection (MI Only); and the QHG213H gel induced a significant improvement in LVIDs relative to the Control gel (B). No difference in diastolic LV dilation (LVIDd) was seen over the monitoring period (C). %FS and LVIDs had statistically significant interaction between factors Group and Time (P<0.0001), but no interaction effect was present for LVIDd. (D) Post-hoc pair-wise comparisons for factor “Days after Injection” for data from (A-C). Data analyzed by two-way repeated measures ANOVA and Holm-Sidak post-hoc analysis and expressed as Mean±SD, N=7/group.*P<0.01, **P<0.001 vs MI Only; #P<0.01 vs Control gel.
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Figure 4-4: Load dependent cardiac function and LV volumes at 6 weeks post MI Steady-state pressure-volume analysis was used for load dependent cardiac analysis after heart excision.
Injection of the Control gel (no peptide) significantly improved cardiac function compared to MI Only (PBS), and injection of the QHG213H gel further improved function (A-C). End systolic LV volume was significantly lower in the QHG213H gel group vs the Control gel and MI Only (PBS) groups, and in the Control gel group vs MI Only (PBS) group (D). The end LV diastolic volume was unchanged between groups (E). Data expressed as Mean±SD, N=7/group, *P<0.05, **P<0.01 vs MI Only (PBS), #P<0.05 vs Control gel.
The load independent end systolic pressure-volume relationship (ESPVR, Figure 4-5 A) and
preload recruitable stroke work (PRSW, Figure 4-5 B) were significantly improved with
injection of gels with or without peptide as compared to the MI Only (PBS) group. However,
addition of the peptide to the gel enhanced its protective effects by inducing further
improvements to these parameters Figure 4-5 A&B).
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Figure 4-5: Load independent pressure-volume analysis at 6 weeks post MI Occlusion pressure-volume analysis was used for load independent cardiac functional analysis after excising
hearts. Significant differences were seen in both end systolic pressure (A) and stroke work (B) between the QHG213H gel group and either the Control gel MI Only (PBS) groups, as well as between the Control gel group and the MI Only (PBS) group. Data expressed as Mean±SD, N=7/group, *P<0.05.
4.2.3. Gross morphology and histology PBS injected (MI Only) hearts appeared larger than Control (no peptide) gel and QHG213H
gel injected hearts; and the extent of remodeling appeared more extensive in the PBS injected
heart cross sections (Figure 4-6 A). Quantification of the scar fraction and MI zone scar
thickness from fixed sections showed that the MI Only (PBS) hearts underwent the most
extensive remodeling; that the Control gel injected group had significantly smaller fractional scar
coverage and significantly greater scar thickness than the MI Only (PBS) group; and that the
QHG213H gel group had significantly smaller fractional scar coverage and significantly greater
scar thickness than either the Control gel or MI Only groups (Figure 4-6 B&C).
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Figure 4-6: Gross heart morphology at 6 weeks post MI Representative whole heart and heart sections (from just below suture to apex of heart, 5 sections per heart)
illustrate differences in gross morphology between groups (A). Fractional MI scar coverage (B) and MI scar thickness (C) showed significant differences between groups, with QHG213H gel having significantly smaller scar coverage and larger scar wall thickness than both Control gel and MI Only (PBS) groups, as determined from gross section images (A). Data expressed as Mean±SD, N=7/group, *P<0.05, **P<0.01 vs MI Only (PBS), #P<0.05 vs Control gel.
Heart sections were histochemically stained and imaged for the different treatment groups.
Differences in scar thickness and area between treatment groups was apparent and there also
appeared to be more healthy tissue (red) within the MI zone in the QHG213H group in
comparison to the other groups based on Masson’s trichrome staining (Figure C-2 A, Appendix
C). Sections were additionally stained for SMA (a marker of smooth muscles cells present in
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mature vasculature), as well as both CD31 and FVIII (markers of endothelial cells and found in
both neo- and mature vasculature) (Figure C-2 B-D, Appendix C). The mean vessel density and
diameter of total vascularization (FVIII+, Figure 4-7 top panel) and mature vasculature (SMA+,
Figure 4-7 bottom panel) were quantified for both the MI and border zones. No difference in
vasculature metrics was seen in the MI zone, or in the FVIII+ vessels in the border zone.
However, there were significantly larger mature SMA+ vessels in the border zone for the
QHG213H gel treatment group as compared with either the MI Only (PBS) and Control gel
groups.
Figure 4-7: Vascularization at 6 weeks post MI Vascularization within the MI and border zones was quantified at 6 weeks from heart cross sections. Vessel
density and diameters were quantified from SMA stained sections to assess mature vasculature, and from FVIII stained sections for total (neo- and mature) vascularization. The QHG213H gel group had significantly larger mature SMA+ vessels in the border zone. Data expressed as Mean±SD; N=7/group.
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4.2.4. Cardiomyocyte survival mechanism TUNEL and cTnT staining was used to determine the number of apoptotic CMs in the MI and
border zones (at 6 weeks post infarct), in order to determine if the functional and morphological
improvements seen with the Control and/or QHG213H gel injections could be attributable to a
reduction in CM apoptosis. Fewer TUNEL+ (apoptotic) cells were visible within the border
region of the Control and QHG213H gel groups as compared to the MI Only group (Figure 4-8
A). Quantification of the TUNEL staining revealed that gel treatment with or without peptide
significantly reduced the fraction of apoptotic cells in the border zone (P=0.005 MI Only vs.
Control gel, P=0.012 MI Only vs QHG213H gel; Figure 4-8 B). In the MI zone, there was no
difference in the apoptotic cell fraction among the treatment groups, however the QHG213H gel
group had a higher proportion of cTnT+ cells than either the MI Only (PBS, P=0.008) and
Control gel (P=0.005) groups (Figure 4-8 C) at six weeks after treatment.
To gain further insight into the CM survival mechanism, RNA isolated from tissues from the
MI and border zones of hearts excised 24 hr after treatment were subjected to RT-qPCR and a
total of 27 genes were assessed, including genes conventionally ascribed as apoptotic,
necroptotic, pro-survival, anti- or pro-inflammatory, and cardiac-specific. Notably, between the
QHG213H gel and MI Only (PBS) groups, there was a 4.50- and 2.29-fold increase in the pro-
inflammatory interleukin (IL)-6 and IL-1β gene expression, respectively; a 7.69-fold decrease in
the pro-apoptotic caspase-9 (CASP9) gene expression; a 2.32-fold decrease in the pro-survival
PI3K gene expression; and a 2.17-fold increase in the anti-apoptotic BCL2 gene expression
(Figure C-3, Appendix C). In the case of all the aforementioned genes, the addition of
QHREDGS peptide to the gel resulted in a larger fold change in the expression levels than the
Control gel (Figure C-3, Appendix C). It must be noted that none of these fold changes were
statistically significant and we can only comment on the trend observed.
Additionally, Western blot analysis was performed to determine if integrin-linked kinase
(ILK) and MAPK were implicated in the CM survival mechanism as both proteins have been
shown to be up-regulated by QHREDGS to promote stem cell survival in recent work (Dang et
al. 2014). In analyzing border zone tissues we did not see a difference in ILK protein levels or
MAPK phosphorylation between the groups, however we did see a significant increase in MAPK
protein expression in the QHG213H peptide gel group over the Control gel and MI Only (PBS)
groups (Figure C-4, Appendix C).
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To determine if the interaction of CMs and cardiac fibroblasts with the QHREDGS peptide
was dependent on the β1-integrin subunit (identified as the receptor in endothelial cells and stem
cells) polyethylene glycol (PEG) hydrogels with and without the QHREDGS peptide were
generated (Miklas et al. 2013; Dang et al. 2014). PEG was chosen as the base hydrogel because it
is non-fouling in short term culture scenarios and does not promote cell attachment; tissue
culture plastic (TCP) was used as a positive control for cell attachment. The peptide
concentration in the PEG hydrogel was 11.7±4.5mg/mL (9.7±3.7nM) (Feric et al. 2014). The
addition of the QHREDGS peptide to the PEG hydrogel significantly increased the number of
adherent cells relative to PEG alone (Figure 4-9A&B). Pre-incubation of either cell type with an
anti-β1-integrin (CD29) antibody reduced CM adhesion to the QHREDGS-PEG hydrogel to a
level comparable to that of the PEG hydrogel without peptide (Figure 4-9A&B).
Figure 4-8: Apoptosis at 6 weeks post MI Sections from hearts extracted at 6 weeks were stained for TUNEL (red), cTnT (green), and DAPI nuclear
stain (blue). Representative images showed fewer TUNEL+ cells in the Control gel and QHG213H gel groups than in the MI Only (PBS) group in the border zone (Scale bars=100µm) (A). Quantification of the fraction TUNEL+ cells (# of TUNEL+ cells/# of DAPI+ cells determined per image) in the border zone showed a significantly smaller TUNEL+ cell fraction for both the Control and QHG213H gel groups relative to the MI
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Only (PBS) group, but no statistical difference between the gel treatment groups (B). In the MI Zone, there was no difference in the TUNEL+ cell fraction between groups, however the QHG213H group showed a significant increase in cTnT expression relative to either the MI Only (PBS) and Control gel groups (C). Data expressed as Mean±SD,* P<0.05, ** P<0.01, N=7/group.
Figure 4-9: Cardiac cell attachment to PEG and PEG-QHREDGS gels Rat NCMs (A) or cardiac fibroblasts (B) were seeded onto PEG only (no peptide) or PEG-QHREDGS disks with or without
the pre-incubation with an anti-β1-integrin (CD29) antibody. After 50 min, the attached cells were counted. The data is normalized to the PEG only gels. Two way ANOVA showed both NCMs and cardiac fibroblasts have significantly more attached cells to PEG-QHREDGS surfaces over PEG alone, but the difference is abolished when pre-incubated with anti-β1-integrin antibody. Data expressed as Mean±SE, * P<0.05, ** P<0.01, N=6/group.
4.3. Discussion
We have demonstrated, by means of in vivo tracking that our chitosan-collagen hydrogel can
localize within the beating heart and can remain in situ for at least two weeks after injection. We
have also shown by chitosan staining of heart sections that 3 weeks after injection hydrogel was
present in ~20% of samples. Moreover, we found a significant improvement in cardiac function
at 6 weeks post MI in an LAD rat model with injection of the chitosan-collagen hydrogel, and
further functional improvements with injection of the QHREDGS conjugated chitosan-collagen
hydrogel. Finally in terms of mechanism, we found that adhesion of rat NCMs to the QHREDGS
peptide is mediated by β1-type integrin receptors; and that in the border zone protein levels of the
pro-survival molecule MAPK was significantly increased by the presence of QHREDGS in the
hydrogel at 24 hr. We observed trends toward up-regulated gene expression of the pro-
inflammatory interleukins IL-6 and IL-1β and the anti-apoptotic protein BCL2, whereas gene
expression of the pro-apoptotic enzyme caspase-9 appeared to be down regulated.
Hydrogel degradation is most often characterized in vitro or by histological staining of
excised in vivo samples (Bouten et al. 2011; Das et al. 2012). While techniques for on-line in
vivo quantification of hydrogel distribution and degradation exist for subcutaneous and intra-
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muscular injections (biotin labeling, in vivo fluorescent imaging systems, MRI, PET, etc), using
these systems for tracking hydrogels in the heart prove extremely difficult for a number of
reasons. These include the dynamic nature of the heart, which continues to beat even when the
animal is sedated, the depth at which the heart is located within the animal, and the often small
volume of material injected therein (~20-200 µL) (Bouten et al. 2011; Ruggiero et al. 2012). The
near infrared dye Dy-Light 800 was used in this study specifically due to its compatibility with
the Kodak In-Vivo Imaging System. This is because this system has the potential to provide on-
line information regarding the localization and lifespan of the hydrogel and the potential to
minimize the number of animals required for the study by eliminating the need for heart tissue
extractions at various time points from which to assess the distribution and degradation of the
hydrogel. Furthermore, the use of a near infrared dye permits deep tissue imaging without
affecting the surrounding tissue as in the case of long term radiolabels used for PET imaging, or
the inaccurate degradation profiles that can result from macrophage uptake of the contrast agents
(e.g. iron oxides or gadolinium) used for MRIs (Ruggiero et al. 2012). However, near infrared
fluorescence imaging has its own limitation in that it provides the lowest spatial resolution.
By imaging the injected labelled hydrogel over a 14 day period, we were able to confirm
localization and retention of the hydrogel within the heart (Figure 4-2), however full
quantification of the degradation rate proved more challenging. Quantification from the images
was not possible because the 2D imaging system could not provide the necessary 3D information
about the volume of injected material and because of distortion in the images due to the constant
beating of the heart over the 40 sec imaging period. Furthermore, we did not fully quantify the
conjugation efficiency, however as the mechanism of DyLight800 conjugation to chitosan is the
same as that for peptide conjugation used we assumed efficiency was about equal at 50%.
Quantification was therefore attempted using excised tissue samples and required two more
assumptions: (i) the bulk of unbound dye was removed from the gel within the first hour
following injection and (ii) the gel present at 1 hr after injection represented the total amount of
gel that was successfully localized. With these assumptions, a comparison of the amount of gel
present at 14 days and the amount present at 1 hr post injection resulted in the information that
approximately 60% of the gel was degraded or removed from the heart over a two week period
(Figure C-1, Appendix C).
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One question that has been raised concerning the in vivo presence of the gel is what happens
to it after being removed from the heart, and specifically why there is presence of fluorescence
(thus chitosan) in the animal’s urine and lack thereof in the kidneys or liver. It is recognized that
the in vivo biodistribution of chitosan is one of the least studied aspects of the compound, with
molecular weight, deacetylation, and particle sizes all having an effect (Kean and Thanou 2010).
Two studies in particular describe the clearance of water soluble sugars of similar size to the
chitosan we used (Yamaoka, Tabata, and Ikada 1993; Onishi and Machida 1999). The first study,
looking into the clearance of dextrans and pollulans of varying MW from the blood finds that
even the largest sugars (100s of kDa) are processed and accumulate in animal excrements within
hours (Yamaoka, Tabata, and Ikada 1993). The second explored the biodistribution of FITC
labelled chitosan (MW ~100kDA) injected directly into the peritoneum in a mouse model. They
found that injected chitosan was completely absorbed from the peritoneal cavity in 14hr, and that
there was rapid renal clearance with 25% being excreted 1hr after injection, and 100% 14 hr post
injection (Onishi and Machida 1999). In our case, initial leaking of injected gel into the chest
cavity would be rapidly cleared as described by Onishi and Machida (1999). For hydrogel that
leaks directly into the left ventricle, it would again be carried directly to the waste processing
organs and again cleared quickly (Yamaoka, Tabata, and Ikada 1993). Similarly, hydrogel that
has successfully gelled and localized within the LV is degraded over time. As there is no
exogenous crosslinking providing structure, or creating large masses of chitosan-collagen
composites, degradation of the collagen holding the hydrogel together releases uncrosslinked
chitosan with a maximum MW of 600kDa which would not be expected to accumulate in vivo
and be quickly cleared as waste through the urine.
We further determined that the amount of QHG213H (peptide modified) hydrogel that
localized in the heart and its lifespan in the heart was sufficient to beneficially affect the cardiac
remodeling process post MI. This is evidenced by the improved cardiac functional outcomes
seen with gel injections with the QHREDGS peptide as compared to both the gel alone and the
PBS injected (MI Only) rats 6 weeks post MI (Figure 4-3, Figure 4-4, & Figure 4-5).
Specifically, in terms of cardiac output parameters, the QHG213H (peptide modified) gel
treatment resulted in a 16% improvement in ejection fraction (from 26% to 42%) and an 8%
improvement in fractional shortening (from 23% to 31%) from the MI Only (PBS) treatment
group (Figure 4-3 & Figure 4-4). Moreover, these improvements in cardiac output with
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QHG213H gel injection translate into an improvement/maintenance in ejection fraction and
fractional shortening from severely impaired (MI Only) to mildly impaired (QHG213H gel), on a
normal–mild–moderate–severe scale of impairment (R. M. Lang et al. 2006). At 6 weeks post-
MI, we observed a significant decrease in the systolic dimension and volumes but did not find
any difference in the diastolic parameters. This may be because while systolic function decreases
significantly due to CM necrosis post-MI, a longer time frame (e.g. 12 weeks) may be required
to observe changes in diastolic parameters due to matrix remodeling and scar expansion.
In terms of the morphological changes induced by QHG213H gel injection 6 weeks post MI,
we observed a decrease in infarct size and retention of infarct wall thickness with gel injections
relative to the PBS injections, with addition of the QHREDGS peptide to the gel providing the
least amount of adverse remodeling. These morphological improvements in the infarct area could
have been responsible for the increased ejection fraction and fractional shortening with gel
injections relative to the PBS injection. The gel may have stabilized the infarct wall and
increased its thickness by acting as a temporary scaffold that altered the surrounding tissue
properties and reduced wall stress. This is unlikely, however, because the storage modulus of the
Control (no peptide) and QHG213H (peptide modified) gels were quantified to be ~55Pa and
~45Pa respectively, well below the reported stiffness of native rat myocardium (1kPa at the end
of diastole to 140kPa at the end of systole) (Venugopal et al. 2012; Reis et al. 2012). Also, we do
not see a difference in end LV diastolic volume between groups, which would be expected if the
hydrogel were aiding LV mechanics and altering the geometry (Figure 4-4). It is therefore more
probable that the hydrogel affected cardiac function post MI by acting at the cellular level,
possibly by promoting cell survival directly via reducing apoptosis or indirectly via increasing
cell-cell contacts.
Indeed, quantifying the CM presence in the MI zone 6 weeks post injection showed a
significantly higher proportion of CMs in the QHG213H injected group as compared to the
Control (no peptide) gel and MI Only (PBS) groups. While there was no difference in apoptotic
cell numbers in this same region at 6 weeks (Figure 4-8), this is not unexpected as apoptosis
would be expected to level off by this time point since most cell death occurs within the first 72
hr post-MI (Sutton and Sharpe 2000). Interestingly, apoptosis was reduced in the border region
of hearts treated by either Control (no peptide) or QHG213H gel as compared to PBS injection at
6 weeks (Figure 4-8), which might suggest that the gel provided an environment that is more
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conducive to cell survival in the long term. Furthermore, the QHG213H gel injection resulted in
larger, mature vessels in the border region at 6 weeks relative to the Control (no peptide) gel and
MI Only (PBS) injection groups (Figure 4-7). This may be because Ang1 angiogenic effects are
mediated through the receptor Tie2, while the peptide QHREDGS is the Ang1 integrin-binding
site(Dallabrida et al. 2005), and may not elicit identical angiogenic responses. Other Ang1-
derived peptides (e.g. Vasculotide) have been demonstrated to bind Tie2 and are angiogenic
(Van Slyke et al. 2009). Moreover, the Ang1-CM interaction is integrin-dependent and Tie2-
independent (Dallabrida et al. 2005); while QHREDGS has been demonstrated to interact with
integrins on endothelial cells (α5β1 and αvβ3) and induced pluripotent stem cells (β1-type) to
promote cell adhesion and survival (Miklas et al. 2013; Dang et al. 2014).It is therefore possible
that the higher CM proportion in the QHG213H injected group could be due to maintained
vascularization, which enabled more cells to acquire the necessary nutrients and oxygen for
survival.
In order to further elucidate the mechanism by which the QHG213H gel injection resulted in
improved cardiac functional and morphometric measures at 6 weeks post MI relative to the PBS
injection, we investigated the effect of treatment on the post MI acute phase (24 hr) wherein
critical apoptotic and immune/inflammatory responses predominate (Figure C-3, Appendix C). A
comparison of the gene expression in the tissue derived from the border zone to that of the MI
zone revealed trend toward increased (2.17-fold) anti-apoptotic BCL2 expression and decreased
(0.13-fold) pro-apoptotic CASP9 expression in the QHG213H gel group relative to the MI Only
(PBS) group. We also observed trend toward increased gene expression of the pro-inflammatory
cytokines IL-6 and IL-1β (4.5- and 2.29-fold, respectively) and decreased (0.43-fold) expression
of PI3K, a pro-survival signaling molecule. However it must be noted that none of these fold
differences were significant, and that the relative contributions of the various signaling pathways
to CM hypertrophy and survival is complex (Nian et al. 2004; Hwang et al. 2001; Fahmi et al.
2013). Hence, while IL-6 is most commonly associated with detrimental post-MI inflammation it
has also been found to promote CM survival in the acute phase of cardiac injury (Fahmi et al.
2013), likely due to the fact that there is a benefit to inflammation in the short term with respect
to repair. Similarly, IL-1β is associated with fibrosis post MI and has classically been viewed as
promoting detrimental post infarct remodeling; however studies in mice have shown that IL-1β
activation is critical in the acute phase of MI and inhibition leads to worse outcomes (Hwang et
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al. 2001). These findings highlight the importance of timing in the activation of various
genes/proteins in post infarct inflammation and long term remodeling (Nian et al. 2004; Fahmi et
al. 2013). Notably, the Control gel (no peptide) had an effect on the expression levels of many of
these same genes (Figure C-3, Appendix C), which might explain the intermediate effects we
observed with injection of the gel without peptide on cardiac function and morphology relative to
that of the QHG213H gel and MI Only (PBS) treatment groups. In general, the addition of the
QHREDGS peptide to the gel seems to exacerbate the effects of the gel alone, however in some
instances the effect of the QHG213H gel appears to be distinctive. For example, expression of
the anti-apoptotic gene BCL2 decreased (0.67-fold) in the Control (no peptide) gel treatment
group relative to the MI Only (PBS) group, whereas the QHG213H gel group had increased (>2-
fold) BCL2 expression relative to the MI Only (PBS) group. These results therefore suggest that
the effect on BCL2 expression is dependent upon the QHREDGS peptide.
We have recently shown that the integrin subunits α5, β3, and β1 and the downstream effectors
ILK and MAPK are involved in the QHREDGS-specific pro-survival responses in stem cells and
endothelial cells (Miklas et al. 2013; Dang et al. 2014). While we were unable to detect up-
regulation of these genes at 24 hr (Figure C-3, Appendix C) nor up-regulation of the ILK protein
or of MAPK phosphorylation between the treatment groups, we did observe a significant
increase in MAPK protein expression in the QHG213H gel group compared to either the Control
(no peptide) gel or the MI Only (PBS) groups (Figure C-4, Appendix C). Here again the timing
of the assessment is important as phosphorylation and translation events can occur within
minutes to hours, whereas transcriptional events can occur within hours to days. Thus choosing
an appropriate time point to capture all of these events, while at the same time minimizing the
number of necessary animals, is a daunting task. Herein we did not observe up-
regulation/activation of the genes/proteins of interest, with the exception of MAPK protein;
however we were limited by our assessment at a single time point (24 hr). As such, the fact that
we did not observe differences between the groups is likely more indicative that we did not
capture the up-regulation/activation events in the tissues effectively, rather than that these events
did not occur. An extensive and individualized time course for the events of interest would be
required to fully delineate the mechanism by which the QHG213H gel promotes CM survival in
the post infarct milieu.
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Using QHREDGS peptide synthesized to a PEG hydrogel we showed that both CMs and
cardiac fibroblasts bind to QHREDGS through β1-type integrin receptors (Figure 4-9A&B). We
have previously demonstrated that QHREDGS interacts with endothelial cells through α5β1 and
αvβ3 integrins to promote cell survival (Miklas et al. 2013). Moreover, this data might provide
some insight into the in vivo mechanism. Two of the four isoforms of the β1-integrin subunit are
present on the CM surface and are involved in cell-matrix interactions. The β1D isoform is
present in adult myocytes and is important for firm anchoring of the myocyte to the matrix,
which provides a mechanical advantage during contraction. The β1A isoform is present in fetal
myocytes and promotes cell mobility and proliferation at the expense of efficient contractility
(M. Sun 2003). On exposure to cytokines such as TNFα (present in high amounts post MI) the
β1D isoform is replaced by the fetal β1A isoform, which can lead to reduced cardiac function and
detrimental remodeling (M. Sun 2003). β1 is also implicated in cardiac fibroblast anchoring to
the ECM (specifically collagen gels), which through further stabilization of the ECM imparts
beneficial effects to CMs (M. Sun 2003). Taken together, it is possible that the binding of cardiac
cells—both myocytes and non-myocytes—to the QHREDGS peptide through β1-type integrins
after MI leads to stabilization of cell-ECM contacts (protects transition of β1D to β1A in CMs
specifically), inhibiting apoptosis, and modulating genes associated with inflammation in the
acute phase (IL-6 and IL-1β), which results in increased CM survival. This then modifies the
chronic inflammation which would normally occur, resulting in a reduction in pathological
remodeling such that impairment to cardiac function is minimized (Figure 4-10).
Figure 4-10: Proposed QHREDGS mediated mechanism It is suspected that binding of cardiac cells to QHREDGS through β1-integrins leads to up-regulation of anti-
apoptotic BCL2 and reduction in pro-apoptotic CASP9, as well as up-regulation of the inflammatory cytokines
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IL-6 and IL-1β. This leads to constructive inflammation in the acute phase and results in early stabilization of the scar and reduced CM loss within the infarct zone, resulting in long term improved cardiac function and reduced detrimental pathological remodeling compared to hydrogel treatment alone or no treatment.
An analysis of animal MI studies performed with treatments ranging from cell sheet
implantation to co-injection of cells with hydrogel post MI, and animal monitoring anywhere
from 4 to 8 weeks post treatment, was performed to see how well the data from our study
compared to that of other groups. Major reported outcomes from these studies are presented in
Table 4-1. The data are presented as a relative change in the outcome of the treatment group to
the control group (in all cases PBS, saline, or media injection) at the end point of the study. For
example, the ejection fraction at 6 weeks post MI in our study was 41.6% for the QHG213H
treatment group, and 26.2% for the MI Only (PBS) control group, giving a 62% increase in the
QHG213H group over the MI Only group at the end of the experiment. In comparison to these
studies, our results are very encouraging in that all our metrics are equivalent to or better than
those reported by the other groups. To properly assess the impact of this work, one must consider
the general simplicity of our treatment compared to others assessed. The mouse embryonic stem
cells delivered in a chitosan-β-glycerol phosphate hydrogel (Lu et al. 2009), the human
embryonic stem cell derived cardiac cells injected in a thymosin β-4 conjugated to PEG hydrogel
(Kraehenbuehl et al. 2011), or the bone marrow stem cells delivered with α-CD/MPEG-PCL-
MPEG synthetic hydrogel (T. Wang, Jiang, et al. 2009) resulted in 22%, 25%, and 77% increases
in ejection fraction respectively. We accomplished a 62% increase in ejection fraction with a
naturally thermogelling hydrogel made from natural components to which a seven amino acid
peptide was immobilized, and without the use of cells. Furthermore, while other hydrogel
formulations (both natural and synthetic) may allow for a higher degree of control of material
properties, the use of exogenous chemical cross-linkers (or harsh cross-linking mechanisms)
raise questions about their safety, whereas the mild, physiologic conditions in which our
formulation gels mimics the natural process of ECM formation in vivo and makes it a safe,
attractive choice for use in the heart as well as to potentially deliver cells in the future
(Nicodemus and Bryant 2008). Finally, the gelling mechanism and gelling time allow for
delivery to the ventricle wall using minimally invasive, simple catheter injection. A full list
describing the studies considered in the comparative-analysis is provided in Table C-2
(Appendix C), and graphical representation of the data is presented in Figure C-5 (Appendix C).
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Table 4-1: Comparative-analysis of reported MI studies
%Relative change†: Median MeanStandard
Deviation# of
Groups QHG213H Scar Thickness 67% 69% 22% 7 53% Fractional Scar Area -48% -52% 21% 5 -34% Fractional Shortening 41% 53% 33% 6 36% Ejection Fraction 43% 54% 40% 6 62% *(Wu et al. 2011; Memon et al. 2005; Landa et al. 2008; Tsur-Gang et al. 2009; Christman, Fok, et al. 2004; Dai et al. 2005; Lu et al. 2009; T. Wang, Jiang, et al. 2009; Fujimoto et al. 2009; Kraehenbuehl et al. 2011) †Data is presented here as %relative change in treatment vs control groups at end point of given studies.
4.4. Summary
SA-2 Assess effect of the developed hydrogel in a rat MI model.
This is the first study to evaluate the effect of the angiopoietin-1-derived peptide QHREDGS
on remodelling post-MI. Injection of the developed QHREDGS peptide modified chitosan-
collagen (QHG213H) hydrogel into the LV of rats with acute MI showed significant
improvement in cardiac morphological and functional measurements at 6 weeks post MI relative
to the Control gel (no peptide) or MI Only (PBS) injections. Scar thickness was improved by
53%, fractional scar area reduced by 34%, fractional shortening improved by 36% and ejection
fraction improved by 62% compared to PBS injected hearts. The in vivo lifespan of the hydrogel
was assessed using an on-line technique and localization of the injected hydrogel to the LV wall
of treated hearts was demonstrated. The retention of the hydrogel in situ for up to two weeks was
also demonstrated and removal of the majority of the hydrogel was determined to occur by 3
weeks post injection. The injected QHREDGS-modified chitosan-collagen hydrogel was also
found to maintain a higher proportion of CMs in the MI zone, and of large mature vessels within
the border zone of infarcted hearts. We postulate that binding of cardiac cells to QHREDGS
peptide via β1-type integrin receptors on the cell surface promotes constructive acute
inflammation, inhibiting apoptosis, and attenuates adverse remodeling resulting in improved
cardiac functional outcomes post MI. While effects of Angiopoietin-1 are considered to be
largely Tie2 receptor mediated, the novelty of our work lies in the expanded insight into integrin-
mediated mechanisms by which Angiopoientin-1 derived peptide, QHREDGS, acts on cardiac
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cells in vivo. The developed delivery system, consisting of collagen:chitosan hydrogel with
covalently immobilized QHREDGS, could also form a basis for a new therapeutic that is fully
chemically defined and based on biocompatible, biodegradable and inexpensive molecules.
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5. Recommendations for future work
This thesis describes the work that has been done towards development of a hydrogel for
cardiac tissue engineering (Specific Aim SA-1), culminating in the publication of Reis et al.
(2012). It continues to describe much of the work done in testing the developed hydrogel in vivo
in a rat model of acute myocardial infarction (Specific Aim SA-2); the original work described in
Reis et al. (2015). Future work must further address SA-2. iii (determining mechanism of action).
5.1. Specific Aim SA-2. i. Show ability to localize in vivo and quantify duration of
its in vivo presence.
Further to the discussion of the in vivo biodistribution of the hydrogel and our data concerning
its presence within the heart one more questions remains: does the peptide modified chitosan
affect any other organs in the body? As seen in the previous discussion, once cleared from the
heart (through degradation of collagen) chitosan and peptide modified chitosan is likely then
processed and rapidly cleared through animals excrements. However, an in depth study of this
has not been done. We have not yet assessed whether addition of the peptide to chitosan makes it
more likely to localize in the kidneys or liver, or even if rapidly cleared does processing it affect
either organ. Recent work in our lab points to the survival promoting activity of QHREDGS
peptide with a number of different cell types (e.g. CMs, FBs, iPSCs, ECs) thus it is likely to
interact with other cells in the body. It is likely that in vitro culture of either liver or kidney cells
in the presence of soluble QHREDGS-chitosan will show if interaction produces any significant
effects, however it would be more interesting to see these effects in vivo in context of post-MI
healing and inflammatory milieu that is present after such insult. Collecting and analyzing these
tissues in tandem with heart tissues during an MI model would provide further insights into the
in vivo biodistribution and life span of our hydrogel.
5.2. Specific Aim SA-2. iii. Determine mechanism of action.
There remain a number of avenues that need to be explored in further determining the
mechanism through which we see improvement after MI with peptide modified hydrogel
treatment. First, the α-integrin sub-unit through which cardiomyocytes attach to QHREDGS
peptide is still undetermined. Similar sub-unit blocking experiments as described in determining
the β-integrin sub-unit should be performed to determine the complement α-sub-unit (many
purified antibodies readily available from BD Biosciences).
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Furthermore, the pathway activated through integrin binding to QHREDGS peptide is not
well or fully characterized. Our results point to higher IL-6 and IL-1β with peptide treatment
(neither were statistically significant), although MAPK activation is implicated from our Western
blot analysis. Recent work assessing inflammatory cytokines in the Border Zone tissues excised
24 hours after treatment were similarly inconclusive and suggests two possibilities: (i) the
hydrogel treatments do not elicit a significantly different inflammatory response than PBS
injection after infarct, or (ii) the time at which animals were sacrificed (24 hrs) was not ideal for
detecting those differences. Full understanding of this activation pathway must be determined.
Cardiomyocytes were the cell type largely studied throughout the current work to understand
their interaction with QHREDGS peptide. However, it is known from our in vitro work that
cardiac endothelial cells and various other non-cardiac cells interact positively with QHREDGS
(Miklas et al. 2013; Dang et al. 2014; Feric et al. 2014). In addition, we recently showed cardiac
fibroblasts bind to QHREDGS peptide through β1-integrins and we have little understanding of
the peptides interaction with inflammatory cells such as monocytes and macrophages.
We have not worked with monocytes before; however work has been done by other groups
into the Ang1/Tie-2 system in human monocyte chemotaxis, attachment, and M1/M2
differentiation (Ahmad et al. 2010; Seok et al. 2013). The reported work found that Ang-1
induces chemotaxis of monocytes in a manner that is independent of both Tie-2 and integrin
binding. In addition, they showed that Ang-1 binding to human umbilical endothelial cells was
partially Tie-2 and integrin dependent, whereas Ang-1 binding to monocytes was independent of
both integrins and Tie-2 (Ahmad et al. 2010). In terms of monocyte differentiation they found
that Ang1 is a potent stimulator of p38 and Erk1/2 MAPKs, induces pro-inflammatory monocyte
activation, and that this is Tie-2 independent. The Ang1 mediated stimulation promoted TNF-α
expression as well as a slight increase in IL-6. It was also shown that Ang1 switched macrophage
differentiation toward a pro-inflammatory M1 phenotype, even in the presence of an anti-
inflammatory (pro M2 phenotype) mediator. As we saw slight increases in IL-6 and MAPK in
our system it may suggest that monocytes do indeed attach to QHREDGS peptide, through
integrins or otherwise, and this may promote pro-inflammatory responses. However, none of our
results showed any statistically different expression levels among groups to suggest M1 vs. M2
macrophage activation.
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The in vitro work also does not take into account the great amount of cross-talk between
various cell types and the complex immune response post-MI that may negate or promote any
effect the peptide may have on any one of the various cell types present. Basic mechanistic
understanding of the anti-apoptotic effect of QHREDGS peptide on various cell types (e.g.
monocytes, CMs, FBs, ECs) could be done in vitro using one of many ischemia models and up-
regulation of various cytokines assessed, however the problem would remain of how these
cytokines affect the post-MI immune/inflammatory response (Wei et al. 2013; Yue et al. 2000;
Ma et al. 2006). This can only be assessed with large scale in vivo models, which themselves
pose many issues. In the simplest case, where our tissue collection time points were not ideal to
assess effect, a large study would be required to sacrifice many animals at many points during
the acute phase after MI (first few days) to better understand the up/down-regulation and
expression of the many proteins involved and find statistical significance. If our original
hypothesized mechanism is true then understanding the exact roles IL6 or IL1β play in terms of
cell recruitment and what those infiltrating cells do post-MI becomes more complex. Both the
control case of MI Only (PBS injection) and the treatment (peptide modified gel injection) elicit
an immune/inflammatory response, as does the MI surgery itself. The classic models of using
knock-out animals or antibodies against specific proteins will not work when both the control
and treatment case elicit similar responses. For macrophage infiltration to MI zone and
assessment of M1/M2 phenotype may require autologus intravenous injections of labelled cells
and tracking outcomes.
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6. Conclusions & Contributions to New Knowledge
As discussed previously, the issue of cardiac disease and the need for clinically translatable
treatment options is one of the most pressing health concerns of developed society. Post MI
pathological cardiac remodeling activities include infarct expansion, ventricular dilation, and
scar tissue formation. While necessary for LV stabilization and immediate survival, these
activities are very disconcerting in the short and long term. Overall, the dilation of the ventricle
has been shown to directly decrease patients’ survival and may ultimately lead to congestive
heart failure. The extent of remodeling is affected by the size of the infarct, how much the infarct
heals, and ventricular wall stresses.
There are routine clinical treatments which attempt to prevent the pathological remodeling
that occurs post-infarction. These methods include both pharmacological methods that aim to
inhibit remodeling pathways, as well as targeting the cause of MI (blocked or partially blocked
coronary arteries). Reperfusion of blocked arteries (through methods such as thrombolysis and
angioplastys) to restore blood flow to the infarcted myocardium is standard procedure which has
been shown to have beneficial effects including limiting detrimental remodeling and partially
healing the myocardium (Sutton and Sharpe 2000; Touchstone et al. 1989; Marino, Zanolla, and
Zardini 1989; Baks et al. 2006). While promising, reperfusion is not without issues as the
mechanism of repair is not fully understood and restoring blood flow to injured cells can lead to
further damage and have severe detrimental effects (Verma et al. 2002). Pharmacological
methods used to treat MI are only partially effective and there are many questions about the
safety of long term use (Sutton and Sharpe 2000; Beltrame 2008; McMurray et al. 2006;
Braunwald et al. 2004), even though they have shown effectiveness in the short term.
Currently, no methods to completely prevent these pathological events exist, however
researchers all over the world are exploring approaches and treatment options that can stabilize
the scar faster, reduce the expansion of the infarct, and decrease ventricular wall stresses that
lead to further detrimental remodeling events. Left ventricular assist devices have improved
greatly over the years and are increasingly used as a bridge to transplantation, and some have
been approved as a destination therapy, providing individuals ineligible for a heart transplant
with a long-term therapy option (Wilson et al. 2009; Kirklin et al. 2013). However this is not the
most desirable option and it would be preferred to prevent heart failure altogether, thus
100
treatments that limit the remodeling that occurs post-infarction will improve the survival and
quality of life of MI patients (Pfeffer and Braunwald 1990).
To address these issues we sought to utilize the novel Ang1 derived peptide QHREDGS along
with two common biomaterials, chitosan and collagen, to create a novel treatment option for MI.
Chitosan and collagen have been used before in the treatment of cardiac dysfunction and tissue
engineering as whole (Lu et al. 2009; Rask, Dallabrida, et al. 2010; Garbern et al. 2011;
Suuronen et al. 2006; Y. Zhang et al. 2008; Wu et al. 2011). However, typical formulations of
chitosan-collagen hydrogels use some sort of crosslinking mechanism to speed gelation, improve
mechanical integrity, and perhaps make them more amenable for use in vivo. Our first success in
this project was demonstrating that combinations of the two components could form
mechanically stable hydrogels, without the use of exogenous chemical crosslinkers, suitable for
CM cell culture, and that specifically a 2.5 mg/mL 1:1 m:m ratio was optimal for culturing
cardiac cells. Furthermore, the functional groups on both chitosan and collagen, as well as their
ionic nature, allows for not only conjugation of various molecules (e.g. QHREDGS) but also
release of small molecules in a drug delivery system. Work concurrent to this showed that a 2:1
collagen to chitosan ratio was optimal for the controlled release of Tβ4 and suitable for the
endothelial and smooth muscle cell growth into vascular structures from epicardial explants
when cultured in the Tβ4 containing gel (Chiu and Radisic 2011; Chiu, Montgomery, et al.
2012).
We were subsequently able to demonstrate that the hydrogel was able to localize at the site of
injection in the LV of live animals and track its in vivo biodistribution on line to show its lifespan
in the heart to be 2-3 weeks. While many would dismiss the relatively low mechanical properties
of our gel (in terms of elastic storage modulus) as being irrelevant for use in the heart we were
successful in demonstrating otherwise. To our knowledge this is the first time a non-exogenously
cross-linked chitosan-collagen blend has been used in the heart. As mentioned previously
varying the ratio of chitosan and collagen also allows for a level of tailoring of the mechanical
and ionic properties of the gels for use in other applications.
The most important findings of this work, however, had to do with the potential of
QHREDGS peptide in treating acute MI as well as furthering the understanding of the underlying
mechanisms through which it acts. Our in vitro work clearly demonstrated stable conjugation of
QHREDGS to chitosan provided long-term signaling as well as measurable improvements in CM
101
performance when encapsulated in peptide modified chitosan-collagen hydrogels. We showed
that CM attachment to QHREDGS is indeed integrin dependant and implicated the β1-integrin
subunit as being integral to attachment and signaling. Our in vivo results clearly demonstrate the
potential of this novel hydrogel in treating acute MI to maintain higher levels of cardiac function
and reduce pathological remodeling. A major contribution of this work is in the understanding of
how both chitosan-collagen hydrogel and peptide-modified hydrogel succeed in producing these
post-MI benefits, in terms of maintaining higher levels of CMs within the developed scar and
reducing scar expansion. From our results it would seem injection of either gel does not induce a
significantly different immune response from PBS injection post-MI, but does result in better
cardiac outcomes. Furthermore, we were able to indentify some trends in inflammatory cytokine
expression that warrants further investigation and raises interesting observations about cell
recruitment and survival mechanisms in the acute phase after infarct.
In looking forward, this work has provided the basis for a novel MI treatment and opened up
new research avenues into how the heart responds to gel treatment after infarct, peptide
immobilization and the long-term impacts of cell exposure to such compounds, and whether this
formulation can have similar successes in larger animal models – leading to the question can it
be successful clinically?
6.1. Publications & Contributions
a. Journal articles
Xiao Y., Reis L.A., Zhao Y., & Radisic M. (2015). Modifications of biomaterials with
immobilized growth factors or peptides for tissue engineering applications. Methods.
Submitted.
Zhang B., Montgomery M., Chamberlain M.D., Wells L.A., Pahnke A., Massé S., Kim J.,
Reis L.A., Abdulah M., Nunes S.S., Nanthakumar K., Sefton M.V., & Radisic M.
(2014). AngioChip: a biodegradable scaffold with built-in vasculature for organ-on-a-
chip engineering and direct surgical anastomosis. Nature Materials. Under Review.
Reis L.A., Chiu L.L.Y., Wu J., Feric N., Momen A., Li R.K., & Radisic M. (2015).
Hydrogels with integrin-binding angiopoietin-1-derived peptide, QHREDGS, for
treatment of acute myocardial infarction. Circulation: Heart Failure. DOI:
10.1161/CIRCHEARTFAILURE.114.001881.
102
Reis L.A., Chiu L.L.Y., Feric N., Fu L., & Radisic M. (2014). Biomaterials in myocardial
tissue engineering. Journal of Tissue Engineering & Regenerative Medicine. DOI:
10.1002/term.1944.
Miklas J.W., Nunes S.S., Sofla A., Reis L.A., Pahnke A., Xiao Yun., Laschinger C., &
Radisic M. (2014). Bioreactor for modulation of cardiac microtissue phenotype by
combined static stretch and electrical stimulation. Biofabrication. 6: 024113.
Miklas J.W., Dallabrida S.M., Reis L.A., Ismail N., Rupnik M., & Radisic M. (2013).
QHREDGS enhances tube formation, metabolism, and survival of endothelial cells in
collagen-chitosan hydrogels. PLOS One. 8(8): e72956.
Chiu L.L.Y., Reis L.A., Momen A., & Radisic M. (2012). Controlled release of thymosin
β4 from injected collagen-chitosan hydrogels promotes angiogenesis and prevents
tissue loss after myocardial infarction. Regenerative Medicine. 7(4): 523-33.
Chiu L.L.Y., Iyer R.K., Reis L.A., Nunes S.S., & Radisic M. (2012). Cardiac Tissue
Engineering: Current State and Perspectives. Frontiers in Bioscience. 17: 1533-1550.
Reis L.A., Chiu L.L.Y., Liang Y., Huynh K., & Radisic M. (2011). A peptide-modified
chitosan-collagen hydrogel for cardiac cell culture and delivery. Acta Biomaterilia.
8(3): 1022-36.
Iyer R.K., Chiu L.L.Y., Reis L.A., & Radisic M. (2011). Engineered cardiac tissues.
Current Opinion in Biotechnology. 22(5): 706-714.
Rask F., Mihic A., Reis L. A., Dallabrida S., Ismail N., Sider K., Simmons C., Weisel R.,
Li R.K., & Radisic M. (2010). Hydrogels modified with QHREDGS peptide support
heart cell survival in vitro and after sub-cutaneous implantation. Soft Matter. 6: 5089-
99.
b. Book chapters
Reis L.A., Chiu L.L.Y., Feric N., Fu L., & Radisic M. (January 2014). Injectable
biomaterials for cardiac repair. In Li R.K. & Weisel R.D. (EDs), Cardiac
Regeneration & Repair (Vol. II, Ch. 3). Cambridge: Woodhead Publishing.
Audet J., Clarke G., Lee W., Ma J., Onishi K., Reis L.A., & Zandstra P. (2013).
Combining in silico and in vitro techniques to engineer pluripotent stem cell fate. In
Micou M.K. & Kilkenny D. (EDs.), A Laboratory Course in Tissue Engineering (1st
ed., p. 173-196). Boca Raton: CRC Press.
103
Odedra D., Chiu L.L.Y., Reis L.A., Rask F., Chiang K., Radisic M. (2011). Cardiac Tissue
Engineering. In J. A. Burdick & R. L. Mauck (EDs.), Biomaterials for Tissue
Engineering Applications: A review of past & future trends (1st ed., p. 562). Vienna:
Springer Vienna.
c. Oral presentations
Reis L.A., Chiu L.L.Y., Wu J., Feric N., Momen A., Li R.K., & Radisic M. (May 2014). A
peptide modified hydrogel therapy for acute myocardial infarction. 2014 Annual
University of Toronto Institute of Biomaterials & Biomedical Engineering Scientific
Day, Toronto, ON, CAN.
Reis, L.A., Chiu, L.L.Y., Wu, J., Momen A., Li, R.K., & Radisic, M. (Oct. 2012). A
peptide modified hydrogel for cardiac tissue preservation after acute myocardial
infarction. 2012 Biomedical Engineering Society Annual Meeting, Atlanta, GA, USA.
Reis, L.A., Chiu, L.L.Y., Wu, J., Momen A., Li, R.K., & Radisic, M. (April 2012). A
peptide modified hydrogel for cardiac regeneration. 5th Annual Regenerative
Medicine Symposium, Toronto, ON, CAN.
d. Poster presentations
Reis L.A., Chiu L.L.Y., Wu J., Momen A., Li R.K., & Radisic M. (May 2013). A peptide
modified hydrogel for treatment after acute myocardial infarction. 2013 Canadian
Biomaterials Society Annual Meeting, Ottawa, ON, CAN.
Reis L.A., Chiu L.L.Y., Wu J., Momen A., Li R.K., & Radisic M. (September 2012). A
peptide modified hydrogel for cardiac tissue preservation after acute myocardial
infarction. An Afternoon of Engineering Innovation, Toronto, ON, CAN.
Reis L.A., Chiu L.L.Y., Wu J., Momen A., Li R.K., & Radisic M. (June 2012). A peptide
modified hydrogel for cardiac regeneration. CIHR 2012 Canadian Student Health
Research Forum, Winnipeg, MB, CAN.
Reis L.A., Chiu L.L.Y., Wu J., Momen A., Li R.K., & Radisic M. (May 2012). A peptide
modified hydrogel to reduce long term effect of acute myocardial infarction. 2012
Annual University of Toronto Institute of Biomaterials & Biomedical Engineering
Scientific Day, Toronto, ON, CAN.
Reis L.A., Chiu L.L.Y., Liang Y., Huynh K., Momen A., & Radisic M. (May 2011). A
peptide modified collagen-chitosan hydrogel for the site-specific delivery of
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cardiomyocytes to the heart. 2011 Annual University of Toronto Institute of
Biomaterials & Biomedical Engineering Scientific Day, Toronto, ON, CAN.
Reis L.A., Huynh K., Momen A., & Radisic M. (June 2010). Development of a peptide
modified collagen-chitosan hydrogel for the site-specific delivery of cardiomyocytes to
the heart. 2010 Canadian Biomaterials Society Annual Meeting, Kingston, ON, CAN.
105
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Appendix A. Reaction schematics and chemical structures
Figure A-1: Polyelectrolytic complexation of chitosan and collagen
(A) Collagen has both carboxyl and amine groups, although it has many more negatively charged carboxyl groups (pI~4.8). In acidic medium most carboxyl groups are quenched and sparse charge along collagen fibers exist. Increasing the pH increases ionic strength and attraction between collagen fibers, and causes spontaneous organization through both covalent bonding and ionic interaction between collagen fibers (only ionic interactions shown). As the reaction is endothermic, an increase in temperature greatly assists and speeds up the gelation process. (B) Addition of chitosan to a collagen mixture greatly increases the number of free, charged amines but does not interfere with the spontaneous organization of collagen fibers (which promotes the formation of a hydrogel). Chitosan associates with collagen fibers through ionic interaction, forming thicker chitosan-collagen complexed fibers. (C) Further addition of chitosan creates even thicker overall chitosan-collagen fibers with chitosan essentially quenching all available negative charge on collagen fibers, and further sequesters collagen fibers. (A-C) Collagen only hydrogels form highly interconnected networks with thinner individual fibers, meaning the hydrogels are easily deformed but do not break apart easily. The addition of chitosan creates larger overall fibers, however reduces the number of cross-links that can be formed between network fibers. The resultant hydrogel resists deformation more than pure collagen and retains network strength to resist tearing. Too much chitosan essentially quenches all available carboxyl groups on collagen, forming very clustered fibers (thick), but very low fibre network interconnectivity. The resultant hydrogel is therefore very brittle. (Inspired by Berger et al. (2004)).
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Figure A-2: Modified chitosans used
Figure A-3: EDC/Sulfo-NHS reaction chemistry for conjugation of biomolecules
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Figure A-4: QHREDGS conjugation to PEG
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Appendix B. In vitro construct CM distribution
Figure B-1: In vitro cell leakage and migration Brightfield images of cell/gel constructs seeded in 12 well tissue culture plates at 1hr (30 minutes post
gelation) and after 120hrs in culture. Low magnification image (top row) and high magnification (bottom) show minimal cell leakage out of the gel during seeding and gelation as well as after 5 days in culture. Scale bars=200µm top, 50µm bottom left, 100µm bottom right.
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Figure B-2: Quantification of in vitro construct viability (A) Live/dead (PI, red/CFDA, green) images of Control and High peptide constructs grown in vitro after
120 hrs. Images were taken at the edge of constructs and in the very center (middle). Scale bars=100µm. (B) Live/dead quantification was performed using ImageJ and shows no significant differences in viability between groups at the two regions, however the edge region has significantly higher viability then in the center of constructs (N=4/group). Data expressed as Mean±SD, ** P<0.01, *** P<0.0001.
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Figure B-3: High QHREDGS hydrogel constructs contain elongated cardiomyocytes with visible cross-striations
High magnification cardiac troponin T immunostaining of in vitro formalin fixed paraffin embedded tissue constructs. cTnT (green), DAPI (blue). Arrow indicates cross-striations. Scale bars=25µm.
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Appendix C. Utility of QHG213H hydrogel for MI treatment published
supplementary data
Table C-1: Genes and primers used in RT-qPCR
Gene Abbr. Forward Reverse
Glyceraldehyde 3-phosphate dehydrogenase
GAPDH GCTGAGAATGGGAAGCTG AGTGATGGCATGGACTGT
B-cell lymphoma 2 BCL2 CTGGTGGACAACATCGCTCTG GGTCTGCTGACCTCACTTGTG
B-cell lymphoma extra large
BCL-XL CATATAACCCCAGGGACAGC GTCATGCCCGTCAGGAAC
Phosphatidylinositol-4,5-bisphosphate 3-kinase
PI3K AGCCACAGGTGAAAATACGG TTTTCTTTCCGCAACAGCTT
BH3 interacting-domain death agonist
BID AGCAGGTGATGAACTGGACC AGACGTCACGGAGCAGAGAT
BCL2-associated X protein BAX TCCAGGATCGAGCAGA AAGTAGAAGAGGGCAACC
Caspase 3 CASP-3 AACCTCAGAGAGACATTCATGG GAGTTTCGGCTTTCCAGTCA
Caspase 9 CASP-9 CTCCTGCGGCGATGC CCACTGGGGTGAGGTTTC
Receptor interacting protein 1
RIP1 GTGGTGAAGCTACTGGGCAT AGGAAGCCACACCAAGATCG
Receptor interacting protein 3
RIP3 CACACCAGCAGGGACATCAT TTTGGTGCGTTCCAGGTGTA
Tumor necrosis factor α TNFα TGCCTCAGCCTCTTCTCATT GCTTGGTTTGCTACGAC
Interleukin 1β IL-1β CTGTGACTCGTGGGATGATG GGGATTTTGTCGTTGCTTGT
Interleukin 6 IL6 TCCTACCCCAACTTCCAATGCTC TTGGATGGTCTTGGTCCTTAGCC
Intercellular adhesion molecule 1
ICAM1 ACAAGTCCGTGCCTTTAGCTC GATCACGAAGCCCGCAATG
Vascular cell adhesion molecule 1
VCAM1 TAAGTTACACAGCAGTCAAATGGA CACATACATAAATGCCGGAATCTT
Matrix metalloproteinase 2 MMP2 AAGTCTGAAGAGTGTGAAGT GTGAAGGAGAAGGCTGATT
Transforming growth factor β1
TGFβ1 TGCTTCAGCTCCACAGAGAA TGGTTGTAGAGGGCAAGGAC
Monocyte chemoattractant proetein 1
MCP1 TTCACAGTTGCTGCCTGTAG TCTGATCTCACTTGGTTCTGG
Macrophage migration inhibitory factor 1
MIF1 CCA GGA CCG CAA CTA CAG CAA GGG CTC AAG CGA AGG TGG AAC
Interleukin 10 IL10 TTGAACCACCCGGCATCT CCAAGGAGTTGCTCCCGTTA
Tissue inhibitor of matrix metalloproteinase 1
TIMP1 CGCGGGCCGTTTAAGG GTTTGCAAGGGATGGCTGAA
Vascular endothelial growth factor
VEGF CCCACGACAGAAGGAGAGCA GCACACAGGACGGCTTGAA
Cardiac troponin T cTnT AGGCTCTTCATGCCCAAC TGTTCTCGAAGTGAGCCTC
α myosin heavy chain αMHC TGATGACTCCGAGGAGCT TGACACAGACCCTTGAGCA
Β myosin heavy chain β MHC CCTCGCAATATCAAGGGAAA TACAGGTGCATCAGCTCCA
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Integrin linked kinase ILK CCCCACGTGTGTAAGCTCAT AGCCCCATAACTGGGGTAGT
Integrin subunit α5 α5 CCAGAGCAAGAGCCGGATAG GCGTTCTTGTCACCCAGGTA
Integrin sub-unit β1 β1 CAATGAGGGTCGTGTTGGGA CCGTTGGACCTATCGCAGTT
Table C-2: Description of studies considered for comparative-analysis
Group/Paper Control Treatment Treatment
Time (After MI)
Monitoring Time (After Treatment)
(Christman et al. 2004) 0.5% BSA in PBS
Skeletal myoblasts in Fibrin glue 1 weeks 5 weeks
(Memon et al. 2005) DMEM injection
Myoblast cell sheets 2 weeks 8 weeks
(Dai et al. 2005) Saline injection Bovine dermal collagen 1 week 6 weeks
(Landa et al. 2008) Saline injection Ca-cross linked alginate 1 week 8 weeks
(Tsur-Gang et al. 2009) Saline injection Ca-cross linked alginate 1 week 8 weeks
(Lu et al. 2009) PBS mESCs in temperature responsive chitosan-β-GP hydrogel
1 week 4 weeks
(Wang et al. 2009) DMEM/F12 medium
Bone marrow stem cells in α-CD/MPEG-PCL-MPEG hydrogel
1 week 4 weeks
(Fujimoto et al. 2009) PBS Thermoresponsive poly(NIPAAm-co-AAC-co-HEMAPTMC)
2 weeks 8 weeks
(Wu et al. 2011) PBS VEGF conjugated PVL-b-PEG-PVL hydrogel
1 week 4 weeks
(Kraehenbuehl et al. 2011)
PBS hESC derived ELCs & SMLCs in Tβ4-PEG hydrogel
1 hour 6 weeks
Our work PBS Chitosan-collagen-QHREDGS hydrogel
Imm. 6 weeks
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Figure C-1: Quantification of hydrogel lifespan Whole heart fluorescence was quantified from homogenized hearts excised 1 hr and 14 days after injection with Dylight-800
labelled chitosan-collagen gel. Assuming that after 1 hr the injected material has gelled fully and that unconjugated dye has been removed along with un-gelled material by this time, the ~60% reduction in fluorescence on day 14 is indicative that ~40% of the labelled hydrogel localized in the heart remained after two weeks.
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Figure C-2: Histological staining 6 weeks post MI (A) Mason’s trichrome, (B) smooth muscle actin (SMA), (C) CD31, and (D) Factor VIII (FVIII) stained
sections of hearts excised 6 weeks post MI. Black outlined sections in (A) are enlarged in (B-D). A large collagenous scar is seen in all groups (A), but there appears to be more SMA, CD31 and FVIII staining in the Control gel and QHG213H gel groups as compared to MI Only (PBS) group, suggesting increased vascularization (B-D). Scale bars = 1 mm (A), 50 µm (B-D), N=7/group.
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Figure C-3: RT-qPCR of MI and border zone tissues from hearts excised 24 hrs post MI Hearts subjected to MI and immediate treatment were excised 24 hrs later, sectioned into MI, border, and remote regions,
snap frozen then ground and used for RT-qPCR. Overall, 27 genes conventionally associated with apoptosis, necroptosis (Necrop), survival, anti- and pro-inflammatory responses, as well as cardiac specific genes were assessed. Data was standardized in samples to reflect changes in gene expression in the progressing border zone to the established MI zone. No significance was determined between groups. Data expressed as Mean±SD, N=5/group.
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Figure C-4: Western blot analysis of border zone tissue from hearts excised 24 hrs post MI
Western blot analysis of border zone heart tissues showed no differences in the levels of ILK or phospho-MAPK (PMAPK) protein expression between groups. The MAPK protein expression was significantly increased in the QHG213H gel group compared to the Control gel and MI Only (PBS) groups. Data expressed as Mean±SD, * P<0.05, N=5/group.
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Figure C-5: Reported cardiac morphological and functional data from studies used in meta-analysis The differences between Control treatment (Saline or PBS injection) and therapy (hydrogel injection, cell treatment, patches)
in various measures of cardiac remodeling and function after acute MI as reported by difference studies are illustrated (Top panel, and bottom left). The relative change in the outcome of these measures between all groups is illustrated (bottom right).
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Appendix D. Unpublished rat MI model supplementary material
After sacrificing, excising, and fixing hearts it was fairly obvious that the MI Only and
Control gel groups were very dilated in comparison to the Sham operated heart (consistent with
the results seen in the mouse MI model described earlier), but that the QHG213H gel injected
heart appeared much more similar in size (Figure D-1 A). Two cross sections of each heart were
taken in the area just below the suture and illustrate the extent of MI in each group (white
scarring, Figure D-1 B) and is made further evident in the Mason’s trichrome staining done on
all sections (Figure D-1 C). The blue collagenous scar deposition seen in the trichrome stained
sections seemed much less apparent in the QHG213H gel group compared to MI Only and
Control gel, and overall tissue cross section appeared very similar to that of the healthy heart
(Sham).
Figure D-1: Heart morphology & histology 3Wk post MI A LAD rat MI model was used to assess hydrogel utility in mitigating remodeling. (A) Ligation site is
clearly marked by the suture and surrounding MI zone (white), and remodeling extent is further apparent in tissue cross sections (B). Scale bars=1mm (A-B). (C) Trichrome images show tissue remodeling and extent of infarct. Scale bars=1mm. Black outlined sections are area of focus for Figure D-2. (D). N=2 Sham, N=6 rest.
Staining for H&E, SMA, F8, and CD31 was also performed on all 3Wk sections (Figure D-2
A-D respectively). SMA expression in the MI zone of all test groups appeared greater than
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endogenous expression seen in the Sham hearts, although expression in the QHG213H group
seemed localized to the gel injection site and appeared to be much less expression than the MI
Only and Control gel groups (Figure D-2 A). F8 and CD31 expression, both endothelial cell
markers used to identify neovascularization, were similar within groups and were consistent with
the trend observed in the SMA stained sections (Figure D-2 C&D). Interestingly, however, in the
QHG213H gel group both F8 and CD31 staining appeared limited to the gel border and
infiltrating cells were not seen within the gel as seen with the SMA staining.
Figure D-2: Further histological staining 3Wks post MI (A) H&E, (B) SMA, (C) F8, & (D) CD31staining illustrate infiltrating myofibroblasts and neovascularization
in the MI zone. Marked differences can be seen between all test groups and the healthy, Sham tissue. MI Only and Control gel groups have very similar expression profiles for all stains, but QHG213H gel group appears to have much more normal expression of all markers, although clear infiltration of SMA + cells into the injection site can be seen. Scale bars=50µm, N=2 Sham, N=6 rest.
Many different metrics were used in an attempt to quantify the extent of remodeling at 3Wks
post MI. First, left ventricle (LV) wall thickness in the MI zone was measured at numerous
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points (~5 locations per sample) to get an average value per section, then sample, and mean
value per group (Figure D-3 A&B). An obvious trend was seen with the MI Only group having
the lowest MI zone LV wall thickness, followed by the Control and then QHG213H gel groups;
however no significant difference between measurements was apparent Figure D-3 B). Adobe
Photoshop CS3 was used quantify MI scar area and total tissue area from trichrome stained
images (shaded areas, Figure D-3 C&D). Values were used to get a measure of fractional scar
coverage (MI scar area/total tissue area) between groups, and compare tissue loss (or
maintenance) by comparing mean total tissue area between groups (Figure D-3 D&F). The trend
of the MI Only group having the greatest fractional scar coverage and lowest mean total tissue
area, and the QHG213H gel group having the best was consistent with the MI zone LV wall
thickness data, however a significant difference was only seen between MI Only and QHG213H
gel groups in terms of mean total tissue area (p=0.007). Both the interior and exterior scar extent
were also measured as a final metric of remodeling at 3Wks, as different groups report many
different metrics and therefore they were performed for completeness (Figure D-4). Both
measures maintained the overall trend seen in Figure D-3, however neither are statistically
significant.
While only one measure (total tissue area) showed statistical significance in the ability of
QHG213H gel to limit remodeling at 3Wks, this time point was only an intermediate and it was
not entirely surprising we did not see more significance between groups. As previously stated, in
both mouse and rat MI models pathological remodeling is not complete until six weeks post MI,
therefore while the trends observed at 3Wks were promising the same metrics at 6Wks were the
truly important measures (Patten et al. 1998; Krzemiński et al. 2008).
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Figure D-3: 3Wk MI quantification measurements (A) MI zone wall thickness measurements were made at various spots (~5) along the scar length and
averaged to get a measurement per sample and (B) mean LV wall thickness between groups. (C) MI scar area was measured and normalized to overall heart tissue area per sample (E) to get a measure of (D) fractional scar coverage between groups, and to compare (F) the maintenance or loss of heart tissue. **P<0.01, one way ANOVA, Bonferroni post tests. Data expressed as Mean±SD, N=6/group.
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Figure D-4: 3Wk MI scar extent assessment Interior (A, red line) compared to LV capacity (perimeter, blackline). Shows clear trend towards lower extent
in QHG213H group (B) which correlates with rest of data. Exterior, same thing (C&D). Data expressed as Mean±SD, N=6/group.
Stained sections from the hearts excised 6 weeks post treatment were imaged using a slide
scanner at Princess Margaret Hospital, allowing for much higher resolution and greater
measuring accuracy when quantifying the extent of remodeling from 6Wk trichrome images
compared to those from 3Wk. The same metrics were used to quantify remodeling extent at
6Wks as done at 3Wks and the same trend remained with QHG213H gel group having the
highest scar thickness, lowest fractional scar tissue coverage, and highest LV tissue area,
followed by Control gel and then MI Only (Figure D-5 A-C, respectively) however the results
were not significant as seen with the quantification done on the gross tissue sections (Figure 4-6).
Two further metrics were assessed; the relative scar thickness compared to healthy LV and the
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infarct expansion index (Figure D-5 D&E respectively). The two graphs show the expected trend
of QHG213H gel performing best, followed by Control gel and finally MI Only (PBS), but there
remained a lack of significance between groups.
Figure D-5: 6Wk MI quantification measurements (from histology pictures)
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Measurements done in the same manner as that for 3Wk samples, however with higher resolution for greater accuracy and more MI zone wall thickness measurements (>10 per sample) to get overall (A) MI zone scar thickness, (B) fractional MI zone scar tissue coverage (with respect to left ventricle (LV) area only), and (C) total LV tissue area. LV area was considered for the 6Wk samples as some RVs were missing from sections, and it was much easier to eliminate RV than it was for 3Wk samples. Data expressed as Mean±SD, N=4/group.
Quantifying differences in vasculature in the MI zone at 3Wks post MI was done using SMA,
F8, and CD31 stained sections. Mean vessel density and diameters between groups were
compared for mature SMA expressing vasculature at 3Wks (Figure D-6 A), or total vasculature
(Figure D-6 B). While images may have hinted at a trend being apparent, the data suggested
there were no differences in vascularization between groups, with large standard deviations
indicating there was a large difference within the groups themselves.
Figure D-6: Vascularization within the MI scar zone at 3Wks post infarct Vascularization within the MI zone was quantified at 3Wks. Vessel density and diameter were quantified
from SMA stained sections to assess mature vasculature within the MI zone (A) and from F8 (B) stained sections for neovascularization. No discernible trend can be seen and there is no significant difference between groups. Data expressed as Mean±SD. N=5 MI Only, N=6 Control, N=3 QHG213H.
The TUNEL label and enzyme are used to label DNA strand breaks for detecting and
quantitating late stage apoptotic cell death at a single cell level. It was used to determine CM cell
death at the MI border zone and see if the functional and morphological improvements seen
could be due to reduced apoptosis and maintenance of healthy CMs. TUNEL/cTNT staining
showed a limited number of apoptotic cells in all groups at 3Wks (Figure D-7 A), and perhaps a
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reduced number in the Control group compared to MI Only and QHG213H groups, but no
definitive conclusions could be made (Figure D-7 B&C). Counting was done using a macro
developed using Olympus CellSens Dimension software tested on TUNEL controls (Figure D-8).
Figure D-7: Apoptosis in the MI boundary zone at 3Wk Late stage apoptosis was assessed by staining sections for TUNEL (red), cTnT (green), and DAPI nuclear
stain (blue). Representative images (A, 40X magnification) from each group show sparse TUNEL positive cells, and there is no difference in number of TUNEL positive cells per image (B) or in fraction of TUNEL positive cells (C). Data expressed as Mean±SD, N=3/group. Scale bars=20µm.
The raw counts made for both DAPI+ and TUNEL+ cells from the 6Wk heart sections are
seen in Figure D-9. The counts were made from multiple images taken per section at two varying
magnifications, 40X (Figure D-9 A) and 10X (Figure D-9 B) as a method to test variability in
counts at the two magnifications, which it appeared there was not. Sample images used to
optimize the macro for counting are displayed in Figure D-10 A, with associated control macro
optimized counts for both DAPI+ and TUNEL+ cells shown in Figure D-10 B. Again, the macro
was optimized for images taken at 40X or 10X magnification to ensure no differences in relative
cell counts between sections with varying magnification.
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Figure D-8: 3Wk TUNEL/cTnT staining controls TUNEL (red), cTnT (green), and DAPI nuclear stain (blue). (A) Representative 40X magnification images of
caspase treated +ve Control, and TUNEL label added only –ve Control. Obvious differences are seen and images were used to calibrate the macro used for counting TUNEL+ve and DAPI+ve cells (B). Data expressed as Mean. Scale bars=20µm.
Figure D-9: 6Wk TUNEL/cTnT staining raw counts
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Raw TUNEL positive and Dapi positive cell numbers per image, used to make Figure 4-8, at (A) 40X and (B) 10X magnifications. Data expressed as Mean±SD, N=4/group.
Figure D-10: 6Wk TUNEL/cTnT staining controls TUNEL (red), cTnT (green), and DAPI nuclear stain (blue). Clear TUNEL positive cells seen at both 40X
and 10X caspase treated positive control (A), and none seen in the negative control images. Images were used to optimize macro for counting TUNEL and DAPI +ve cell numbers (B). Data expressed as Mean. Scale bars=20µm 40X, 100µm 10X.
