Development of an In Vitro Outer Annulus Fibrosus-

Cartilage Endplate Model and its Response to Dynamic

Mechanical Loading

by

Jasmine E. Chong

A thesis submitted in conformity with the requirements

for the degree of Masters of Applied Science

Institute of Biomaterials and Biomedical Engineering

University of Toronto

© Copyright by Jasmine E. Chong 2019

ii

Development of an In Vitro Outer Annulus Fibrosus-Cartilage Endplate Model

and its Response to Dynamic Mechanical Loading

A thesis submitted in conformity with the requirements for the degree of

Masters of Applied Science

Institute of Biomaterials and Biomedical Engineering

University of Toronto

© Copyright by Jasmine E. Chong 2019

ABSTRACT

Tissue-engineering research to date has neglected to address the annulus fibrosus (AF)-

cartilage endplate (CEP) interface of the intervertebral disc. It was hypothesized that in vitro-

formed outer AF and cartilage tissues will integrate in co-culture to form an interface

containing ECM that resembles the native interface, and that can be mechanically and

compositionally enhanced through dynamic loading. 2-week old in vitro OAF tissues

integrated with 3-day old in vitro cartilage by 1 week of co-culture, and resembled the

distributions of collagen type I, collagen type II, and aggrecan seen in the native interface.

The apparent tensile strength of the in vitro interface increased between 2 and 4 weeks of co-

culture. Dynamic mechanical loading did not enhance the mechanical strength, appearance,

or composition of the interfacial matrix. This study demonstrated the ability to engineer a

biological AF-CEP interface model, and will assist in defining how future tissue-engineered

disc replacements should be developed.

iii

ACKNOWLEDGEMENTS

I would like to thank my supervisors Dr. Rita Kandel and Dr. Paul Santerre for their guidance

and encouragement over the course of my degree. I would also like to thank my committee

members Dr. Craig Simmons and Dr. Marc Grynpas for their input and assistance during the

completion of my project.

It has been a pleasure to work with all members of the Kandel Lab, who have not only provided

me with assistance when in need, but have given me countless memories I will always look back

on with a great big laugh. A special thanks goes to Jian Wang at the Faculty of Dentistry. Never

before have I been so pleasantly surprised when we serendipitously discovered our mutual

obsession with street photography, which never failed to keep the conversation interesting and

enlightening during those long and would-be tedious mechanical testing sessions.

Thanks to my parents, who have constantly supported me in more ways than one: To my mother

for always keeping me very well fed, and to my father whom I could always count on to be my

own personal machine shop technician or to let me borrow tools, electrical tape, and random

pieces of plexiglass during my own personal desperate times of lab-related MacGyver-ing. I

hope that whatever I achieve in the future pays off for you both.

iv

TABLE OF CONTENTS

Abstract ........................................................................................................................................... ii

Acknowledgements ........................................................................................................................ iii

Table of Contents ........................................................................................................................... iv

List of Tables ............................................................................................................................... viii

List of Figures ................................................................................................................................ ix

List of Supplementary Figures ........................................................................................................ x

List of Appendices ......................................................................................................................... xi

List of Abbreviations .................................................................................................................... xii

CHAPTER 1: Introduction ............................................................................................................. 1

1.1 Introduction ........................................................................................................................... 1

1.2 Hypothesis ............................................................................................................................. 5

1.3 Objectives .............................................................................................................................. 5

CHAPTER 2: Literature Review .................................................................................................... 7

2.1 Intervertebral Disc ................................................................................................................. 7

2.1.1 Embryonic Development of the Intervertebral Disc .................................................... 10

2.1.1.1 Extracellular Matrix Components of the Developing Intervertebral Disc ............. 14

2.1.2 Nucleus Pulposus.......................................................................................................... 15

2.1.3 Annulus Fibrosus .......................................................................................................... 17

2.1.4 Cartilage Endplate ........................................................................................................ 18

v

2.2 The Annulus Fibrosus-Cartilage Endplate Interface ........................................................... 19

2.3 Degenerative Disc Disease .................................................................................................. 23

2.3.1 Current Treatment Strategies for Degenerative Disc Disease ...................................... 26

2.3.1.1 Injectable Biomolecules ......................................................................................... 27

2.3.1.2 Cell-Based Therapies ............................................................................................. 28

2.4 Intervertebral Disc Tissue-Engineering .............................................................................. 32

2.4.1 AF Tissue Engineering ................................................................................................. 33

2.4.2 Electrospun Aligned Nanofibrous Polymer Scaffold for Multilamellar Tissue-

Engineered AF ....................................................................................................................... 37

2.4.3 Tissue-Engineered IVD-CEP Integration ..................................................................... 40

2.5 The Role of Mechanical Stimulation on Extracellular Matrix Composition ...................... 42

CHAPTER 3: Development of an In Vitro Outer Annulus Fibrosus-Cartilage Endplate Model

and its Response to Dynamic Mechanical Loading ...................................................................... 53

3.1 Introduction ......................................................................................................................... 53

3.2 Methods ............................................................................................................................... 56

3.2.1 Isolation of OAF cells .................................................................................................. 56

3.2.2 Generation of aligned nanofibrous PU-ADO scaffold constructs ................................ 57

3.2.3 Formation of hyaline cartilage tissue in 3D culture ..................................................... 58

3.2.4 Fluorescent labelling of primary chondrocytes ............................................................ 59

3.2.5 Co-culture of in vitro-formed OAF and cartilage tissues ............................................. 60

vi

3.2.6 Determination of mechanical loading magnitude ......................................................... 62

3.2.7 Dynamic mechanical loading of OAF-CEP constructs ................................................ 65

3.2.8 Pull-apart testing of OAF-CEP constructs ................................................................... 66

3.2.9 Histological characterization of native and in vitro OAF-CEP interface ..................... 69

3.2.10 Immunohistochemical characterization of the OAF-CEP interface model ................ 69

3.2.11 Quantification of DNA, collagen, and proteoglycan contents ................................... 70

3.2.12 Statistical analysis ...................................................................................................... 72

3.3 Results ................................................................................................................................. 73

3.3.1 Co-culture of in vitro OAF on cartilage tissue generated an integrated tissue

interface ................................................................................................................................. 73

3.3.2 Chondrocytes remained localized to the cartilage layer beneath the interface ............ 78

3.3.3 Immunohistochemical characterization of the bioengineered OAF-CEP

interface ................................................................................................................................. 79

3.3.4 Apparent mechanical strength of the in vitro OAF-CEP interface increased

over time ................................................................................................................................ 82

3.3.5 No significant difference in OH-Pro or GAG content between 2 and 4 weeks of

co-culture ............................................................................................................................... 84

3.3.6 A compressive force of 0.4 N resulted <10% radial strain in OAF tissues .................. 86

3.3.7 Dynamic mechanical stimulation did not significantly increase the apparent

tensile strength, collagen content, or proteoglycan content in OAF-CEP interface

constructs ............................................................................................................................... 88

vii

CHAPTER 4: Discussion .............................................................................................................. 92

4.1 Summary ............................................................................................................................. 92

4.2 Discussion and Study Limitations ....................................................................................... 93

4.2.1 Use of in vitro-formed OAF tissue on aligned angle-ply PU-ADO nanofibrous

scaffolds and hyaline-like cartilage are suitable to generate an OAF-CEP model................ 93

4.2.2 Requirements for successful co-culture of in vitro OAF and cartilage tissues ............ 95

4.2.3 Characterization of the OAF-CEP interface ................................................................. 98

4.2.4 Failure testing and apparent mechanical strength of the OAF-CEP interface ............ 102

4.2.5 Dynamic mechanical loading of the in vitro OAF-CEP interface .............................. 108

4.3 Recommendations for Future Work .................................................................................. 113

4.4 Conclusions ....................................................................................................................... 116

CHAPTER 5: References ........................................................................................................... 118

viii

LIST OF TABLES

Table 1: Literature review summary of mechanical loading effects on IVD and

cartilage cells and tissue……………………………………………………………43-48

ix

LIST OF FIGURES

Figure 2.1: Schematic diagrams of the native IVD……………………………………………... ..9

Figure 2.2 Schematic diagrams of IVD development ……………………………………….13-14

Figure 2.3 DIC image of AF fibre bundle insertion into the native CEP……………………….. 19

Figure 2.4 DIC image of AF fibres pulling out VB osteons during tensile testing……………...22

Figure 3.1: Schematics showing cell isolation and tissue culture methods…………………..61-62

Figure 3.2: Determination of compressive load parameters for OAF + agarose plug………..64-65

Figure 3.3: Loading of OAF-CEP constructs in a Mach-1™ mechanical tester………………... 66

Figure 3.4: Pull-apart failure testing and OAF surface area measurement.………………….68-69

Figure 3.5: Schematic diagram of OAF-CEP interface preparation for biochemical analysis…..72

Figure 3.6: Histological appearance of in vitro and native OAF-CEP interface …………….75-76

Figure 3.7: Histological appearance 3-day old, 5-day old, and 7-day old cartilage tissues

interfaced with OAF or cultured alone……………………………………………77-78

Figure 3.8: CFDA-labelled chondrocytes in OAF-CEP constructs……………………………..79

Figure 3.9: Immunostaining of in vitro and native OAF-CEP interface for collagen type I,

collagen type II, and aggrecan…………………………………………………….81-82

Figure 3.10: OAF-CEP apparent stress and histological appearance at failure ……………..83-84

Figure 3.11: Biochemical analysis of DNA, GAG, and OH-Pro contents between 2 and

4-week old interface constructs………………………………………………………85

Figure 3.12: Determination of loading parameters by mechanical testing of OAF

and agarose plug constructs…………………………………………………….....87-88

Figure 3.13: Effect of dynamic loading on mechanical strength, histological

appearance, and ECM composition at the OAF-CEP interface……………………90-91

x

LIST OF SUPPLEMENTARY FIGURES

Figure S1: Appearance of cultured OAF-CEP constructs with and

without physical tubing………………………………………………………..150-151

Figure S2: Method of scaffold strip cutting……………………………………………………152

Figure S3: Fluorescent CFDA labelling of chondrocytes…………………………………….. 153

Figure S4: Set-up diagrams for dynamic loading…………………………………………154-155

Figure S5: Method of cryotape sectioning………………………………………………...156-157

Figure S6: Nanofiber alignment confirmation by SEM………………………………………. 157

Figure S7: Resemblance between in vitro and native OAF by immunostaining……………....158

Figure S8: Histological appearance of variations in construct integration…………….............159

Figure S9: In vitro cartilage grown alone in AF-optimized media versus control

Ham’s F12 conditions……………...……………...………………………………160-161

Figure S10: Appearance of paucicellular cartilage layer by immunostaining…………………161

xi

LIST OF APPENDICES

APPENDIX A: Method Development for In Vitro OAF-CEP Construct Assembly ................. 149

APPENDIX B: Supplementary Figures ...................................................................................... 150

xii

LIST OF ABBREVIATIONS

ADO Anionic dihydroxyl oligomer

AF Annulus Fibrosus

CEP Cartilage endplate

CFDA 5,6-carboxyfluorescein diacetate

CPP Calcium Polyphosphate

DAPI 4’6-diamidino-2-phenylindole

DAPS Disc-like angle ply structure

DDD Degenerative disc disease

DIC Differential interference contrast

DMEM Dulbecco's modification Eagle's medium

DMSO Dimethyl sulfoxide

DNA Deoxyribonucleic acid

DTT Dithiothreitol

ECM Extracellular matrix

EDTA Ethylenediaminetetraacetic acid

FBS Fetal bovine serum

GAG Glycosaminoglycan

GDF-5 Growth differentiation factor-5

H&E Hematoxylin & eosin

HA Hyaluronic acid

HSC Hematopoietic stem cell

hMSC Human mesenchymal stem cell

xiii

IAF Inner annulus fibrosus

IgG Immunoglobulin G

iPSC Induced pluripotent stem cell

ITS Insulin-transferrin-selenium

IVD Intervertebral disc

MMP Matrix metalloproteinase

MRI Magnetic resonance imaging

MSC Mesenchymal stem cell

NP Nucleus pulposus

OAF Outer annulus fibrosus

OH-Pro Hydroxyproline

PBS Phosphate buffered saline

PCL Polycaprolactone

PLA Polylactic acid

PLGA Poly(lactic-co-glycolic acid)

PRP Platelet-rich-plasma

PU Polycarbonate urethane

PU-ADO Polycarbonate urethane-anionic dihydroxyl oligomer

SEM Scanning electron microscopy

SLRP Small leucine-rich proteoglycan

TBS Tris buffered saline

TE Tissue-engineered

VB Vertebral body

1

CHAPTER 1: INTRODUCTION

1.1 Introduction

The intervertebral disc (IVD) is a fibrocartilaginous tissue that permits flexibility and load

transmission between adjacent vertebral bodies in the spine. It consists of a central gel-like

nucleus pulposus (NP) and a surrounding annulus fibrosus (AF). Relative to the rest of the disc,

the NP contains the highest content of collagen type II and aggrecan, a proteoglycan that

provides the tissue with compressive-resistant properties. The content of both of these

extracellular matrix (ECM) macromolecules progressively decreases, while collagen type I

increases, towards the outermost part of the AF. The AF is a complex concentric, multi-

lamellated, angle-ply fibrous structure; collagen fibres of successive lamellae are oriented at

opposing ~30° angles to each other. It is further subdivided into inner (IAF) and outer (OAF)

regions. The IAF contains collagen type I, collagen type II, and rounded cells. The OAF is more

organized than the IAF and consists mostly of highly aligned collagen type I fibres and elongated

fibroblastic cells. Cartilage endplates (CEP) composed of hyaline cartilage are located cranially

and caudally to each disc and provide anchorage of the soft IVD into the hard vertebral bone

(VB). Within the CEP, a layer of calcified cartilage lies immediately adjacent to the VB which

acts to further strengthen the disc-cartilage-bone anchorage system1. The CEP and the periphery

of the OAF contain small blood vessels, which likely provide routes of nutrient transport to the

IVD via diffusion2-4. Despite peripheral contacts with these small capillaries, overall the IVD is

an avascular tissue that is housed in a low-oxygen environment. At approximately 6000

cells/mm3, the human IVD has an even lower cell density than articular cartilage, which is in

itself known to be a tissue of low cellularity2. Poor vascularity in conjunction with low cellularity

2

severely limits the IVD’s capacity for self-repair, and degenerative processes once initiated are

irreversible.

Degenerative disc disease (DDD) is associated with lower back pain5,6, and at advanced stages

can lead to decreased disc height, AF tearing, and disc herniation7. Conservative treatment of

back pain symptoms with physical therapy, analgesics or muscle relaxants does not target the

underlying tissue pathology of DDD. Surgical approaches for treatment of advanced DDD

include discectomy, spinal fusion, and prosthetic disc replacement, however these do not

guarantee long-term pain relief. These procedures have been reported to have low overall success

rates and may lead to complications such as adjacent disc degeneration8,9.

Cell-based strategies for DDD that have been implemented at the clinical level include

intradiscal placement of autologous bone marrow-derived mesenchymal stem cells (MSCs)10,11,

allogeneic bone marrow-derived MSCs12, passaged allogeneic juvenile chondrocytes (NuQu®,

ISTO Technologies Inc.)13, and autologous cultured disc cells (chondrotransplant® DISC,

co.don)14. Although the results from these studies suggested that such cell-based treatments were

able to reduce pain and disability scores over time, the long-term outcomes of these studies past

2 years, including adverse effects, are unknown. It is unclear whether the injected/implanted cells

exerted their effects by targeting underlying degenerative processes. It is also not known whether

the implanted cells acquired and/or sustained a disc cell-like phenotype, nor whether they

promoted the appropriate regional ECM production; these limitations are especially true for the

intradiscal implantation of marrow-derived MSCs and passaged juvenile chondrocytes. Although

in vitro studies have induced MSCs to differentiate towards AF and NP cell-like phenotypes15-17,

3

acquisition and maintenance of the correct phenotypes are difficult to achieve. Hyaline cartilage-

producing chondrocytes are phenotypically distinct from IVD cells18,19, and intradiscal

implantation of these cells might have unknown long-term adverse effects. While autologous

disc cells might seem like a solution to circumvent the phenotypic challenges presented by MSCs

and chondrocytes, intradiscal placement of these cells by injection is not expected to restore the

exact structural organization of the native disc. Therefore, a bioengineered implant solution

composed of IVD cells seeded within a biomimetic IVD scaffold structure might be able to

replicate the form and function of the native IVD in such a way that prosthetic replacements or

intradiscal cellular injections cannot achieve.

A major challenge of IVD tissue engineering is accurate reproduction of the complex AF

structure and organization. Furthermore, studies to date that have generated whole TE-IVD

implant models have simplified AF construction by overlooking the distinct inner and outer

regions. AF cell-seeded collagen hydrogels have been used for AF generation, but they have not

shown the ability to mimic every individual structural feature of the AF within a single IVD

construct. Photochemical cross-linking of successive collagen layers around an NP-like hydrogel

core was achieved in order to mimic the multi-lamellar feature of the AF, but was unable to

generate aligned or angle-ply features20. Collagen gels were also contracted around an NP-like

hydrogel core to produce circumferential alignment of collagen fibres, however an angle-ply

structure was not achievable by this method21. Other IVD models have made use of AF cell-

seeded porous lamellar silk22 and micropatterned bacterial cellulose scaffolds23 in IVD implant

construction, which were able to address the circumferential, multilamellar, and aligned angle-

ply structural components. However, production of biological materials such as collagen, silk, or

4

bacterial cellulose may not be as efficient or fine-tunable as the production of a synthetic

polymer, which may be considered to be a better alternative for reproducible large-scale AF

scaffold generation. Current TE whole IVD implant models have utilized electrospun

nanofibrous aligned polymer scaffolds to mimic all the structural components of the AF24-28.

Strips of aligned nanofibrous scaffold can be fashioned into concentric multilamellar angle-ply

constructs. In comparison to scaffolds containing randomly oriented nanofibers, directional

alignment of nanofibrous polymer scaffolds promotes alignment and morphological elongation

of seeded AF cells, which subsequently enhances the production of AF-specific ECM

components29. Aligned nanofibrous scaffolds made from electrospun polycarbonate urethane

with added anionic dihydroxyl-oligomer (PU-ADO) have been shown to support the

maintenance of separately cultured IAF and OAF cell phenotypes, as well as the distinct region-

specific ECM accumulation of each respective AF region30,31.

Tissue engineering research has made much progress in the generation of a structurally accurate

TE IVD, however ensuring appropriate and stable integration of TE implants with surrounding

tissue in vivo remains difficult. In vitro culture of bovine AF cell-seeded nanofibrous angle-ply

structures showed accumulation of collagen and GAG-containing ECM, however upon

implantation into rats these constructs extruded from the disc space28. Although an acellular

porous polymer that was subsequently incorporated into the model acted as a bony endplate and

improved integration upon implantation of the whole TE IVD28, inclusion of a CEP layer to more

accurately model the anchorage that normally occurs between the native IVD and VB was

neglected. Achieving integration of engineered IVD and cartilage tissues prior to implantation is

a necessary next step towards the development of a more complete disc replacement model.

5

Determination of applicable treatments to enhance the integration process in vitro is also of

interest in order to maximize the potential for implant success. Dynamic mechanical loading is

one such treatment that has been shown to have anabolic effects on ECM by stimulating

macromolecule gene and protein expression, and potentially by reducing ECM degradation32-53.

While the anabolic effects of dynamic compressive and tensile loading have been shown in

cartilage and IVD tissues and cells, the effect on the IVD-CEP interface is unknown. The current

thesis investigates the development of a bioengineered nanofibrous scaffold-based AF-CEP

interface model, and the potential for dynamic mechanical stimulation to enhance interfacial

tissue formation and strength. OAF cells will be used in the generation of this model as a primary

step towards producing an AF-CEP model that has differentiated inner and outer regions.

1.2 Hypothesis

In vitro OAF tissue grown on aligned nanofibrous multi-lamellated PU-ADO scaffolds will

integrate in co-culture with in vitro cartilage tissue grown on a porous Teflon membrane, to 1)

form an interface containing ECM that resembles the native OAF-CEP junction in its cellular

composition and immunolocalization of collagen type I, collagen type II and aggrecan, and 2)

that can be mechanically and compositionally enhanced through dynamic loading.

1.3 Objectives

1) Generate an in vitro OAF-CEP interface model using primary bovine OAF cells, primary

bovine articular chondrocytes, and an electrospun nanofibrous PU-ADO scaffold

6

2) Characterize the OAF-CEP interface tissues in vitro and determine the effect of culture

time on their development and integration through histological examination and

mechanical axial tensile testing

3) Determine the effect of dynamic compressive loading on in vitro OAF-CEP interface

ECM accumulation and mechanical strength

7

CHAPTER 2: LITERATURE REVIEW

2.1 Intervertebral Disc

The IVD is located between adjacent vertebral bodies and functions to permit flexibility and load

transmission within the spine. It is composed of a central collagen type II and proteoglycan-rich

nucleus pulposus (NP) that is contained by a surrounding annulus fibrosus (AF). Hyaline

cartilage endplates (CEP) located superiorly and inferiorly attach each disc to adjacent vertebral

bodies (Figure 2.1). The IVD is an avascular tissue located in a low oxygen environment,

however it obtains its nutrients via diffusion from small blood vessels that can be found around

the outermost periphery of the AF2. Blood vessels also extend from the subchondral bone,

through the bony vertebral endplates and into the CEP, however there is not clear evidence that

the latter extend from the CEP into the disc4,54. Human cranial vertebral endplates were recently

found to contain a significantly higher number of marrow openings in comparison to caudal

endplates, suggesting that the IVD relies more heavily on blood supply originating from the

adjacent cranial vertebra versus the caudal vertebra55. The average water content in the human

disc is approximately 75%, with ~10% higher content in the NP when compared to the AF56,57.

Relative to the NP, the AF has approximately double the collagen as determined by

hydroxyproline content per dry weight, and half the proteoglycan content as estimated by

spectrophotometric assay via dimethylmethylene reactivity57. The ratio of proteoglycan to

collagen in the NP is far greater when compared to the AF and CEP, which were found to be

27:1, 1.6:1, and 2:1 respectively in healthy human adult tissues by glycosaminoglycan (GAG) to

hydroxyproline dry weight18.

8

The main collagens that comprise the IVD are the fibrillar collagen types I and II, which are

characteristic structural proteins of tensile-resistant fibrous tissues and compressive-resistant

cartilage tissues, respectively58. Collagen type I fibrils pack parallel to each other in an aligned

fashion giving these fibres tensile strength, whereas collagen type II fibrils are less organized and

interact with ECM proteoglycans to resist compressive loads59,60. Collagen type I is highly

abundant in the OAF and decreases centrally as the NP is approached, while collagen type II is

most highly abundant centrally in the NP and decreases laterally towards the OAF58. Collagen

type III has been found to be mainly pericellular within the disc, although its role is unclear61.

Collagen type VI is found in the translamellar bridges of the AF and may have a role in collagen

bundle lubrication62. Collagen type X, which is a marker of hypertrophic chondrocytes involved

in mineralization, can be found in the native disc-vertebra interface63.

Proteoglycan content of the IVD is highest within the nucleus pulposus and lowest in the OAF,

with percent water content generally following the same trend56. The main proteoglycan found in

the IVD and CEP is aggrecan, which is similar to collagen type II in its regional distribution and

in its role to provide compressive resistance64. It exists within the extracellular matrix (ECM) in

the form of aggregates, whereby large numbers of aggrecan molecules are bound to hyaluronic

acid. Its osmotic properties allow it to swell and resist compression when hydrated. Lubricin, a

lubricating proteoglycan that is known to be produced by superficial zone articular

chondrocytes65, has also been identified in human IVD. While lubricin staining was found to be

distributed diffusely throughout the human disc, the NP and IAF stained more intensely

compared to the OAF66. In contrast, lubricin in mature goat IVD was not found in the NP nor the

IAF but was rather localized to the OAF lamellae and interlamellar junctions, and to the

9

intracellular regions of AF and NP cells in primary monolayer culture67. The contrasting results

between the two studies may be attributed in part to species differences and/or effects of

degeneration in the cadaveric adult human specimens. Despite these differences, both studies

highlighted that the presence of lubricin might have a functional role in IVD biomechanics.

Lubricin is known to reduce friction at the articulating surfaces of synovial joints68. IVDs from

lubricin knockout mice were found to have greater rotational stiffness when compared to wild

type controls, indicating that lubricin might have a similar role in the IVD to facilitate proper

mechanical function69.

Figure 2.1. Schematic diagrams of the native IVD. (A) The IVD consists of the nucleus

pulposus (NP) and the annulus fibrosus (AF), and is attached to the vertebral bone (VB)

through hyaline cartilage endplates (CEP). A layer of calcified cartilage (CC) lies between

the hyaline CEP and VB. (B) The multilamellar AF is divided into inner (IAF) and outer

(OAF) regions. Collagen fibres between successive lamellae are oriented at opposing 30°

angles.

10

2.1.1 Embryonic Development of the Intervertebral Disc

In embryonic development of the vertebral column, the NP originates from the rod-like

notochord and its surrounding notochordal sheath, while the AF, CEP, and vertebral bodies arise

from sclerotomal cell condensations of the mesoderm7. Prior to the formation of cell

condensations, sclerotome cells are round in shape70. By approximately day 30 of human

embryogenesis, sclerotome cells have aggregated around the notochord to form a metameric

pattern of loose and dense condensations that will later give rise to the vertebral bodies and the

AF, respectively (Figures 2.2A-D)7,70. The densely packed sclerotome cells elongate to adopt a

fibroblastic morphology and align into sheets of oppositely angled cells (Figure 2.2E)70,71.

Concentric alignment of these cells precedes the deposition of ECM and gives rise to the

lamellar structure seen in the fully developed AF71. This highly organized structural architecture

is established before the vertebral column experiences any external mechanical stress caused by

movement70. The notochord regresses and eventually disappears at the level of the developing

vertebral bodies, while it bulges within the intervertebral tissue (E16 in the rat)70,71. It has been

suggested that mechanical force from the developing vertebral bodies puts pressure on the

intervertebral notochord and causes it to bulge70. This notochordal bulging might promote the

differentiation of the fibrocartilaginous AF tissue into its inner and outer regions, as these two

developmental events coincide with each other71. At E15 in the rat there is very little ECM, the

notochord is a continuous rod-like structure, and differences in morphology between IAF and

OAF are discernable72. After E16, AF differentiation is marked by the orientation of fibroblastic

AF cells into angle-ply lamellae and the appearance of collagen types I and II in the OAF and

IAF respectively, which appear as histologically distinct tissues58,73. Hayes and colleagues

11

showed the presence of cell-cell contacts by electron microscopy and positive immunostaining

for the adherens junction components cadherins and vinculin, in cells of the developing rat OAF

at E1572. Positive immunostaining for integrin α1 and β1 subunits as well as vinculin indicated

that cell-ECM contacts were also present in the form of focal adhesions in the developing OAF

at this stage. Actin was diffuse throughout at E15, however by the time of notochordal bulging at

E16, actin stress fibres were prominent and longitudinally oriented within the aligned fibroblastic

cells of the developing OAF. Pericellular fibronectin was diffuse throughout the disc until E16,

at which point it organized longitudinally around OAF cells and appeared to mirror the

intracellular organization of actin stress fibres. The authors suggested from the collective

evidence that cell-cell and cell-ECM contacts are required to establish aligned AF cell

orientation and actin stress fibres, which in turn mediate the longitudinal organization of

pericellular fibronectin that ultimately guides the assembly of aligned collagen deposition72.

SEM images from a later study showed that the orientation of extracellular collagen fibrils

between AF cells did indeed follow the path of longitudinal intracellular stress fibres73.

Cadherins in the developing OAF was no longer detectable by E17 and stress fibres were absent

by birth, suggesting that the mediators of cell-cell interactions and cellular orientation were no

longer required by the time of ECM deposition. Vinculin and integrin subunit expression

persisted throughout development, suggesting that cell-matrix interactions were necessary for the

continued organized deposition of ECM after the establishment of cellular orientation72.

The vertebrae and CEP embryonic cartilage anlagen develop from the loose sclerotomal cell

condensations. At the region of the developing CEP, cells are aligned horizontally, parallel to

adjacent vertebral bodies70. In humans, AF collagen fibres form interconnections with the CEP

12

by the seventh embryonic month, and CEP collagen fibres are oriented parallel to adjacent

vertebral bodies74. The fibrils of the OAF are joined with the uncalcified region of the peripheral

CEP cartilage that will later ossify to become the vertebral rim74. CEP thickness is greatest in the

immature spine and decreases with age as the outermost region becomes calcified75.

13

Figure 2.2. Schematic diagrams of IVD development. (A) Loose (future VB, CEP) and

dense (future AF) sclerotomal cell condensations form around the rod-shaped notochord

(future NP). (B) The notochord bulges within the developing AF and constricts within the

14

developing VB cartilages, which also coincides with differentiation of the AF. (C) The

notochord regresses in the region of the developing VB cartilages. (D) The fully developed

IVD and ossified VB. (E) Elongated cells of the developing OAF align into sheets of

oppositely angled cells (tangential section, representative of rat E1671).

2.1.1.1 Extracellular Matrix Components of the Developing Intervertebral Disc

As mentioned in the previous sections, AF differentiation is marked by the appearance of

collagen type I in the OAF and collagen type II in the IAF at the time of notochordal bulging58.

Collagen type III, which plays a role in collagen type I fibrillogenesis23, was also found to be

present throughout the developing rat disc prior to notochordal bulging; however, after E16 it

remained localized with collagen type I in the OAF while it was gradually lost from the IAF58,71.

Collagen type IV, which is known to normally occur in basement membranes76, was present in

the rat notochordal sheath and OAF at E1658. By E20, it was lost from the notochordal sheath

and remained in the OAF and pericellularly in the IAF. The function of collagen type IV in the

developing IVD is unclear.

Aggrecan, the major proteoglycan of the IVD, was present in the cartilages of the developing rat

vertebral bodies, and the IAF region by E17. By birth, aggrecan remained colocalized with

collagen type II in the IAF58. Several small leucine-rich proteoglycans (SLRPs) have been

reported to be found in the developing and mature IVD such as decorin, biglycan, fibromodulin,

and versican58,73. Decorin and biglycan, which are involved in collagen fibrillogenesis, were

found by immunostaining in developing rat NP, vertebral body cartilages and IAF by E19, and at

a relatively weaker intensity in the OAF by E2173. Increased biglycan was also noticeable in the

15

IAF and at the AF-vertebra interface by E21. Both decorin and biglycan were associated with

CEP cells at 4 months after birth73. Immunoblotting showed an increased presence of decorin in

the AF of 2 and 10-year old sheep compared to undetectable amounts in fetal samples77,

suggesting that its content further increases with maturity. Fibromodulin is the most predominant

SLRP found in the developing AF and likely has roles in fibrillogenesis and lamellar

organization73. Fibromodulin was localized to the rat IAF and AF-vertebra interface at E17. Its

distribution extended into the vertebral body cartilages by E19 and to the OAF by E21, and

persisted throughout the AF after birth73. Versican was reported to be present in ovine AF

interlamellar junctions and translamellar bridges, and might have a possible role in collagen

bundle lubrication62,77. Immunostaining showed that native young bovine IAF contained more

versican than OAF, which was reflected by the finding that primary cultured bovine IAF cells

accumulated more versican over time compared to OAF cells30. However, Hayes et al. described

the distribution of versican in the developing rat IVD which was at first predominant in the IAF

and posterior OAF at E17, but absent from the IAF by birth58. As such, it was reported here that

versican colocalized with collagen type I and was only present in the OAF. The inconsistent

findings regarding versican localization in inner versus outer AF may be in part due to

differences between species.

2.1.2 Nucleus Pulposus

The NP is the central region of the IVD and is mainly subjected to compressive loading. It is a

hydrated gel-like tissue that is mainly composed of collagen type II and aggrecan30,78. The

collagen fibril network organization in the NP is irregular79. Tensile failure testing suggested that

the collagen fibres of the NP are longer and more highly convoluted in the central region

16

compared to the peripheral region80. Elastin in the NP forms an organized network of fibres that

run both radially towards the AF and vertically to anchor into the CEP81. At the interface

between the NP and CEP, elastic fibres in the NP are parallel to each other and run vertically

towards the endplates81.

The cell population in the NP is mixed, consisting of small chondrocyte-like NP cells and

clusters of large notochordal cells that have vacuoles and dense cytoplasmic actin82,83.

Differential gene expression has been shown between these two cell populations84. Although NP

cells are often considered to be “chondrocyte-like”, it has been suggested that these cells are

differentiated notochordal cells and therefore do not originate from the same mesenchymal

lineage as chondrocytes. Porcine NP notochordal cells subjected to hydrostatic pressure

transitioned into small NP cells, and time-lapse imaging of rabbit NP notochordal cells also

showed differentiation into smaller chondrocyte-like cells85,86. Aging may be involved in the

progressive loss of notochordal cells in certain species, including humans. Loss of notochordal

cells with age has been shown in some species such as mongrel dogs and sheep, while they

persist with maturity in other species such as rat and rabbit82. Older, more degenerated NP

tissues from non-chondrodystrophoid dogs were shown to have a loss of vacuole-containing,

actin-rich notochordal cells compared to NP from younger, healthier discs83. Human notochordal

cells were found to be present in clusters in early fetal development (embryonic week 8),

however were undetectable in adult human specimens87. Degenerative processes are also likely

to be involved in notochordal cell loss. An in vivo mouse model of degeneration by needle

puncture showed a decrease of NP notochordal cells and an increase of NP chondrocyte-like

cells within 2 weeks of injury compared to unpunctured controls88.

17

2.1.3 Annulus Fibrosus

The AF experiences axial compression; longitudinal, radial, and circumferential tension; shear

stress between adjacent lamellae and at the interfacing CEP; and torsion. It is a heterogeneous,

multi-lamellar structure comprised of aligned collagen fibres that are alternately angled at

approximately 30° with respect to the vertical axis89. It is further distinctly subdivided into the

outer (OAF) and inner (IAF) regions. The OAF is composed of aligned, highly organized type I

collagen fibre bundles and elongated fibroblast-like cells, whereas the IAF is less organized and

contains proteoglycans, type I and type II collagen, and rounded cells7,90. OAF cells have a

polarized, spindle-shaped morphology and express cell polarity proteins Par3, claudin-1, and

claudin-11 to a higher degree as compared to IAF or NP cells91. Directed ECM secretion by

polarized OAF cells may in part be responsible for the organized aligned structure of OAF

lamellae91.

The spaces between adjacent lamellae, or interlamellar junctions, are proteoglycan-rich and

contain organized networks of elastic fibres as shown by both orcein staining and elastin

immunostaining62,92. The interlamellar elastic fibre network is loose compared to the densely

packed collagen fibres of the lamellae; tension and shear testing demonstrated that failure

occurred at the interlamellar matrix and not within the lamellae93. The interlamellar junctions

form continuous connections with proteoglycan-rich translamellar cross-bridges, which are

predominantly found in the mid-AF and appear as transverse discontinuities within the

lamellae62. These interconnecting cross-bridges become more numerous with increasing age,

may function to prevent delamination, and have been suggested to be the remnants of regressed

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blood vessels62,94,95. In addition to containing collagen type I and elastin, the interlamellar

junctions and translamellar cross-bridges contain localized collagen type VI, aggrecan, and

versican62. While the function of collagen type VI in these regions is unknown, it has been

suggested that it is a marker of fibrocartilage differentiation96. Aggrecan and collagen VI have

also been found between collagen fibre bundles in the fibrocartilaginous region of the bovine

deep digital flexor tendon which experiences both tension and compression during locomotion,

suggesting that these molecules might have potential roles in resisting both compressive and

tensile forces within the IVD96,97. It has also been suggested that aggrecan and versican within

the interlamellar matrix may serve to lubricate collagen fibre bundles and prevent damage to

them during loading62.

2.1.4 Cartilage Endplate

A hyaline cartilage endplate (CEP) lies on either side of the IVD (cranially and caudally),

adjacent to the vertebral bone. Collagen fibres are aligned horizontally with respect to the

adjacent disc and vertebral body79. This layer, which serves as the interface between the disc and

the bone, is reported to have a thickness of approximately 0.6-0.8 mm in adult humans and

resembles articular cartilage in that it contains chondrocytes within an ECM rich in collagen type

II and proteoglycans57,98. An uncalcified cartilaginous region is adjacent to the disc, and a

mineralized region is adjacent to the vertebral bone63. The transition from uncalcified to calcified

cartilage within the CEP occurs sharply in a step-wise fashion rather than in a progressive

gradient99. Collagen fibres within the CEP are oriented horizontally57. Collagen fibres of the AF

insert vertically through the CEP (Figure 2.3), into the calcified region and into the vertebral

bone, forming the AF-CEP interface that is crucial for the distribution of mechanical

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load63,100,101. Vertical insertion of AF fibres through the CEP confers tensile-to-shear stress

transfer, whereby the direct application of tensile stress (from spinal extension/flexion/lateral

bending) is transferred into shear stress along the length of the AF fibre which is embedded

within the cartilaginous matrix100,102. It has also been proposed that splitting of AF fibre bundles

into sub-bundles within the CEP matrix further enhances anchorage and resistance to tensile

stress by increasing the AF fibre-CEP matrix contact area100,101.

Figure 2.3. Vertical insertion of AF fibres (A) into the CEP at the tidemark (TM) and into

the vertebral bone (VEP) at the cement line (CL). From Sapiee et al. Spine J. (2018).

Differential interference contrast microscopy. Reused with the permission of Elsevier,

License Number 4504970208529.

2.2 The Annulus Fibrosus-Cartilage Endplate Interface

The number of structural studies on the native AF-CEP interface is limited, but has been on the

rise in recent years since 20121,63,99-103. Nosikova et al. found that AF collagen fibres pass

through the CEP and insert into the calcified cartilage layer in both immature and mature bovine

20

discs. While fibres from both the IAF and OAF insert through the hyaline CEP and into the layer

of calcified cartilage, the outermost OAF fibres also curve laterally towards the outer aspect of

the vertebral bone to merge with the cartilage63. The AF-CEP interface contains AF cells that are

aligned with the inserting collagen fibres, as well as chondrocytes that are clustered within a

proteoglycan-rich ECM63. AF collagen fibre bundles embedded within the CEP exhibit a wave-

like crimped morphology, a feature which allows for flexibility in the AF when it undergoes

tension (i.e. affords higher mechanical compliance, as crimps will uncoil under tension to

accommodate deformations)100,101. Upon insertion into the CEP, this characteristic crimp is

absent. A structural study by Paietta et al. of human disc-vertebra samples aged 56 days to 84

years found that while the hyaline CEP is present throughout maturity in the IAF region, the

adult OAF region only contains calcified cartilage and subchondral bone, indicating that the

hyaline CEP is remodeled and mineralized with age. The layer of calcified cartilage was uniform

in the infant and thicker in the OAF and vertebral rim regions with increasing maturity99. Elastic

fibres have also been reported to anchor the AF into the CEP104.

There is conflicting evidence in the literature regarding the degree of AF fibre insertion past the

CEP, however recent studies appear to confirm that insertion indeed extends to the vertebral

bone1,100-103. Rodrigues et al. stated in 2012 that fibre bundles from the mid-AF region in ovine

discs branched into sub-bundles and inserted into the CEP, with no further insertion into the

vertebral endplate100. Berg-Johansen et al. concluded from tensile testing, histology, and

scanning electron microscopy (SEM) imaging of adult human cadaveric thoracic spine

specimens, that AF fibres did not insert into the subchondral bone, that the CEP-vertebral bone

interface lacked structural interconnectivity and that it was therefore weakly integrated103.

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Despite these findings, compelling evidence argues against a lack of integration at the interface.

After initially reporting that AF fibres insert through the CEP only, Rodrigues et al. found in

2015 that mature ovine AF sub-bundles did indeed insert through the CEP, into the calcified

cartilage and, to a limited extent into the vertebral endplate bone1. Sapiee et al. recently proposed

the presence of robust anchorage of AF fibres through the CEP and into the vertebral bone in

ovine lumbar specimens, and furthermore that this anchorage is independent of CEP

mineralization102. In this latter study, tensile testing and SEM imaging were performed to show

that AF and CEP fibrils interconnected with the vertebral endplate, where they inserted between

osteons and amongst the bony fibrils. After selective demineralization of endplate regions,

tensile testing of vertebral segments showed that some osteons of the vertebral endplate

remained attached to the pulled-out AF sub-bundles. The results indicated that AF fibres were

structurally integrated with the vertebral bone independent of the CEP mineralized layer. The

authors suggested that four modes of AF fibril insertion are generally present in their integration

with the bony endplate: Perpendicular insertion between bony fibrils; oblique insertion between

bony fibrils; abrupt ending at the bony layer without insertion; and intermingling with bony

fibrils in an irregular fibrillar network102. Adding to the evidence for structural integration shown

in sheep, a human study by Junhui et al. similarly found that AF fibre sub-bundles extended

through the CEP and into the vertebral endplate101. The conflicting findings between the two

human studies may be attributed to the use of older cadaveric human specimens (49-65 years) by

Berg-Johansen et al. as compared to the younger specimens (22-57 years) studied by Junhui et

al. All human specimens examined by Berg-Johansen et al. and were in fact graded above 1, at a

range between 2-4 (average 2.6) according to the Thompson Grade scale of 1 (“healthy”) to 5

(“severely degenerated”)103. In addition, Junhui et al. produced numerous high magnification

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histological images compared to the limited low-magnification images provided by Berg-

Johansen et al. As such, the evidence supporting structural integration at the human AF-CEP-

bone interfaces was arguably superior and more thorough than the evidence for a lack of

integration. High magnification SEM images in the ovine study by Sapiee et al. also provided

compelling proof to support not only that the AF integrates with the CEP, but that its insertion

and interconnectivity with the vertebral bone play important roles in the anchorage system

(Figure 2.4).

Figure 2.4. AF fibres can be seen pulling out bony osteons of the vertebral endplate (VEP)

with them during tensile testing. From Sapiee et al. Spine J. (2018). Differential

interference contrast microscopy. Reused with the permission of Elsevier, License Number

4504970208529.

23

In addition to being the site where attachment and mechanical load transfer occur between the

disc and vertebral bone100,101, the CEP may also directly modulate IVD maintenance through the

inhibition of matrix-degrading enzymes and inflammatory cytokines by secreting modulatory

molecules105,106. Compared to in vitro-formed NP tissue grown alone, NP tissue grown in the

presence of native CEP or in vitro-formed cartilage showed increased proteoglycan content and

decreased protein expression of the matrix metalloproteinase MMP-13, likely due to decreased

gene and protein expression of the proinflammatory cytokine tumor necrosis factor-α105. Trauma

to the CEP may also promote IVD degeneration by inducing both necrotic and apoptotic IVD

cell death, and by increasing collagenase gene expression106. Although the mechanisms of any

direct signalling effects between the CEP and IVD are unknown, it is critical to consider their

integration in order to move towards developing a more complete IVD implant model given the

structural, mechanical, and biological relationships that are known to exist between them.

2.3 Degenerative Disc Disease

Degenerative disc disease (DDD) is associated with lower back pain5,6 and can be attributed to a

combination of different factors such as genetic predisposition, mechanical overloading,

immobilization, and aging107-109. It has been suggested that males are more susceptible to IVD

degeneration earlier and to a higher degree compared to females108. The average cell density of

the human IVD has been reported to be approximately 6000 cells/mm3, approximately 2.5 times

lower than that of articular cartilage which in itself is known to be a tissue of relatively low

cellularity2. Decreased cellularity was found in the AF and NP of old-aged (≥80 years) compared

to middle-aged (35-50 years) human discs55, adding support to the idea that decreases in cell

number with age further reduces the disc’s ability to maintain synthetic function, thus leading to

24

degradative changes in the matrix. In addition, the tissue is poorly vascularized and the dense

matrix likely provides a barrier to cell entry and migration. The capacity for IVD self-repair is

therefore low and the process of degeneration once initiated is irreversible. Normal matrix

remodelling, whereby homeostatic matrix macromolecule degradation and synthesis during

growth and development occur, is impaired in DDD. An increase in denatured collagen type II

and a decrease in procollagen type II was found in Thompson grade 4 degenerated human discs

when compared to grade 1-3 discs, indicating that more advanced degeneration is in part due to

an imbalance in matrix turnover that favours collagen type II degradation over synthesis110. In

addition, increased fibrotic matrix synthesis was evidenced by the finding that procollagen type I

levels, indicative of collagen type I synthesis, were significantly higher in grade 4 discs

compared to grade 2 or 3 discs110. Higher degenerative grade was also associated with decreased

GAG content and a correlated decrease in water content when compared to lower degenerative

grades110. Thinning of collagen fibrils in human AF and NP has been associated with increasing

age and may also be a result of degenerative processes in the IVD55.

Decreased protein content and a loss of uniform distribution of aggrecan, the major proteoglycan

of the IVD, is seen in the AF of degenerated IVDs111. This is mirrored by a decrease in aggrecan

gene expression, as well as downregulation of aggrecanase genes that may be important in the

regulation of normal aggrecan turnover111. Interestingly, human IVD aggrecan extracts inhibited

human neuroblastoma cell line neurite growth and repelled chick dorsal root ganglia explant

growth in a concentration-dependent manner, implying that degenerative aggrecan loss may

partially contribute to increased innervation of the disc resulting in pain112. Decreased aggrecan

content also prevents the tissue from maintaining hydration and may contribute to enhanced AF

25

tearing that is seen with degeneration111. AF tearing may initiate and/or accelerate further

degeneration113. Using magnetic resonance imaging (MRI) and the Pfirrmann MRI grading scale

for disc degeneration114, it was shown in human patients that the presence of annular tears was

associated with higher disc degeneration grades and significantly increased disc degeneration

grades at a 2-year follow-up compared to discs without annular tears113. Advanced degeneration

is marked by an overall reduction in structural IVD integrity in the form of reduced disc height,

annular tears, bulging, and herniation7. Other degenerative changes to the disc and its

environment include SLRP fragmentation, increased apoptotic cell death, increased matrix

degrading enzymes, and an increase in inflammatory cytokines115-118.

Mechanical overloading has been implicated in the progression of DDD. Mechanical overloading

of goat IVDs by axial compression led to increased collagen disorganization and a reduced

proteoglycan to collagen ratio when compared to unloaded control IVDs119. Increasing time and

magnitude of static compressive loading has also been linked to apoptotic cell death120,121. While

unloaded control and sham IVDs showed minimal apoptotic cell death, the relatively lowest

mechanical stress levels induced cell death in the IAF, and the relatively higher stress levels and

longer durations led to more extensive cell death in the CEP and NP121. The number of apoptotic

CEP chondrocytes in organ cultured mouse IVDs increased as the degree of static compressive

loading increased from 0 to 1.0 MPa120.

Pathologic changes of the CEP in relation to aging and disc degeneration include reduced

thickness, point defects, and reduced permeability leading to decreased nutrient transport122-124.

Reduced disc nutrition and subsequent degeneration may also result from the obstruction of

26

vascular channels that pass from the vertebral bone marrow to the CEP. Decreased marrow

contact channel number and size in vertebral endplates was linked to increased Thompson

degenerative grade125. Endplates having fewer channel openings per unit surface area also had

decreased proteoglycan content in the NP region when compared to endplates having more

openings. Likely due to calcification, the loss of channel openings presumably leads to decreased

capillary blood supply to the endplates and consequently decreased disc nutrition and waste

removal by diffusion, which may then promote degenerative changes to the disc125. Ariga et al.

found structural degeneration and an increase in apoptotic cells in the CEP of naturally aging and

surgically-induced spondylotic rats, and suggested that pathologic changes in the CEP precede

degenerative changes in the disc126.

2.3.1 Current Treatment Strategies for Degenerative Disc Disease

Conservative treatments such as analgesics, muscle relaxants, and physical therapy are used to

treat the symptoms of back pain but not the underlying tissue pathology127,128. Current surgical

treatments for degenerated and herniated discs include discectomy, spinal fusion, and prosthetic

disc replacement. Removal of herniated disc tissue by discectomy reduces disc height and does

not guarantee long-term pain relief, as up to 25% of patients may experience recurrent back pain

and disability 2 years postoperatively14,129. Prosthetic disc replacement is a newer procedure that

maintains the motion segment and is theoretically meant to overcome some of the limitations of

spinal fusion9. However, complications arising from prosthetic disc replacement include adjacent

segment degeneration, prosthesis migration, and wear of the implant material130. One meta-

analysis reported total disc replacement as having a higher success rate, superior clinical

outcomes, and a lower incidence of adjacent segment degeneration when compared to

27

discectomy and fusion131. Other reports have shown contradictory findings and suggested that

both procedures have low overall success rates, no significant clinical outcome differences, and

similar levels of adjacent disc degeneration8,9,132.

2.3.1.1 Injectable Biomolecules

Clinical trials have investigated the efficacy of injectable biomolecules for DDD therapy such as

growth differentiation factor-5 (GDF-5) and platelet-rich-plasma (PRP). In vivo animal studies in

rabbits133 and mice134 have shown that a single intradiscal injection of GDF-5 was associated

with increased disc height and improve degeneration grade compared to untreated discs. A Phase

II clinical trial has been conducted to assess the safety and tolerability of recombinant human

GDF-5 intradiscal injection, the results of which have not yet been published135. While GDF-5 is

now being investigated clinically as a potential biological treatment for DDD, it should be noted

that multiple intradiscal injections of GDF-5 in mice did not have an effect on disc height and in

fact elicited an inflammatory response134. Further work is required to assess its clinical efficacy

and safety. PRP is an autologous source of cytokines, growth and differentiation factors, and

adhesion molecules that are known to be involved in tissue healing136. A bovine study

investigating the effects of PRP administration on AF cells in vitro and whole discs ex vivo found

increased GAG production when compared to untreated cells and discs137. PRP-treated AF cells

also showed increased DNA content over time when compared to untreated cells, implying a

PRP-induced increase in cell proliferation137. However, results from intradiscal PRP injection in

rabbits indicated that PRP did not increase IVD cell proliferation but rather decreased

apoptosis138. Preliminary clinical trials evaluating the efficacy of intradiscal PRP injection have

found improved patient-reported outcomes of back pain and function over follow-up periods of

28

6-12 months139,140. One double-blind, randomized control study reported significant

improvements in self-reported pain, physical function, and satisfaction in patients who received a

single intradiscal PRP injection compared to control patients who received an injection of

contrast agent141. Any biological effects of PRP on patient disc tissue are unknown as they were

not evaluated in these studies. Although radiographic evidence reported by Akeda et al. showed

that PRP-treated discs did not show further ossification or height reduction over time, there were

no differences between treated (L4/L5 level) and untreated (L3/L4 level) discs of the same

patients139. Conclusions about the clinical efficacy of PRP in the intervention of disc

degeneration therefore cannot be drawn. It is still unknown whether PRP specifically targets the

underlying pathologies of DDD, alters pain, or even whether it is an appropriate treatment for

advanced stages of degeneration.

2.3.1.2 Cell-Based Therapies

Cell-based therapies aim to replace cells, restore normal disc biology, and are perhaps of

potential interest for more advances stage of DDD. Clinical intradiscal implantation of

autologous and allogeneic stem cells, allogenic chondrocytes, and autologous disc cells have

been investigated.

Autologous mesenchymal stem cells (MSCs), which can be isolated from autologous bone

marrow or adipose tissue, are of great interest due to their multilineage differentiation potential

and have been shown to be safe and feasible for the treatment of DDD142. Autologous

hematopoietic stem cell (HSC) delivery by direct intradiscal injection of bone marrow aspirate

was not found to be an effective clinical intervention as none of the patients reported reductions

29

in discogenic pain in the year following treatment143. Patients further underwent spinal fusion or

disc replacement surgery after study completion. Although the authors of this study claimed to

have injected HSCs, the marrow aspirate was not processed to isolate specific cell fractions, nor

was the cell population characterized. A more thorough clinical study utilized an isolation and

centrifugation processing system to obtain an autologous bone marrow aspirate progenitor cell

concentrate, which was then injected intradiscally144. The cell preparation was also characterized.

The cell preparation contained progenitor cells characteristic of MSCs as determined by in vitro

assay of fibroblast colony-forming units, and was positive for MSC markers CD90, CD105, and

CD34. Results suggested that bone marrow aspirate containing higher MSC concentrations

compared to those containing lower MSC concentrations led to greater decreases in patient-

reported pain and disability, and to improvements in disc hydration in some patients as

determined by MRI. Despite encouraging results, treatment efficacy appeared to depend on the

patient-specific MSC numbers in each autologous preparation, suggesting that this treatment

might not be a universally appropriate for all patients. Two human trials of intradiscal autologous

mesenchymal stem cell implantation have been published10,11. Yoshikawa et al. presented case

studies of 2 women who received intradiscal graft implants consisting of collagen sponges

seeded with autologous bone marrow-derived MSCs that had been precultured (2-week culture in

one case; 4-week culture in other case) and confirmed for cell proliferation. 2 years following

surgery, patients in both cases reported improved pain and disability scores and showed evidence

of improved disc hydration as determined by T2-weighted MRI signal intensity10. In a pilot

clinical investigation by Orozco et al., ten patients diagnosed with chronic back pain and lumbar

DDD received intradiscal injection of autologous bone marrow-derived MSCs that had been

precultured for 7-10 days11. Patient-reported pain and disability scores improved significantly 3

30

months following the procedure, with further improvements at 6 and 12 months. MRI revealed

that although disc height was not significantly different after each follow-up period, treated discs

showed increased hydration after 12 months. A more recent clinical study was reported to be the

first to see a significant association between MRI metrics and improvements in pain scores after

intradiscal injection of culture-expanded bone marrow-derived autologous MSCs145. 85% of the

treated patients showed decreased disc bulging by MRI imaging, which was related to decreased

pain scores relative to baseline reports. Allogeneic MSCs, which was reported to be more

convenient than autologous MSCs with respect to immediate availability for use in the clinic,

were investigated in a randomized clinical trial12. MSCs were derived from the bone marrow of

healthy donors that were not immune-matched to recipients, expanded for 27 days, and injected

intradiscally. Treated patients reported improvements in pain and disability scores as compared

to the sham-treated control patients at a 12-month follow up, and no significant immune

responses were found12,146. Although MRI imaging showed an improvement in degenerative

grade for MSC-treated compared to control patients, disc height and water content were not

shown to be significantly improved12. It is not shown from these clinical studies whether the

intradiscally transplanted MSCs resolved the underlying degenerative processes. Evidence from

a large-animal (porcine) study suggested that intradiscal delivery of MSCs slows down, rather

than resolves, degeneration147. It is also not known whether the implanted/injected cells acquired

and sustained a disc cell-like phenotype. MSCs have been induced in vitro to differentiate

towards both AF and NP cell-like phenotypes, supporting the idea of pre-culturing MSCs

towards a disc cell-like phenotype prior to implantation15-17. However, acquisition of the correct

phenotype, long-term phenotype maintenance, and the determination of optimal differentiation

conditions remain challenging.

31

Induced pluripotent stem cells (iPSCs) are also of interest as a possible cell source for DDD

treatment. Human somatic cells can be reprogrammed into iPSCs, which then have the capacity

for directed pluripotent differentiation148. Recent in vitro work has demonstrated the ability to

reprogram human NP cells isolated from DDD patients into iPSCs, and subsequent induction of

these iPSCs back into NP-like cells149. Despite this potential, there are currently no published

clinical studies that investigate the efficacy of iPSCs in humans150. Clinical use of both MSCs

and iPSCs poses further safety concerns such as potential for tumour and osteophyte

formation150,151.

Alternatively, a primary disc cell source is perhaps more attractive due to the difficulty of

inducing MSCs to express the correct phenotype. Results from a preliminary clinical study

suggested that intradiscal injection of passaged allogeneic juvenile chondrocytes (NuQu®, ISTO

Technologies Inc.) led to improved disc contour, improved disc height, and diminished annular

tears at 6-month follow-up13. Scores of patient-reported pain and disability significantly

decreased up to 12 months after treatment. However, clinical use of this allogeneic cell treatment

may be limited by the availability of cadaveric juvenile donors, and by costly cell expansion and

cryobanking procedures152. Additionally, chondrocytes produce hyaline cartilage and are

phenotypically distinct from NP cells and certainly from fibrocartilage-producing AF cells18,19. It

is unknown whether intradiscal implantation of these cells will have long-term adverse effects.

Autologous cultured disc cell transplantation (chondrotransplant® DISC by German company

co.don) was investigated in a clinical trial by the name of EuroDisc14. The chondrotransplant®

32

DISC procedure involved removal of herniated disc tissue during regular discectomy, disc cell

isolation and expansion in culture, and intradiscal injection of expanded cells back into the

patient 3 months after the initial discectomy153. Treated patients received discectomy and

subsequent cell transplantation while control patients received discectomy alone. Pain and

disability scores were lower in patients who received the treatment compared to control patients

at a 2-year follow up. Treated discs also showed a higher percentage of normal fluid content

when compared to control discs at 2-year follow up14. However, statistical significance was not

reported for either outcome, and follow-up was not extended past 2 years. While the results of

this study provide some evidence to support the use of autologous disc cells as a treatment, cell

administration by injection is not expected to restore structural organization of the normal disc.

A tissue-engineered therapy that combines autologous cells and a structural template might

therefore be a solution to replicate both form and function of the normal native disc.

2.4 Intervertebral Disc Tissue-Engineering

Research focused on tissue-engineered IVD implant models is currently in the pre-clinical

stages. IVD tissue engineering research has focused on disc models that include only AF and

nucleus pulposus components20,154-157, or the CEP alone158. Many bioengineered IVD models do

not distinguish OAF from IAF and utilize cells from the AF as a whole, thereby failing to

address the fact that the AF is a heterogenous tissue. Engineered IVD constructs have been

generated using collagen hydrogels20,154,159,160 as well as three-dimensional porous scaffolds

made from silk22,156,161, synthetic polymers162, and synthetic polymer/Bioglass composites163.

Regenerative strategies under investigation for NP tissue engineering are hydrogel or scaffold-

based164-166. Given that the gel-like NP lacks the structural complexity of the AF and is relatively

33

homogeneous in comparison, the greatest challenge in disc tissue-engineering research arguably

remains in accurately modeling the circumferential, aligned angle-ply, multi-lamellar structure of

the AF.

2.4.1 AF Tissue Engineering

Various approaches have been undertaken in an attempt to reproduce the specific structural

complexities of the AF. A biphasic IVD model by Choy and Chan mimicked the lamellar AF

structure by repeated photochemical cross-linking of up to 10 layers of collagen gels around a

collagen-GAG co-precipitate NP-like core20. Compared to the engineered constructs with fewer

collagen layers (1-4 layers), those with the highest number at 10 layers showed mechanical

properties similar to native disc. Moriguchi et al. constructed and assessed in vivo implantation

of a tissue-engineered (TE) canine IVD implant model made from a NP cell-seeded alginate gel

core surrounded by two layers of AF cell-seeded collagen gels, which contracted around the NP

core over a 2-week in vitro culture period160. Implanted TE IVDs stained positively for

proteoglycan and collagen by 4 weeks. At 16 weeks post-implantation, AF cells had a distinctly

elongated morphology compared to the rounded NP cells and appeared to align with the

concentric lamellae160. It is unknown from this study alone whether the aligned and elongated

phenotype expressed by AF cells in TE IVDs in vivo was a result of being seeded in a lamellar

structure, or if it was due to the innate behaviour of these cells regardless of scaffold structure.

Evidence that lamellar collagen fibre alignment promotes AF cell alignment was demonstrated

by Bowles et al. in a circumferential construct21. Collagen gels were contracted around a mandrel

to induce circumferential alignment of collagen fibrils. Unlike uncontracted collagen gel disks,

the contracted annular gels were shown to promote alignment and elongation of seeded ovine AF

34

cells. While these models did highlight the circumferential lamellar component of the native AF,

they did not address nor assess the angle-ply alignment of collagen fibres within lamellae.

Work using silk protein has highlighted the importance of having lamellar features in a scaffold

for proper AF tissue formation, and has also been able to incorporate angle-ply features. Park et

al. compared porous to lamellar silk-based scaffolds and found that the latter was superior in

promoting matrix molecule synthesis by porcine AF cells156. However, the freeze-dry method of

lamellar scaffold construction did not reproduce the alignment nor the circumferential structure

that is characteristic of native AF tissue. Bhunia et al. further generated silk-based porous

aligned lamellar scaffolds by using liquid nitrogen to unidirectionally freeze an aqueous silk

fibroin solution22,161. The mechanical strength of these scaffolds was shown to be adjustable by

blending together different ratios of silk fibroin from different silkworm species, and was shown

to modulate the composition of deposited matrix161. These scaffolds were fashioned into

concentric constructs, and lamellar pore alignment was confirmed to be at opposing angles of

~30° by SEM. Over a 2-week culture period, silk fibroin constructs showed alignment of seeded

primary porcine AF cells along the directionally aligned (±30°) lamellar pores, accumulation of

collagen type I and sulfated GAGs, and increased expression of Col1α1, aggrecan, and SOX9.

Constructs seeded with human MSCs (hMSCs) also showed increased expression of the same

genes22. There are a few limitations to these studies. Silk for use as a biomaterial requires

sourcing from cocoons as well as boiling and washing steps for the removal of sericin, a glue-

like substance that otherwise elicits an immune reaction upon implantation161,167. While the

authors stated that macrophage invasion was negligible after 4 weeks of subcutaneous

implantation in mice22, immune cells were still present at this timepoint suggesting a persisting

35

inflammatory response to the implant. The results from the cultured hMSC constructs suggested

chondrogenic differentiation and there was no evidence to say that these cells acquired an AF

cell phenotype; the genes and matrix macromolecules that were reportedly upregulated are

known to be expressed by chondrocytes and are not specific markers of AF cells. Cell

morphology was also not assessed so it not known whether these scaffolds, despite having

directionally aligned lamellar features, were supportive of the AF cell phenotype over the culture

period.

A recent study by Yang et al. used microfluidic technology to generate bacterial cellulose sheets

patterned with microgrooves oriented at 30° angles, which were then stacked at alternating angle

orientations and rolled to create AF-like angle-ply implant structures. In vitro experiments

showed higher elongation of rat AF cells on micropatterned compared to non-patterned

scaffolds, and confirmed cell directional alignment along the ±30° orientation angles. Scaffolds

wrapped around a collagen type II-based hydrogel NP-like core were implanted into the disc

space of discectomized rats and held in place by needles and an external fixation device.

Although the authors reported robust in vivo integration of the acellular TE IVD implant with the

adjacent rat endplate, tissue formation between lamellae, and a resemblance to the intact native

rat disc space after 3 months23, it appeared from the limited histology and MRI imaging that the

TE AF lamellar walls were parallel rather than perpendicular to the endplate, and that the implant

was displaced and at least partially extruded from the disc space. In vitro AF cell experiments

did show that rat AF cells elongated and aligned to the micropatterned scaffolds, however it is

unknown from the in vivo implantation study whether the endogenous cells (of unknown origin)

in the newly formed tissue exhibited these same organizational and morphological properties that

36

are characteristic of the AF cell phenotype. The nature and composition of the newly formed

tissue in vivo was also not evaluated, and it is unknown whether the implant was conducive to

AF-like matrix accumulation.

In contrast to biological materials like silk, bacterial cellulose, or collagen, a biocompatible

synthetic polymer that can be produced in a laboratory setting might be considered a more

efficient and controllable alternative for large-scale, reproducible production of a scaffold

suitable for AF tissue engineering. Xu et al. proposed a method of AF scaffold construction

whereby the synthetic polymer polycaprolactone (PCL) was wet-spun around a 3 mm-diameter

rotating mandrel to generate layers of aligned concentric microfibres168. In vitro culture of rabbit

AF cells on the aligned scaffolds showed cell infiltration, and increased collagen and GAG

accumulation over the 21-day culture period. Content of both collagen types I and II were

significantly increased from day 1 to day 21 of culture, however collagen type I content was

significantly higher than collagen type II throughout the culture period suggesting that the

scaffolds supported and maintained an AF tissue-like phenotype. Consistent with these findings

was the significantly higher fold increase of collagen type I and aggrecan gene expression

compared to that of collagen type II expression over the 21-day culture period. However, this

model did not replicate all physical features of the native AF, as the method of circumferential

fibre deposition did not allow for an angle-ply structure to be produced.

It is challenging to model all of the native AF structural features (circumferential, multilamellar,

aligned, angle-ply) simultaneously within a single bioengineered construct. However, it is

evident from the above studies that a scaffold that encapsulates them all would be instrumental in

37

appropriate AF cell maintenance and tissue formation. The following section discusses

electrospun nanofibrous scaffolds in the construction of a synthetic biocompatible implant that

encapsulates all the structural complexities of the native AF, supports AF cell phenotype

maintenance, and promotes AF tissue formation.

2.4.2 Electrospun Aligned Nanofibrous Polymer Scaffold for Multilamellar Tissue-

Engineered AF

Electrospinning provides a means of producing sheets of aligned nanofibrous polymer scaffolds

that mimic the aligned collagen fibres of native AF tissue and that can subsequently be cut,

stacked at alternate angles, and rolled into concentric angle-ply structures24,25. AF cells seeded

onto these scaffolds align with the nanofibers, allowing for the resulting matrix deposition to be

more closely matched to native tissue in their physical appearance and properties. AF tissue

matrix components such as type I collagen, aggrecan, and GAG were increased with the use of

aligned nanofibrous scaffolds when compared to scaffolds composed of randomly-oriented

nanofibers29. Increased nanofiber alignment achieved through a mandrel rotation speed of 1400-

2000 rpm resulted in elongated morphology of seeded AF cells and AF-derived stem cells,

similar to the shape of OAF cells observed in native tissue29,169. Yang et al. generated IVD

implants with an NP-like hydrogel core surrounded by aligned multilamellar angle-ply

nanofibrous sheets of co-electrospun PCL, poly(lactic-co-glycolic acid) (PLGA), and collagen

for increased hydrophilicity. In vivo results at 1 and 3 months post-implantation in rats showed

that these PCL/PLGA/collagen-containing acellular constructs were hydrated and maintained

disc height at levels comparable to native disc. In contrast, constructs made of hydrophobic

PLGA alone showed poor hydration and decreased in height170. While this study demonstrated

38

the importance of an appropriately hydrophilic biomaterial for AF tissue-engineering, the

acellular nature of the construct was not demonstrative of a fully functionalized IVD implant.

Martin et al. developed an IVD model referred to as a Disc-like Angle Ply Structure (DAPS),

which was generated from electrospun PCL arranged in a multilamellar ±30° angle-ply

concentric structure26-28. Early work on the model determined that incorporation and subsequent

dissolution of a sacrificial polymer layer within electrospun PCL DAPS construct created

interlamellar spaces thereby improving cell infiltration in vitro26. DAPS made from electrospun

PCL strips alone were densely packed and showed poor infiltration of bovine AF cells in vitro

due to the tight spaces between scaffold layers. Acellular DAPS wrapped around a hyaluronic

acid (HA)-based hydrogel NP-like core were implanted into caudal rat spines27. Collagenous

matrix accumulation within the acellular DAPS implant was shown in vivo over the course of 12

weeks, however this process relied on endogenous cell infiltration. In comparison to an acellular

biomaterial implant which may also have poor endogenous cell infiltration in vivo, a cell-based

AF implant consisting of pre-formed engineered tissue might have improved compositional

characteristics and post-operational outcomes26,27. TE IVDs comprised of bovine AF cell-seeded

DAPS and NP cell-seeded HA hydrogel cores showed accumulation of GAG and collagen over a

15-week culture period in vitro, and were subsequently implanted into rat caudal spines28. This

study did not include a control experiment comparing the in vivo maturation of cell-seeded

DAPS implants to that of acellular implants, so it cannot be concluded from here that cell-seeded

implants perform with superiority in vivo. However, it was demonstrated that pre-seeded DAPS

accumulated ECM over time in vitro, thereby ensuring the presence of tissue upon implantation

and eliminating the reliance on endogenous cell infiltration.

39

Electrospun nanofibrous scaffolds for AF tissue generation have been made from various

polymers such as polylactic acid (PLA)171, PCL24,25,169,172,173, and polycarbonate-urethane

(PU)29,174,175. Electrospun nanofibers are smaller in diameter compared to microfibres, which

have previously been produced for AF construction by methods such as wet-spinning. Wet-

spinning of PCL showed that decreased microfiber diameter was achievable by increasing

mandrel rotation speed, however this resulted in fusion between microfibers, low scaffold

porosity, and consequently poor AF cell infiltration168. Smaller electrospun polymer fibre

diameter (≤ 1µm) improved human umbilical vein endothelial cell adhesion and promoted cell

elongation compared to larger fibre diameter (7 µm), on which seeded cells exhibited a rounder

morphology and low proliferation176. Electrospinning of PU for AF generation has produced

fibres with diameters of 200-400 nm177. PU in particular is a promising candidate for AF tissue

generation due to its tensile strength, elasticity, biocompatibility, and non-toxicity to AF cells177.

Although both aligned nanofibrous PCL and PU scaffolds seeded with AF cells produce ECM

that mimics the composition of native AF tissue, in comparison of both, PU has been found to

have better resistance to stretch and higher initial cell seeding efficiency178. Its initial modulus is

also tailored to be closer that of native ECM than the more crystalline PCL material.

Additionally, PLA forms acidic degradation products at concentrations that readily lower pH

levels and might lead to a decrease in ECM synthesis179,180. Increasing the surface energy of PU

scaffold by incorporating anionic dihydroxyl-oligomer (ADO) into the electrospun polymer was

found to increase scaffold hydrophilicity as indicated by a decreased water contact angle, and to

enhance AF cell attachment and collagen accumulation without significant cytotoxic

effects175,177. PU-ADO based polymer scaffold is particularly attractive for tissue engineering

40

purposes due to: (1) Its elastic properties and tensile strength; (2) its reproducibility, which

allows for reliable large-scale production; (3) its hydrophilic surface chemistry which mimics the

surface characteristics of proteins that are otherwise costly and that may have varying genetic

factors that cannot be controlled for; and (4) its biocompatible byproducts175,177. Iu et al. showed

that both IAF and OAF cells cultured separately on electrospun nanofibrous PU lamellar

scaffolds were able to maintain their respective phenotypes and produce appropriately distinct

ECM30,31.

2.4.3 Tissue-Engineered IVD-CEP Integration

One function of the native CEP is to provide a transitional substrate for IVD anchorage into the

underlying vertebral bone. There are currently few if any studies that have specifically generated

a cell-based tissue-engineered IVD-CEP interface model that contains a hyaline cartilage

component and that also recapitulates the structure of the aligned multilamellar AF. One in vitro

study by Tsai et al. incorporated an endplate component into an IVD model using human MSC-

derived osteoblasts induced to undergo chondrogenic differentiation, however the AF-endplate

interface itself was not the main focus of the study and was not characterized40.

Implantation studies have examined the performance of nanofibrous TE AF construct

implantation in rats, however these acellular implants relied on the slow processes of post-

operative endogenous cell infiltration and in vivo tissue integration26,27,170. Martin and colleagues

acknowledged that while implantation of an acellular DAPS into rat caudal spines showed

endogenous cell infiltration, integration with adjacent vertebrae did not occur and implant

extrusion from the disc space was evident by 12 weeks27. Poor integration and extrusion of a

41

cell-seeded DAPS implant from the intervertebral space also occurred28. Cell-seeded DAPS 5

weeks post-implantation displayed proteoglycan loss, lack of integration with adjacent vertebrae,

and extrusion from the disc space regardless of either a 5 or 10 week pre-culture period28. This

was an indication that tissue integration must be considered going forward in the development of

a feasible bioengineered IVD implant. Subsequently, an acellular porous polymer endplate

component allowing for bony ingrowth was incorporated into the implant model in an attempt to

facilitate integration and generate mechanical continuity along the vertebra28. The engineered

endplates appeared to mitigate the loss of proteoglycan and improved integration with adjacent

vertebrae after implantation, which emphasized the importance of having integrated components,

particularly endplates, in an IVD implant model. However, the cartilaginous layer that lies

between the IVD and the bony endplate within the native disc space, which is in itself essential

for disc-vertebra anchorage and integration, was disregarded in this DAPS-endplate model. The

distributions of collagen types I and II were also diffuse and not distinctly localized within the

DAPS in vitro, thus the engineered AF did not have distinct inner or outer regions28. To resolve

these issues that arise from the innate complexity of the native IVD, a more complete implant

model would require distinct IAF and OAF regions, as well as an integrated CEP layer that

would further be able to integrate with bone.

Inclusion of a cartilaginous endplate component is desired for a more complete IVD implant

model; however, without a pre-formed IVD-CEP interface prior to implantation, appropriate

tissue integration in vivo might be difficult and slow. For example, integration of cell-seeded

collagen gel AF-NP constructs with native rat endplate showed discontinuities 6 weeks after

implantation, and continuous integration did not occur until 6 months post-operatively154.

42

Integration of tissue-engineered AF constructs specifically with an engineered CEP has not been

an area of focus in current literature. Integration of engineered disc and cartilage tissues prior to

implantation might be a solution to circumvent poor implant integration in vivo. However,

culture times may be long, and sub-optimal in vitro conditions might not allow for maximal

integration potential. Therefore, it would be of great use to determine applicable treatments that

would be able to enhance or accelerate the integration process in vitro. Mechanical stimulation is

one method that has been shown to be of potential use for the enhancement of tissue matrix

accumulation and mechanical properties.

2.5 The Role of Mechanical Stimulation on Extracellular Matrix Composition

It has been shown that in IVD tissue, IVD cells, tissue-engineered IVDs, and tissue-engineered

cartilage, certain frequencies and amplitudes of dynamic tensile and compressive strain can have

anabolic effects on matrix components by stimulating protein and gene expression, and may play

a role in reducing matrix protein degradation32-42. Table 1 summarizes the findings of the current

literature on compressive, tensile, and static loading on IVD and cartilage cells and tissue.

43

Table 1. Literature review of studies that have investigated the effect of cyclic tensile strain, cyclic compressive strain, and static

compression on IVD cells, IVD tissue, and tissue-engineered IVDs and cartilage. Timepoints for cell/tissue analysis post-loading

were assumed to be immediate unless otherwise stated.

DYNAMIC COMPRESSION

Authors Source species, age

Tissue/Cell Type

Culture conditions Load Conditions Harvest time post-loading

Findings

Cells

Kasra et al. 2003

Adult rabbit OAF cells Monolayer expansion to P1, seeded in culture dishes overnight before loading

1 Hz, 0.3 MPa, (low); 20 Hz, 1.7 MPa (high); 30 mins/day for 3-9 days

Immediately Compared to unloaded control: High loading collagen synthesis; low loading had little effect37

Kasra et al. 2006

Pig, 10 months old

AF cells Primary cells in 3D alginate beads, cultured 24 hrs before treatment

1, 3, 5, 8, or 10 Hz; 1.0 MPa; 30 mins a day for 3 days

Immediately Compared to unloaded control: Maximum collagen synthesis and release at threshold of 5 Hz36

Sharma et al. 2007

Goat, 2-3 months old

Full thickness articular chondrocytes

Primary cells cultured in alginate beads for 3 days prior to loading. 10 day total culture time.

0.66 Hz; 1.2 or 2.4 MPa; Once for 4 hrs

Immediately Compared to unloaded control: Dose-dependent GAG content, metalloprotease activity43.

Korecki et al. 2009

Young (4-6 months) and mature (18-24 months) cow

AF cells Tissues kept in culture for 4-6 days before cell isolation. Monolayer expansion under 9 days to P1, seeded in 3D alginate gel culture

0.1, 1, or 3 Hz; 2-12% vertical displacement of initial gel height; 2 hrs/day for 7 days

Immediately Compared to unloaded control: COL1, COL2, ACAN, and MMP3 gene expression were affected by age and loading frequency. Significant interactions found between age and loading frequency44.

Hee et al. 2010

Human adolescent idiopathic scoliosis patients

IAF tissue/cells

IAF tissues received compression treatment. Cells subsequently isolated, cultured 7 days in suspension

1 Hz; 0.2 MPa (moderate) or 0.4 MPa (high) 2 hrs/day for 7 days

7 days Compared to unloaded control: Moderate load GAG and collagen content; high load GAG content42.

44

Native Tissue

MacLean et al. 2004

Mature rat, >12 months old

Whole IVDs N/A (in vivo experiment) 0.01, 0.2, or 1 Hz; 0.2 or 1.0 MPa; Once for 2 hrs

Immediately Compared to unloaded sham control discs: AF catabolic gene expression at higher loading amplitude across all frequencies; 1 Hz AF COL1 at low and high amplitudes, ACAN at low amplitude45

MacLean et al. 2005

Mature rat, >12 months old

Whole IVDs N/A (in vivo experiment) 1 Hz; 1 MPa; Once for 0.5 or 4 hrs

Immediately Compared to unloaded sham control discs: Longer duration (4 hrs) AF COL1, COL2, MMP-3, MMP-13 gene expression46

Wang et al. 2007

Rabbit, 12 weeks old

Whole IVDs In vitro whole organ culture 0.1 or 1 Hz; 0.5-1 MPa Once for 6 hrs

Immediately Compared to unloaded control: IVD and CEP collagen, aggrecan gene expression47.

Korecki et al. 2008

Mature bovine, 18-24 months old

Whole caudal IVDs

In vitro whole organ culture 1 Hz; 1.0 MPa (low), 2.5 MPa (high); 1 hr/day for 5 days

Immediately Compared to static loaded control: COL1 gene expression in AF by high loading, COL1 by low loading48.

Li et al. 2016

Pig, 3-4 months old

Whole IVDs (AF tissue analyzed)

In vitro whole organ culture 1 Hz; 0.1, 0.2, 0.4, 0.8, or 1.3 MPa for 2 hrs/ day, 7 days; or 0.4 MPa for 1, 2, 4, or 8 hrs/day, 7 days

Immediately Lower compressive magnitudes and durations or maintained PG and aggrecan staining intensity, GAG and collagen content, and ACAN and COL1 gene expression. Highest compressive magnitude and longest duration led to catabolic changes in ECM deposition and gene expression41

Engineered Tissue

Davisson et al. 2002

Bovine, ~3 months old

Tissue-engineered cartilage

Primary chondrocytes cultured on fibrous polymer scaffolds for 3 weeks

0.001 or 0.1 Hz; 10% or 50% displacement + 5% oscillation amplitude; Once for 24 hrs

Immediately Compared to unloaded control: Dynamic compression total protein and S-GAG synthesis38.

45

Waldman et al. 2004

Bovine, 6-9 months old

Tissue-engineered cartilage

8-week total culture period on calcium polyphosphate substrate surface prior to loading

1 Hz; 5% displacement; 400 cycles every other day for final 2 or 4 week culture period

Immediately Compared to unloaded control: 4 week loading tissue thickness, proteoglycan and collagen content, stiffness and load bearing capacity49.

Waldman et al. 2006

Bovine, 6-9 months old

Tissue-engineered cartilage

1, 8 or 15 days in culture on calcium polyphosphate substrate surface prior to loading

1 Hz; 9.81 mN (1 g) amplitude; Once for 15, 30, or 60 min

Immediately Compared to unloaded control: Early cultures (1 day old) most responsive to loading; greatest proteoglycan and collagen content from 30-min loading50.

Tsai et al. 2011

Human discectomy and hip arthroplasty patients, 35-72 years old

AF, NP cells co-cultured with hMSCs; seeded in electrospun polymer construct

P2 and P4 AF and NP cells co-cultured with hMSCs, seeded overnight on scaffold strips.

1 Hz; 10% strain; 1 hr/day for 21 days

Immediately Compared to unloaded control: SOX9, ACAN gene expression in loaded engineered AF compared to unloaded, trend towards COL1, COL2 expression; GAG content unaffected, although deposition on scaffold suggested to be enhanced40.

Hudson et al. 2015

Sheep, 14-months old

AF cell-seeded collagen gel + NP cell-seeded alginate gel IVD construct

AF and NP cells cultured for 2 weeks in suspension, trypsinized, seeded in collagen and alginate gels, respectively.

1 Hz; 1%, 2.5%, 5%, or 10% displacement; 1 hr on, 1 hr off, 1 hr on every other day for 2 weeks

Immediately Compared to unloaded control: Collagen, GAG, and viscoelastic properties (equilibrium and instantaneous moduli); full effects observed at 5% strain39.

CYCLIC TENSILE STRAIN

Authors Source species, age

Tissue/Cell Type

Culture conditions Load Conditions Harvest time post-loading

Findings

Cells

Rannou et al. 2003

Rabbit (age unspecified)

AF cells P1 cells cultured on flexible bottomed plates

1 Hz; 1% or 5% strain; Once for 0.5-24 hrs

Immediately Compared to static control: 5% strain newly synthesized proteoglycan content; Mechanism post-translational and nitrite oxide-dependent51

46

Sowa and Agarwal 2010

Rat, 10-12 weeks old

AF cells P2 cells plated on col1-coated silicone membranes, grown to 95% confluence; cells maintained in fresh media 18 hrs before loading

0.05 Hz; 6% strain; Once for 4 hrs

Immediately Compared to unloaded control: IL-1β-induced gene expression of inflammatory response mediators (iNOS, TNF-α) and matrix degrading enzymes (MMP-3, MMP-13)52

Gilbert et al. 2010

Human, 29-66 years old

Degenerated and non-degenerated AF cells

Monolayer; Passage ≤ 6 on BioFlex culture plates

0.33 Hz (below physiologic), 1 Hz (physiologic); 10% strain; Once for 20 min

0-24 hrs Compared to unloaded control: 1 Hz load catabolic gene expression (MMP-3, ADAMTS-4) in nondegenerative AF cells; 0.33 Hz COL1 and COL2 gene expression, MMP-9 expression33

Gilbert et al. 2011

Human, 37-57 years old

Degenerated and non-degenerated AF cells

Monolayer expansion; P≤ 6 on BioFlex culture plates

1 Hz; 10% strain; Once for 20 min

0-24 hrs Compared to unloaded control: Catabolic gene expression (MMP3, ADAMTS4) in nondegenerative cells; effect modulated by cytokines IL-1 and IL-432

Li et al. 2011

Bovine, 7 days old

OAF cells 4 day monolayer culture until confluent; P1 cells reseeded in monolayer on col1-coated Flex I culture plates

1 Hz; 10% strain; Once for 1 hr

Immediately Compared to unloaded control: OAF cell elongation; β-actin, β-tubulin gene and protein expression; COL1, COL2 gene expression; differentially regulated expression of matrix-degrading enzymes34

Native Tissue

Li et al. 2017

Porcine, 2-4 months old

Whole IVDs In vitro whole organ culture 0.1, 0.5, 1, 3, or 5 Hz; 0.4 MPa; 2 hrs/day for 7 days

Immediately Compared to unloaded control: GAG content and COL1/COL2/ACAN gene expression or maintained at 0.1-3 Hz, at 5 Hz. Greater catabolic gene expression profile (MMP-3, ADAMTS-4, TIMP-1, TIMP-3) at 5 Hz compared to 0.1-3 Hz35.

Engineered Tissue

Fotticchia et al. 2018

Adult human Marrow-derived MSCs seeded in electrospun PCL construct; 2% NP-like agarose plug

Monolayer cultured cells trypsinized once, seeded between double layer of electrospun PCL; cultured 2 days on PCL before loading

1 Hz; 5% or 10% vertical displacement; 8 hrs/day for 3 days

Immediately Compared to static control: Cell viability in both test conditions53.

47

STATIC COMPRESSION

Authors Source species, age

Tissue/Cell Type

Culture conditions Load Conditions Harvest time post-loading

Findings

Cells

Chen et al. 2004

Porcine, 4-5 months old

AF cells Monolayer expansion to P1, cultured in alginate discs 24 hrs prior to loading

25% compressive strain; Once for 2, 8, or 30 hrs

Immediately Compared to unloaded control: COL1, COL2, ACAN, and SLRP gene expression at 30-hr time point181.

Sharma et al. 2007

Goat, 2-3 months old

Full thickness articular chondrocytes

Primary cells cultured in alginate beads for 3 days prior to loading. 10-day total culture time.

1.2 or 2.4 MPa; Once for 4 hrs

Immediately Compared to unloaded control: Dose-dependent GAG and collagen content; metalloprotease activity43.

Native Tissue

Ohshima et al. 1995

Cow, 2 years old

Whole IVDs In vitro whole organ culture 0.5-15 kg; 1-8 hrs

Immediately Disc hydration with increasing load. Proline and sulfate incorporation (indicative of matrix synthesis rate) lowest at 0.5 kg, doubled at 5-10 kg, decreased at 15 kg. Maximum synthesis rate occurred when hydration level in vitro was similar to what is naturally found in vivo182.

Lotz et al. 1998

Mouse, 12 weeks old

Whole IVDs N/A (in vivo experiment, force applied by external compression device)

0.4, 0.8, or 1.3 MPa; 1 week

Immediately or after 1-month recuperation post-loading

Compared to unloaded control: Separation of NP cell clusters; AF lamellar disorganization (greater effect in inner & middle AF compared to OAF). Collagen type II mRNA expression in all regions. Cell death in peripheral NP and IAF at low load, further extended into AF and CEP at high load183.

Lotz and Chin 2000

Mouse, 12 weeks old

Whole IVDs N/A (in vivo experiment, force applied by external compression device)

0.4, 0.8, or 1.3 MPa for 7 days; or 1.3 MPa for 1 or 3 days

Immediately Compared to unloaded control: Cell death in IAF of loaded discs. Increased loading and time led to apoptosis in CEP. Further apoptosis into NP at highest load & time (1.3 MPa, 7 days)121.

48

Ariga et al. 2003

Mouse, 7 weeks old

Whole IVDs In vitro whole organ culture 0-1.0 MPa; Once for 24 hrs

Immediately Compared to unloaded control: Apoptotic cell death in CEP and AF-CEP interface with increasing static load120.

Wang et al. 2007

Rabbit, 12 weeks old

Whole IVDs In vitro whole organ culture 0.5-1 MPa Once for 6 hrs

Immediately Compared to unloaded control: IVD and CEP collagen and aggrecan gene expression; apoptotic cell death highest at 1 MPa static compression47.

Paul et al. 2013

Adult goat, 3-5 years old

Whole IVDs with attached CEPs

In vitro whole organ culture 0.6 MPa; 16 hrs/day for 7, 14, or 21 days

Immediately Compared to unloaded control: Cell viability and density in OAF at each time point; less pronounced cell loss in NP. COL1 gene expression in NP, in AF. Catabolic and inflammatory genes. GAG content after 21 days184.

Engineered Tissue

Davisson et al. 2002

Cow, ~3 months old

Tissue-engineered cartilage

Primary chondrocytes cultured on fibrous polymer scaffolds for 3 weeks

10%, 30%, or 50% displacement; Once for 24 hrs

Immediately Compared to unloaded controls: 50% static compression total protein and S-GAG synthesis38.

49

Cyclic tensile strain applied to AF cells at 1 Hz has been shown to decrease catabolic gene

expression, upregulate collagen gene expression, and have a role in proteoglycan production

modulation32-34,51. Compared to uncompressed controls, dynamic compression of immature

porcine IVDs at low frequencies of 0.1-3 Hz maintained or promoted ECM synthesis, while

compression at a relatively higher frequency of 5 Hz led to decreased ECM molecule

accumulation and increased catabolic gene expression35. Similarly, a study on the dynamic

compression of porcine IVD cells in 3D culture found that among uncompressed controls and

discs under loading frequencies of 1-10 Hz, decreased collagen synthesis and maximum collagen

degradation occurred at a frequency threshold of 5 Hz36. It is suggested that a loading frequency

of 0.1–1 Hz provides a “physiologic” range by which dynamic mechanical loading may be

favourable for IVD matrix stimulation and repair185. Frequencies below the physiologic value of

1 Hz may cause a shift towards matrix catabolism33.

In addition to compressive frequency, matrix accumulation appears to be dependent on

compressive magnitude and duration. Hee et al. found that a moderate level of 0.2 MPa dynamic

compression increased human AF and NP collagen and GAG content compared to uncompressed

control cells, whereas a higher compression magnitude of 0.4 MPa decreased content42. In

contrast, Kasra et al. found that high dynamic loading amplitude and frequency (1.7 MPa, 20 Hz)

led to greater collagen synthesis in rabbit OAF cells compared to unloaded cells, while low

amplitude and frequency (0.3 MPa, 1 Hz) had little effect37. However, it was not possible from

the study design to determine whether the apparent effect on collagen synthesis was due to

loading amplitude or loading frequency. Maclean et al. found that catabolic gene expression in

the AF of an in vivo rat model was upregulated across all loading frequencies (0.01–1 Hz) when

50

compared to sham controls, but only by the higher loading magnitude tested (1.0 MPa)

suggesting that AF matrix metabolism may be more dependant on compressive magnitude rather

than frequency45. The same research group found that in comparison to 30-minute dynamic

compression, a longer duration of 4 hours significantly upregulated gene expression of collagen

type I, collagen type II, MMP-3, and MMP-13 in the AF compared to unloaded sham controls46.

Li et al. demonstrated that prolonged dynamic compressive stimulation for 8 hours led to

catabolic changes in matrix deposition and gene expression in comparison to uncompressed

controls, while stimulation between 1 and 4 hours had anabolic effects41. There may therefore be

a duration threshold that is optimal for matrix synthesis and/or maintenance.

In contrast to dynamic compression, static compression has been shown to predominantly have

catabolic effects on matrix molecules38. Static mechanical loading of IVD explants has been

found to lead to CEP chondrocyte apoptosis, which may further contribute to IVD

degeneration120. Acute ex vivo vibratory stimulation of IVDs transiently increased anabolic gene

expression and matrix molecule accumulation, however repeated long-term vibration-type

stimulation interestingly resulted in degenerative changes186-188.

There is a gap in the current literature regarding the effect of dynamic loading on nanofibrous

angle-ply multilamellar electrospun constructs seeded with AF cells. Two studies have

investigated the effect of dynamic mechanical loading on cell-seeded electrospun IVD models,

however the conclusions that can be drawn from them are perhaps limited due to their use of

MSCs40,53. Fotticchia et al. recently showed that in comparison to static control conditions,

dynamic tensile strain of a double layered AF-like electrospun PCL-based construct decreased

51

cell viability of seeded human bone marrow-derived MCSs. The AF construct was wrapped in an

angle-ply fashion around a central NP-like agarose plug, which was then subjected to 5% or 10%

dynamic vertical displacement. In this design, the radial stress exerted from the agarose plug

transferred into circumferential stress in the AF layer, thus causing the electrospun cell-seeded

fibres experience tensile strain. Sinusoidal compression was applied at 8 hours a day for 3 days

and was found to decrease cell viability at both amplitude conditions, suggesting that dynamic

mechanical stimulation may be detrimental to an angle-ply, lamellar IVD construct53. However,

it must be noted that dynamic compression was applied to the constructs only 2 days after initial

cell seeding. The purpose of engineering a scaffold-based IVD is to form tissue, and 2 days may

not be sufficient time for cells to form substantial tissue. Had the cell-seeded constructs been

given time to form tissue over a longer culture period, results may have differed. The 8-hour

application time was also excessive compared to other studies (see Table 1), and was previously

found to lead to catabolic changes in whole IVDs41. Additionally, the reduced cell viability in

response to cyclic compression may have been related to the fact that marrow-derived MSCs do

not normally experience the same types of forces that AF cells experience in the spine.

Tsai et al. also applied dynamic compression to engineered IVDs composed of electrospun

PLA/PCL angle-ply layers seeded with AF cells and MSCs, and a central NP-like synthetic

hydrogel matrix (PuraMatrix) seeded with NP cells and MSCs. An “optimal ratio” of co-cultured

human AF or NP cells and human MSCs was determined by gene expression analyses and

seeded in their respective scaffolds. Dynamic compression of these constructs was suggested to

promote homogenous cell and matrix distribution within the scaffolds. It was also suggested

from gene expression analyses that dynamic loading inhibited the differentiation of MSCs

52

towards hyaline chondrogenesis and promoted fibrochondrocyte (AF-like) differentiation40.

However, they did not provide convincing evidence of anabolic effects on the AF ECM through

their histology images, and immunohistochemistry to identify specific matrix molecule

accumulation was not performed. In addition to evaluating the effect of dynamic loading, the

difficulty of co-culturing AF cells and MSCs in an appropriate culture medium to induce an AF-

like phenotype added further complexity to the system. Investigation of dynamic compression on

an electrospun angle-ply multilamellar construct cultured with AF cells alone would provide a

more reflective determination its effectiveness for promoting AF tissue growth, and would fill

the gap in the literature.

The current literature demonstrates that the effects of dynamic mechanical stimulation on matrix

maintenance and synthesis are dependent on a complex set of parameters. The IVD in vivo is

constantly under mechanical load, however the role of dynamic compression in the formation,

maintenance, and mechanical strength of the specialized AF-CEP interface is unknown.

Development of a bioengineered nanofibrous scaffold-based AF-CEP model and placing it under

mechanical stimulation will help in furthering this understanding.

53

CHAPTER 3: DEVELOPMENT OF AN IN VITRO OUTER

ANNULUS FIBROSUS-CARTILAGE ENDPLATE MODEL AND

ITS RESPONSE TO DYNAMIC MECHANICAL LOADING

3.1 Introduction

Back pain is a chronic condition that presents direct healthcare costs estimated to exceed 20

billion dollars per year in the United States alone189. It is associated with degenerative disc

disease (DDD) and pathologic changes in the vertebral cartilage endplate5,126,190,191. Degenerative

processes in these tissues are irreversible due to their inherent lack of regenerative capacity192.

Conservative treatments for DDD-related back pain such as physical therapy and pain

medication do not target the underlying degenerative pathologies. Surgical interventions include

discectomy, fusion of vertebral segments, or prosthetic disc replacements; however, these

treatments do not guarantee long-term pain relief and may lead to further complications such as

adjacent segment degeneration8,129. Current research is focused on investigating and ultimately

implementing a cell-based tissue-engineered implant solution that replaces the degenerate native

IVD thereby restoring function, and minimizing further degeneration in adjacent discs193.

The intervertebral disc (IVD) consists of a central nucleus pulposus (NP) and a surrounding

annulus fibrosus (AF). Hyaline cartilage endplates (CEP) are located on the cranial and caudal

aspects of each disc. The IVD is a relatively avascular tissue that is contained in a low oxygen

environment. It obtains nutrients primarily via diffusion from small blood vessels that extend

from the vertebral bone marrow into the hyaline portion of the CEP, and from vessels that are

located at the periphery of the AF2,4,54. It is a tissue of relatively low cellularity; the average cell

54

density of the human IVD has been reported to be ~6000 cells/mm3, approximately 2.5 times

lower than that of articular cartilage which in itself is known to be a tissue of low cellularity2.

While the NP is gel-like and does not possess a high level of organization, the AF is a

multilamellar angle-ply structure that is further organized into inner and outer regions, each

having their own distinct composition. The collagen fibres within successive AF lamellae are

alternately oriented at ±30° angles with respect to the vertical axis89. The main collagens of the

IVD are the fibrillar collagen types I and II, which provide tensile strength and compressive

resistance respectively58. The outer AF (OAF) is composed of highly aligned collagen type I

fibres and elongated fibroblast-like cells. The inner AF (IAF) is less organized and composed of

collagen type I, collagen type II, proteoglycans, and rounded cells7,90. Proteoglycan content of

the IVD is highest within the NP and lowest in the OAF, with percent water content generally

following the same trend56. The main proteoglycan found in the IVD and CEP is aggrecan,

which is similar to collagen type II in its regional distribution and in its role to provide

compressive resistance64. The CEP is the site of attachment and mechanical load transfer

between the relatively soft IVD fibres and stiff vertebral bone (VB)1,100-102. A layer of calcified

cartilage within the CEP lies adjacent to the VB, which serves to provide further anchorage of

the disc into the bone1,63. AF fibres insert into the CEP and also curve laterally around it at the

periphery63. With increasing age, the hyaline CEP becomes mineralized in the region between

the OAF and vertebral rim99. Tissue engineering of this interface remains a key challenge in

establishing a functional biological disc replacement.

Tissue-engineered cell-based IVD models that have been developed to date have included AF

and NP components only, and have lacked any in vitro-formed CEP-like cartilage tissue

55

components20,154-157,194. Concentric multilamellar angle-ply structures have successfully been

generated to mimic the complexity of the AF using biomaterials such as silk fibroin22,161,

bacterial cellulose23, and electrospun nanofibrous polymer scaffolds24,25,27,28,31,169,170,195. Tissue-

engineered NP-like hydrogel and scaffold components have been generated from biomaterials

such as polysaccharide, glycosaminoglycan, collagen, and synthetic

polymers20,21,27,28,39,154,155,157,160,170. Both acellular20,27,170 and cell-seeded tissue engineered NP-

like cores21,28,39,160,194 have been used in the generation of biphasic AF-NP constructs. However,

to the best of our knowledge, there are no studies describing the generation of an integrated cell-

based IVD-CEP interface model. The importance of integrated tissues within an engineered IVD

implant model was demonstrated by Martin et al. who implanted a multilamellar angle-ply AF

structure into rat spines, and found that the implant extruded from the disc space due to a lack of

appropriate stabilization between the native vertebral bodies28. Subsequent incorporation of an

acellular porous polymer endplate at the ends of the angle-ply structure improved the integration

of the implant within the native disc space28, however this model lacked a cartilaginous layer to

accurately mimic what is seen in the native AF-VB region63. The CEP, which is in itself essential

for AF fibre anchorage into the native VB, should be included in an implant model in order to

establish proper biomechanics during loading, to facilitate nutritional diffusion, and to mimic the

integration of tissues that naturally occurs in vivo. Integration of in vitro-generated tissues upon

implantation in vivo has been shown to be difficult and slow; integration of cell-seeded collagen

gel AF-NP constructs with native rat endplate showed discontinuities 6 weeks after implantation,

and continuous integration did not occur until 6 months post-operatively154. Therefore,

generation of an IVD-CEP interface in vitro prior to implantation might be a way to circumvent

poor in vivo integration during the post-implantation period.

56

In a previous study, an optimized AF tissue culture system was developed whereby distinct and

robust IAF and OAF tissues were generated using a single culture medium cocktail31. IAF and

OAF cells seeded separately on aligned nanofibrous polycarbonate urethane-anionic dihydroxyl

oligomer (PU-ADO) scaffolds maintained their respective phenotypes when cultured in DMEM

media supplemented with fetal bovine serum, insulin-transferrin-selenium (ITS), proline,

dexamethasone and pyruvate31. Additional work has also validated a 3-dimensional (3D) culture

method to generate in vitro hyaline cartilage from primary chondrocytes196. Having these two

tissues will allow for the development of a co-culture model to address the issue of tissue

integration. Furthermore, it will facilitate investigations into the effect of dynamic loading as a

potential treatment to enhance integration between bioengineered interfacial IVD tissues.

Thus the hypothesis of this study was that the placement of in vitro OAF tissue formed on

aligned nanofibrous multi-lamellated PU-ADO scaffolds in direct-contact and co-culture with in

vitro-generated cartilage tissue formed in 3D culture will generate an integrated interface

containing extracellular matrix (ECM) that resembles native OAF-CEP. Additionally, it was

hypothesized that this interface can be mechanically and compositionally enhanced through

dynamic compressive loading.

3.2 Methods

3.2.1 Isolation of OAF cells

Intervertebral discs were aseptically excised from 6-9-month old bovine caudal spines within 4

hours of death and placed in Ham’s F12 media. OAF was isolated from the rest of the disc and

57

fine-diced. Tissues were enzymatically digested in 0.3% protease (P5147, Sigma) for 1 hour at

37°C, followed by 0.2% collagenase A (COLLA-RO, Roche Diagnostics) overnight at 37°C.

Cells from digested tissue were filtered through a cell strainer and centrifuged at 600 RCF for 8

minutes. Cell pellets were washed 3 times in Dulbecco's modification Eagle's medium (DMEM;

4.5 g/L glucose; Multicell) supplemented with 5% fetal bovine serum (FBS; Multicell). Viable

OAF cells were counted by staining with trypan blue and a hemocytometer, and re-suspended in

DMEM supplemented with 5% FBS at a density of 1.6 x 106 cells/mL (Figure 3.1.A).

3.2.2 Generation of aligned nanofibrous PU-ADO scaffold constructs

Anionic dihydroxyl oligomer (ADO) and polycarbonate urethane (PU) were synthesized as

described previously175. Briefly, in a 20 mL glass scintillation vial, ADO was added to PU at a

concentration of 0.15% by dry weight of PU. 1,1,1,3,3,3-hexafluoro-2-propanol (105228,

Aldrich) was added to yield 18% concentration by dry weight of PU-ADO to volume of solvent.

The solution was sealed tightly, protected from light and allowed to fully dissolve overnight in a

rotator at room temperature. Viscosity was standardized by allowing the solution to flow out of a

1 mL syringe (plunger removed) fitted with a blunt fill 18G x 1½” needle, and ensuring that the

time taken for the volume to decrease from 0.7 mL to 0.6 mL was 73 seconds across

formulations. Viscosity was adjusted if necessary, by either adding more solvent or by allowing

the solution to evaporate under a fume hood.

PU-ADO aligned nanofibrous scaffolds were produced by electrospinning for 8 hours onto a

1150 rpm rotating mandrel, with an applied potential difference of 17kV at an infuse rate of 0.5

mL/hr. The scaffold was cured in an Isotemp Vacuum Oven (Model 281A, Fisher Scientific)

58

overnight at 45°C. Nanofibers were imaged using scanning electron microscopy and structural

alignment was quantified using Fast Fourier Transformation analysis on ImageJ software197.

Scaffold sheets were cut into 3mm x 50mm strips at a 30° angle relative to nanofiber direction,

and sterilized by gamma irradiation (2.5 Mrad) (Supplemental Figure S2).

Irradiated scaffold strips were soaked in 10 µg/mL fibronectin (f1141, Sigma-Aldrich) in sterile

phosphate buffered saline without Ca2+ or Mg2+ (PBS-/-; Multicell) overnight at 37.5°C. Two

fibronectin-coated strips with opposing fibre angle orientation were stacked and rolled around

TYGON 3350 silicone tubing (4 mm outer diameter, Saint-Gobain) to create 6-layered

concentric constructs. Eight constructs were secured by pins within a spinning bioreactor flask,

and seeded with 1 x 106 primary OAF cells per individual construct in AF-optimized complete

media (DMEM supplemented with 1% ITS [315-080-ZL, Wisent], 1% sodium pyruvate [11360-

070, Gibco], dexamethasone [final concentration 0.0002 mM; D2915, Sigma], L-proline [final

concentration 0.08 mg/mL; P5607, Sigma], 1% Antibiotic-Antimycotic solution [450-115-EL,

Multicell], 10% FBS). After 2 days, the full volume was taken out of the bioreactor in order to

remove all unattached cells, and replaced with fresh complete media and ascorbic acid (final

concentration 100 µg/mL; Sigma-Aldrich). 2/3 of the conditioned media volume was replaced

with complete media and fresh ascorbic acid every 2 to 3 days thereafter. Tissues were

maintained in dynamic culture at 35 rpm for up to 2 weeks (Figure 3.1A).

3.2.3 Formation of hyaline cartilage tissue in 3D culture

Full-thickness cartilage was harvested from 6 to 9-month-old bovine metacarpal-phalangeal

joints within 4 hours of death and placed in Ham’s F12 medium (Multicell). Tissues were

59

digested in 0.25% protease for 1 hour at 37°C, followed by 0.1% collagenase A overnight at

37°C. Cells were filtered through a 100 μm cell strainer followed by a 40 μm cell strainer,

pelleted and washed 3 times in media before counting with trypan blue and a hemocytometer.

Cells were placed in cryomedia (50% DMEM, 40% FBS, 10% DMSO) at 5 x 106 cells per 2 mL

cryotube and stored at -80°C until use. 3, 5, or 7 days before initiation of co-culture with OAF

tissues (Day 0), frozen chondrocytes were thawed, washed immediately in Ham’s F12 media,

and re-counted. 1 x 106 viable cells were seeded on type II collagen-coated Millicell culture plate

polytetrafluoroethylene membrane inserts (0.4 µm pore size, 12 mm diameter; Millipore) in

Ham’s F12 supplemented with 10% FBS. After 2 days, the media was replaced and

supplemented with ascorbic acid (100 µg/mL final concentration). Ham’s F12 media with

ascorbic acid was replaced every 2 to 3 days until transfer into the experimental conditions at 3,

5, or 7 days (Figure 3.1B). Experimental conditions consisted of either cartilage co-cultured with

OAF, or cartilage grown alone. In co-culture experiments, only DMEM complete media was

used. In cartilage-alone experiments, tissues were grown in either DMEM complete media or in

control Ham’s F12 media. All experiments were supplemented with 10% FBS.

3.2.4 Fluorescent labelling of primary chondrocytes

Prior to seeding, select cultures of freshly isolated chondrocytes were incubated with

carboxyfluorescein diacetate (CFDA; Vybrant CFDA SE Cell Tracer Kit V12883, Invitrogen) as

described previously198. A 2.5 μM dye solution was prepared by diluting CFDA stock solution in

PBS-/-. Chondrocytes were gently resuspended in 2.5 μM CFDA at a density of 1 x 106 cells/mL

and incubated for 10 minutes at 37°C. Cells were pelleted and washed twice in Ham’s F12 for 30

minutes at 37°C. Examination of cell aliquots under fluorescent microscopy (Olympus 1X81

60

microscope and OptiGrid imaging) confirmed that >95% of cells were labelled (Supplemental

Figure S3). Labelled cells were resuspended in Ham’s F12 supplemented with 10% FBS and

seeded as described above (Figure 3.1B). OAF-CEP constructs assembled with CFDA-labelled

chondrocytes were imaged with differential interference contrast (DIC) and fluorescent

microscopy using an Olympus 1X81 microscope and OptiGrid imaging device.

3.2.5 Co-culture of in vitro-formed OAF and cartilage tissues

Medium was aspirated from 3, 5, or 7-day old cartilage tissues in their membrane insert holders

and transferred into sterile 6-well plates. 2-week old OAF tissues with the inner TYGON tubing

were placed on the cartilage within each cartilage-containing membrane holder, ensuring

physical contact between the tissues. 5 mL of DMEM complete media (containing ITS, sodium

pyruvate, dexamethasone, L-proline, antibiotics, 10% FBS) with 100 µg/mL ascorbic acid was

pipetted in the plate well, around the outside of each holder. 180 μL of complete media with

ascorbic acid were then placed inside of each holder, which ensured submersion of the OAF

without causing it to float. After allowing the OAF to attach to the cartilage for 48 hours, media

in each 6-well plate was topped up to a final volume of 10 mL. Half of the conditioned media

volume was replaced with fresh complete media and ascorbic acid every 2-3 days. Co-cultured

tissues were harvested at various times up to 4 weeks after initiation of static co-culture (Figure

3.1C). See Appendix (Section 5.1) for method development (note that the described method of

OAF-CEP assembly in this section was the second of two methods that are reported on in

Section 5).

61

62

Figure 3.1. Schematics showing cell isolation and tissue culture methods. (A) Primary

bovine OAF cells were isolated and seeded on angle-ply PU-ADO nanofibrous constructs,

and maintained in dynamic culture for 2 weeks. (B) Articular chondrocytes were cultured

on type II collagen-coated membrane inserts in a 24-well plate. (C) 2-week OAF and 3-day

cartilage tissues were co-cultured in a 6-well plate. Results comparing OAF-CEP constructs

containing 3, 5, or 7-day old developed cartilage informed the decision to perform all

further experiments with 3-day cultured chondrocytes (due to the formation of a

paucicellular layer in 5 and 7-day old cartilage, described in Results, Section 3.3.1).

3.2.6 Determination of mechanical loading magnitude

When the IVD is compressed the NP disperses the forces radially to the AF199. To model this

effect, various compressive forces were applied to an OAF construct containing a central agarose

plug. The radial expansion of the OAF was determined to identify the compressive parameters to

be used in the mechanical stimulation experiments. Three batches of 3% agarose were prepared

separately. 2 plugs from each batch were tested in 6 individual OAF samples. Each agarose plug

(5 mm height x 4 mm diameter) was placed in the centre of a circularized multi-lamellated OAF

tissue formed after 2 weeks of culture. To acquire photographs of the underside of OAF samples

during compression, a custom camera stand was built by hot-gluing 2 pieces of plexiglass (height

= 26 cm) to the short edges of a standard single-well plate lid. The stand was placed under the

indenter of an Instron 4301 mechanical tester machine. An Olympus OM-D E-M10 II mirrorless

camera with a 35 mm fixed focal length lens was placed beneath the stand at a fixed position

throughout the imaging acquisition process, and operated by remote control. The height of the

camera stand allowed for proper lens focusing. The OAF construct and agarose plug were placed

63

atop the transparent plastic surface and imaged (Figures 3.2A, B). Bottom view photographs

were taken during rest and under varying compressive forces in order to determine if each load

was sufficient to cause OAF radial distension during compression of the agarose plug. Agarose

plug samples were also tested alone to determine their point of failure (Figure 3.2C). Lateral

view photographs were taken concurrently in order to determine the point at which the indenter

went too far and began to compress the OAF (Figure 3.2D). ImageJ software was used to

measure 10 individual diameters per sample at fixed 18° angle increments in (Figures 3.2E, F).

Due to the irregularity of tissue thickness within a single OAF construct, these measurements

were used to determine the percent diameter change from rest. Findings from previous

studies33,34,39,40,44,51-53 (Section 2.5, Table 1), in combination with the presently described

determination of loading parameters, were used to establish the experimental loading schedule.

64

Figure 3.2. Determination of compressive load parameters for OAF + agarose plug. (A)

The indenter of the Instron 4301 was positioned over the top of the agarose plug, while the

camera was maintained in a fixed position and operated by remote control. (B) Image

acquired from camera set-up. (C) Agarose plug samples were tested under compression

and photographed alone to determine their point of failure. (D) Lateral photographs were

taken of OAF + agarose plug at each compressive force. (E) Diameter measurements of

65

OAF + plug at rest and (F) under compression using ImageJ. Diameters were measured at

fixed 18° angle increments (yellow arrows, red boxed data) to obtain a total of 10 length

measurements for each sample (blue boxed data).

3.2.7 Dynamic mechanical loading of OAF-CEP constructs

During the third week of a 4-week co-culture period, OAF-CEP constructs were subjected to

dynamic compressive stimulation 3 times (Monday, Wednesday, Friday) using parameters

determined in 3.2.6. The TYGON tubing was removed from 2-week co-cultured OAF-CEP

constructs under sterile conditions. A 3% agarose plug (5 mm height x 4 mm diameter) was

carefully inserted into the centre of each OAF-CEP construct using thin tweezers. Four

constructs contained in their plastic insert holders and conditioned media were placed in the

outside corner wells of a 24-well plate, which was covered with a custom-designed jig consisting

of titanium alloy loading platens and a Teflon plate cover (Supplemental Figure S4). The jig

indenters (diameter = 8.8 mm) were lowered onto the agarose plug and compressive forces were

applied by a Mach-1™ mechanical tester (Biomomentum) (Figure 3.3A).

An axial compression force of 0.4 N per sample was applied at 1 Hz for 1800 cycles (Figure

3.3B). Immediately after mechanical stimulation, samples were returned to their culture plates

and replenished with complete media. Control constructs did not receive mechanical stimulation

but otherwise were treated in the same manner. All constructs were maintained in culture for

another week without stimulation.

66

Figure 3.3. (A) Custom-designed jig loaded in a Mach-1™ mechanical tester. 1 - Load cell;

2 - loading platens; 3 - custom culture plate top; 4 - standard 24-well culture plate

containing OAF-CEP samples. (B) Schematic diagram illustrates the OAF-CEP construct

with an agarose plug in a 24-well plate well during dynamic compression.

3.2.8 Pull-apart testing of OAF-CEP constructs

To determine failure stress, axial tensile strength testing was conducted using an Instron 4301

mechanical tester. Samples were kept hydrated in conditioned media until immediately before

tensile testing. Each sample was prepared at room temperature prior to each individual tensile

test. Excess media was blotted away from each tissue in order to ensure proper adhesion with

glue to the jig. The top and bottom of the OAF-CEP construct was fixed to the machine jig with

Krazy Glue gel. The glue was allowed to cure for 5 minutes under a preload of ~0.1 N (Figure

3.4A). Tensile strain was applied immediately after the 5-minute preload period, at 1 mm per

minute with a 50 N capacity load cell (full scale = 2.5 N) until complete separation of the OAF

and cartilage tissues occurred (Figure 3.4B). Tensile testing for each sample was completed

within approximately 10 minutes of the beginning of sample preparation. Force-displacement

67

graphs were generated and the maximum force value in Newtons was calculated. To obtain

normalized kPa stress values, the contact surface area of each detached OAF was measured.

OAF tissue contact surface area was painted with black ink to allow for visualization. Care was

taken to ensure that only the apparent contact area was marked. Inked samples were

photographed with an Olympus OM-D E-M10 II camera, and the surface area was measured by

manually tracing the inked area using ImageJ software (Figures 3.4C, D). Each sample was

traced 3 separate times to obtain an average surface area value. Samples that did not fully

separate during tensile testing were saved for histology in order to examine the pattern of early

failure at the interface. These samples were not included in the calculation of tensile strength.

68

Figure 3.4. (A, B) Pull-apart failure testing of OAF-CEP constructs with an Instron 4301

Tensile Tester until complete separation of the tissues. (C) After tensile testing, the surface

area of OAF tissues pulled apart from in vitro cartilage were marked with black ink and

photographed next to a ruler. Each image was calibrated in ImageJ using the distance in

pixels corresponding to 10 mm (yellow line). (D) After calibration, the outer and inner

circumferences were manually traced using the polygon selection tool. The value of the

inner area was subtracted from the value of the outer area to obtain a final measurement

for the blackened area. Yellow arrowhead indicates the unmarked side wall of the OAF

69

tissue that was not in contact with cartilage. Red arrowhead indicates shadow produced by

camera flash.

3.2.9 Histological characterization of native and in vitro OAF-CEP interface

Harvested in vitro tissues were fixed in 10% buffered formalin for 1 hour and immersed in 30%

sucrose overnight. Native bovine caudal IVDs with attached CEP and VB were fixed in 10%

buffered formalin for 48 hours, decalcified in 0.5 M EDTA for 4 weeks, and immersed in 30%

sucrose for 48 hours. In vitro and native samples were embedded in optimal cutting temperature

compound embedding medium and kept frozen at -30°C until cryostat sectioning. 7 µm-thick

frozen sections were mounted on cryotape200 (Cryofilm type 2C, SECTION-LAB Co. Ltd.) and

subsequently adhered onto glass slides using UV-activated optical glue (Norland Optical

Adhesive 61, Norland Products Inc.) as previously described201. Glued sections were allowed to

cure under a biosafety cabinet UV lamp for 1 hour (Supplemental Figure S5). Adhered sections

were stained with hematoxylin and eosin (H&E) or toluidine blue, and imaged under light

microscopy.

3.2.10 Immunohistochemical characterization of the OAF-CEP interface model

For collagen type I and type II immunostaining, cryotape-adhered sections were subjected to

antigen retrieval by sequential enzyme digestion as follows: Pepsin (2.5 mg/mL in TBS pH 2.0;

P7012, Sigma) for 10 minutes at room temperature followed by two washes in PBS-/- for 10

minutes; Trypsin (2.5 mg/mL in TBS; T7409, Sigma) for 30 minutes at room temperature

followed by one wash in PBS-/- for 5 minutes; Hyaluronidase (25 mg/mL in PBS-/-; H3506,

Sigma) for 30 minutes at 37°C followed by one wash in PBS-/- for 5 minutes. For aggrecan

70

immunostaining, sections underwent antigen retrieval with hyaluronidase only for 30 minutes at

37°C followed by one wash in PBS-/- for 5 minutes. Sections were blocked with 20% goat serum

and 0.1% Triton X-100 in PBS-/- for 1 hour at room temperature, washed for 5 minutes in PBS-

/-, and incubated overnight at 4°C with primary antibodies reactive with collagen type I (1:75

dilution; ab34710, Abcam), collagen type II (1:75 dilution; MAB8887, Millipore), and aggrecan

(1:750 dilution; AHP002, Invitrogen). Rabbit and mouse IgG were used as negative controls.

Antibody diluent consisted of 10% goat serum and 0.1% Triton X-100 in PBS-/-. After overnight

incubation, sections were washed in PBS-/- (once for 5 minutes, twice for 15 minutes) and

incubated for 1 hour at room temperature with fluorescent labelled secondary antibodies

suspended in antibody diluent. Goat anti-mouse IgG (1:1000 dilution; A11032, Invitrogen) or

goat anti-rabbit IgG (1:1000 dilution; A11008, Invitrogen) secondary antibodies were used

depending on the source of the primary antibodies. Sections were washed three times in PBS -/-

(once for 5 minutes, twice for 15 minutes) and incubated with DAPI (1:10 000 in PBS-/-) for 5

minutes to stain nuclei. Slides were washed three times in PBS-/- for 5 minutes, and coverslipped

with PermaFluor mounting medium (TA-030-FM, Thermo Scientific). Slides were protected

from light and allowed to dry for 48 hours before visualization under a Zeiss Axioplan

fluorescent microscope.

3.2.11 Quantification of DNA, collagen, and proteoglycan contents

In vitro OAF-CEP tissues were removed from the insert holder using a scalpel. To isolate the

tissue interface region, 1 mm of the OAF tissue in contact with the cartilage layer was cut with a

scalpel under a dissecting microscope (Figure 3.5A). 3-mm and 7-mm biopsy punches were used

to cut the cartilage inside and around the OAF tissue, respectively (Figure 3.5B). In vitro

71

cartilage and OAF tissues cultured alone (i.e. not in co-culture) for the same time period were cut

in the same manner and used as control samples. Excess tissue from each sample was discarded,

and the remaining interfacial tissues were digested in 40 µg/mL papain (P3125, Sigma) in

digestion buffer (35 mM ammonium acetate, 1 mM EDTA, 2 mM DTT pH 6.2) for 48 hours at

65°C. DNA content was determined by a fluorometric assay (355 nm excitation, 460 nm

emission) using Hoechst 33258 dye as previously described202. Calf thymus DNA (D1501,

Sigma) was used to generate a standard curve. Sulphated glycosaminoglycan (GAG) content was

determined using a dimethylmethylene blue binding assay and spectrophotometry (λ = 525 nm).

Bovine trachea chondroitin sulphate A (C8529, Sigma) was used to generate a standard curve.

To determine hydroxyproline content, papain-digested samples were hydrolyzed in an equal

volume of 6 N hydrochloric acid at 110°C for 18 hours, followed by neutralization with the same

volume of 5.7 N sodium hydroxide. Hydroxyproline (OH-Pro) content was quantified using the

chloramine-T/Ehrlich’s reagent assay and spectrophotometry (λ = 560 nm). Cis-4-Hydroxy-L-

proline (219959, Aldrich) was used to generate a standard curve. Hydroxyproline and GAG

values were normalized to DNA.

72

Figure 3.5. Schematic diagram illustrating how in vitro OAF-CEP interface was prepared

for biochemical analysis (red lines). (A) 1 mm of OAF was cut above the cartilage layer. (B)

3-mm and 7-mm biopsy punches were used to cut the cartilage surrounding the interface.

Excess tissue was discarded. Control tissues (cartilage or OAF alone) were prepared in the

same manner.

3.2.12 Statistical analysis

Each experiment was performed in triplicate and was independently repeated at least 3 times

using cells from different animals. Tensile testing was performed on at least N = 10 samples for

each experimental condition. Results are expressed as mean ± SEM. Differences between two

groups were determined using an unpaired T-test and considered to be statistically significant at

p < 0.05.

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3.3 Results

3.3.1 Co-culture of in vitro OAF on cartilage tissue generated an integrated tissue interface

Six-layer angle-ply OAF tissues remained in place (i.e. did not dislodge from the cartilage layer)

upon the addition of media on day 2 of co-culture, indicating that some degree of attachment

between the two tissues had occurred by this time. The two tissues were integrated as early as 1

week of co-culture (Figure 3.6B). Histology confirmed the presence of OAF tissue composed of

elongated cells aligned parallel to the PU-ADO scaffold (Figures 3.6C, D), which was confirmed

by SEM to contained aligned nanofibers (Supplemental Figure S6). Interestingly, the OAF cells

were aligned parallel to the cartilage layer at the interface (Figure 3.6E), which was similar to the

alignment seen in the native fetal OAF-CEP interface (Figure 3.6G). Scaffold layer ends of the

model were oriented downwards perpendicular to the cartilage layer (Figure 3.6E, F), similar to

what occurs in the native interface. OAF tissue alone did not stain for proteoglycan (Figure

3.6D), however cartilage tissue rich in toluidine blue-positive ECM was present (Figure 3.6F).

2-week developed OAF tissue integrated with cartilage tissues that were pre-grown for 3, 5, and

7 days. At the interface, a paucicellular layer of ECM developed in the superficial aspect of the

cartilage of 5-day old CEP-like tissue (Figure 3.7B). This layer appeared thicker and more

distinct in constructs formed with 7-day pre-grown cartilage (Figure 3.7C). 7-day in vitro tissues

forming cartilage alone (i.e. not co-cultured) did not have this layer prior to placement in co-

culture. Similar to 3-day old cartilage alone (Figures 3.7D, E), 7-day old cartilage alone

contained evenly distributed cells throughout the tissue (Figures 3.7F, G). 3-day, 5-day and 7-

day old cartilage grown for an additional 2 weeks in AF-optimized DMEM complete media

appeared to have accumulated a thicker top layer of paucicellular ECM than did the cartilage

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grown in control Ham’s F12 media (Figures 3.7H-M). The paucicellular layer that was seen at

the interface between OAF and older in vitro cartilage tissues was therefore not a pre-existing

characteristic of the pre-grown cartilage, and appeared to develop from prolonged culture with

AF-optimized DMEM media. These findings informed the decision to perform all further

experiments with chondrocytes grown for 3 days, in order to avoid the appearance of an

interfacial ECM layer with sparse cellularity.

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Figure 3.6. (A) Lateral and top macroscopic views of 2-week co-cultured OAF-CEP tissue

constructs. Constructs are shown as they appeared both before and after removal from

plastic membrane insert holders. (B) 2-week old OAF and 3-day old cartilage integrated by

1 week of co-culture. Arrow indicates the outermost edge of the in vitro OAF. (C) OAF

tissue formed between layers of PU-ADO scaffold in the in vitro OAF region, and (D) did

not stain for proteoglycan. (E, F) In vitro-formed interface of OAF-CEP co-cultured for 2

weeks. (G) Native fetal bovine OAF-CEP interface. Yellow arrowheads indicate directional

orientation of OAF cells along PU-ADO scaffold layers (C, D), parallel to the cartilage

layer at the in vitro interface (E), and parallel to the CEP of native fetal interface (G). # =

OAF; * = cartilage; S = PU-ADO scaffold. H&E (B, C, E, G) and toluidine blue (D, F)

stained. Light microscopy. Scale bars represent 100 µm.

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Figure 3.7. Histological appearance of in vitro formed interface between 2-week old OAF

and (A) 3-day old , (B) 5-day old, and (C) 7-day old cartilage tissues co-cultured for 1 week.

Histological appearance of 3-day (D, E) and 7-day (F, G) in vitro-formed cartilage prior to

placement in co-culture. 3-day (H, I), 5-day (J, K), and 7-day (L, M) pre-grown cartilage

grown alone for an additional 2 weeks in control Ham’s F12 media + 10% FBS (H, J, I) or

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complete DMEM AF-optimized media (I, K, M). S = PU-ADO scaffold; # = OAF; * =

cartilage; arrowheads = paucicellular layer of ECM. H&E (A-C, D, F, H-M) and toluidine

blue stained (E, G). Light microscopy. Scale bars represent 100 µm. NB: OAF-CEP

constructs (A-C) were assembled without physical support of tubing, which led to

convolution of scaffold layers and a lack of perpendicular insertion into the cartilage layer.

3.3.2 Chondrocytes remained localized to the cartilage layer beneath the interface

To determine if the integration was related to cells migrating in both directions (i.e. chondrocytes

migrating into the AF tissues or vice-versa), the chondrocytes were labelled with green

fluorescent CFDA dye. At two weeks, when there was sufficient integration to handle the

constructs, examination of histological sections under fluorescent microscopy showed that the

fluorescent labelled cells were localized to the cartilage layer and did not appear to migrate up

into the OAF (Figure 3.8). The morphological features of spindled cells resembling OAF cells

suggested that these cells were growing down into the cartilage layer to form the interface.

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Figure 3.8. Representative histological sections of the OAF-CEP interface at 2 weeks of co-

culture showing integration of the two tissues by (A) light microscopy and (B) green stained

chondrocytes as visualized by fluorescent microscopy. CFDA-labelled chondrocytes

remained localized to the cartilage layer of the OAF-CEP construct. # = OAF; * =

cartilage; S = PU-ADO scaffold. (A) H&E; (B) DIC/fluorescent microscopy. Scale bars

represent 100 µm. NB: The green signal that is visible within the OAF scaffold was due to

autofluorescence and not to the presence of CFDA in this region.

3.3.3 Immunohistochemical characterization of the bioengineered OAF-CEP interface

Collagen type I was localized to the OAF (green arrows; Figure 3.9C), and collagen type II was

localized to the cartilage as expected (yellow arrows, Figure 3.9C), similar to the localization in

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native interface where collagen type I was localized to the OAF region, and collagen type II to

the CEP region (Figure 3.9J). Compared to 2-week constructs, 4-week constructs showed more

pronounced collagen type I staining in the OAF above the interface, as well as distinct overlap

between the two collagen types indicated by the yellow staining (Figure 3.9 G). This overlap was

also characteristic of the native interface, where yellow staining was also visible (Figure 3.9J).

The main difference between the in vitro model and the native tissue was the diffuse collagen

type II staining in the OAF region above the in vitro cartilage-like layer (white arrows; Figures

3.9C, G). This collagen type II-positive region corresponded with the histological images to an

intermediate region containing elongated cells, suggesting that this tissue was AF-like (white

arrows; Figures 3.9B, F). The ECM of this zone appeared less dense (i.e. showed lighter eosin

staining) than the above OAF tissue surrounding the scaffold, suggesting that it was less

developed and arose during the co-culture period. The denser, more heavily stained OAF tissue

around the scaffold only stained for collagen type I as expected (green arrows; Figures 3.9C, G)

and was not detected in any other OAF regions distant from the interface region (Supplemental

Figure S7).

Aggrecan was localized to the cartilage layer and was only weakly if at all present in the

interfacial OAF region of the 2-week in vitro interface (Figure 3.9D). For interfaces developed

over 4-weeks in vitro, the presence of aggrecan extended past the cartilage layer and up into the

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OAF tissue (Figure 3.9H). The presence of aggrecan in the OAF region at 4 weeks resembled the

presence of aggrecan in the native interface (Figure 3.9K).

Figure 3.9. Representative histological appearance of the in vitro interface following

staining by H&E (A, B, E, F) and corresponding immunostaining for collagen type I

(green) and type II (red) (C, G) and aggrecan (D, H) of 2-week (A-D), and 4-week (E-H)

constructs. Representative histological staining by H&E (I) and corresponding

immunostaining for collagen types I and II (J), and aggrecan (K) of the native young

bovine OAF-CEP interface. Nuclei were stained with DAPI (blue). Green arrows = in vitro

OAF tissue; White arrows = intermediate in vitro OAF-like tissue; Yellow arrows = in vitro

cartilage tissue; White arrowheads = presence of aggrecan in in vitro OAF region. # = in

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vitro OAF; * = in vitro cartilage; S = PU-ADO scaffold. Light microscopy (A, B, E, F, I) and

fluorescent microscopy (C, D, G, H, J, K). Scale bars represent 100 µm.

3.3.4 Apparent mechanical strength of the in vitro OAF-CEP interface increased over time

Pull-apart failure testing revealed that 2-week cultured OAF-CEP constructs a possessed

measurable interfacial mechanical strength of 11.9 ± 1.8 kPa, which was significantly increased

to 21.0 ± 2.2 kPa at 4-weeks of co-culture (p = 0.002) (Figure 3.10A). Histological examination

revealed that the failure occurred between the OAF and cartilage layer in all constructs (Figures

3.10B, C). In addition to this predominating failure pattern, mid-substance failure was also

observed whereby the failure occurred either within the cartilage (Figure 3.10D) or the within the

OAF (Figure 3.10E) rather than between the two tissues at the interface. These mid-substance

modes of failure were rare and were only each observed once among the 6 samples examined.

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Figure 3.10. (A) Apparent stress at failure of 2-week compared to 4-week cultured OAF-

CEP constructs. Each plotted point represents the apparent stress at failure of a single

sample. 2-week OAF-CEP N = 16 samples, 4-week OAF-CEP N = 14 samples, from n = 5

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sets. Data are expressed as mean ± SEM. **p = 0.002. Failure during tensile testing

occurred between the OAF and cartilage tissues of 2-week (B) and 4-week (C) co-cultured

constructs (N = 6 samples examined). Rare mid-substance failures occurred directly in

either the cartilage (D) or the OAF (E) versus the mid-section of the interface (N = 1 for

each case). In these cases, the cartilage remained attached to the OAF in some regions (D),

or OAF remained attached to the cartilage in some regions (E). Arrowheads indicate

regions of failure between tissues. # = OAF; * = cartilage. Toluidine blue. Light

microscopy. Scale bars represent 100 µm.

3.3.5 No significant difference in OH-Pro or GAG content between 2 and 4 weeks of co-

culture

DNA content did not significantly change between 2 and 4-week OAF-CEP constructs (p =

0.384). DNA content also did not significantly change between control OAF tissues cultured

alone for 2 or 4 weeks (p = 0.400), or between control cartilage tissues cultured alone for 2 or 4

weeks (p = 0.498, Figure 3.11A). There was no significant difference in GAG content relative to

DNA (p = 0.252, Figure 3.11B), or OH-Pro content relative to DNA (p = 0.230, Figure 3.11C)

between 2 and 4 weeks of co-culture.

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Figure 3.11. (A) Biochemical analysis of DNA content of 2-week and 4-week OAF-CEP

interface tissues (p = 0.384), OAF alone (p = 0.400), and cartilage alone (p = 0.498). (B)

GAG normalized to DNA of 2-week and 4-week OAF-CEP interface tissues (p = 0.252). (C)

OH-Pro content relative to DNA of 2-week and 4-week OAF-CEP interface tissues (p =

0.230). Data pooled from n = 4 biological replicates. Total sample size for 2-week OAF-CEP

sample size N = 10; 4-week OAF-CEP sample size N = 11. Results expressed as mean ±

SEM.

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3.3.6 A compressive force of 0.4 N resulted <10% radial strain in OAF tissues

Compression at 0.6 N consistently led to failure of the 3% agarose plug or complete

displacement from its position beneath the indenter (Figure 3.12A). It also led to crushing of the

OAF + agarose plug samples (Figure 13.12B). For this reason, forces of no more than 0.5 N were

tested on OAF + agarose plug samples. A compressive force of 0.4 N was selected because it

enabled a mechanically stable agarose plug. The condition of 0.4 N yielded an average OAF

tissue distension not exceeding 10% (average percent diameter change from 6 measured samples

= 4%, SD = 3.56) (Figures 3.12C, D).

A review of the literature suggested a maximum of 10% strain as appropriate to induce anabolic

effects in in vitro IVD and cartilage tissues33,34,39,40,44,51-53. Most mechanical stimulation studies

have utilized a load frequency of 1 Hz (Section 2.5, Table 1) which corresponds to a

physiological locomotion frequency51. Kasra et al. described a loading treatment of 30 minutes a

day for 3 days for timely experiment completion and for minimization of contamination risk36.

As such, the following loading schedule was established: 0.4 N per construct, 1 Hz sinusoidal

compression for 1800 cycles, once every other day (Monday, Wednesday, Friday) for 1 week.

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Figure 3.12. (A) Agarose plug testing under the indenter of the Instron 4301. (B)

Compression of OAF + agarose under 0.6 N crushed the tissues. (C) Lateral and bottom

view photographs were taken concurrently of 2-week old OAF tissue constructs containing

a central agarose plug at rest and under varying compressive forces. Lateral view photos

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(upper panel) are not to scale. Red arrows = OAF diameter bottom view measurement for

each compressive force in ImageJ (example of only one 90° angle measurement for N = 1

sample is shown). d = Measured bottom diameter in unit image pixels (px). ∆d = Unit

change in diameter compared to the resting diameter. % change = Percentage of diameter

increase in comparison to the resting diameter. (D) Mean distribution of diameter

measurements taken from N = 6 OAF samples under 0.4 N load. Each plotted point

represents one of ten diameter measurements for each sample.

3.3.7 Dynamic mechanical stimulation did not significantly increase the apparent tensile

strength, collagen content, or proteoglycan content in OAF-CEP interface constructs

4-week co-cultured OAF-CEP constructs that received dynamic loading every other day during

the third week of a 4-week culture period did not significantly differ in apparent tensile strength

compared to 4-week cultured OAF-CEP constructs that received no loading (p = 0.111) (Figure

3.13A).

There were no obvious differences in the presence of proteoglycan as indicated by toluidine blue

staining between 2 and 4-week constructs (Figures 3.13B, E). Immunostaining of 4-week old

loaded and unloaded OAF-CEP constructs did not show a change in collagen type I, collagen

type II (Figures 3.13C, F) or aggrecan distribution (Figures 3.13D, G) at the interface. 30

minutes of 1 Hz, 0.4 N dynamic compressive loading every other day for 1 week led to the

destruction of the interface in n = 2 out of the 3 constructs examined (Figures 3.13H-J).

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There was no significant difference in the quantification of DNA content (p = 0.329), GAG

relative to DNA (p = 0.347), or OH-Pro relative to DNA (p = 0.179) when comparing unloaded

and loaded 4-week OAF-CEP constructs (Figures 3.13K-M).

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Figure 3.13. (A) Apparent stress at failure between unloaded control and loaded OAF-CEP

constructs. Each plotted point represents the apparent stress at failure of a single sample.

Data pooled from n = 4 biological replicates. 4-week unloaded control total sample size N =

10; 4-week Dynamic Loaded total sample size N=11. Data expressed as mean ± SEM. 4-

week unloaded control (B-D) and loaded OAF-CEP constructs (E-G). Histology (B, E)

corresponds to the immunostaining for collagen type I (green), collagen type II (red) (C, F),

and aggrecan (red) (D, G). Dynamic loading led to destruction of the interface in n = 2 out

of 3 samples examined as shown by histological staining (H, I) and immunostaining (J). # =

OAF; * = cartilage; S = PU-ADO scaffold. Toluidine blue stained, light microscopy (B, E,

H, I). Fluorescent microscopy (C, D, F, G, J). Scale bars represent 100 µm. NB: The

separation at the interface (C) was due to tissue processing during the immunostaining. (K)

DNA content, (L) GAG/DNA content, and (M) OH-Pro/DNA content between unloaded

control and loaded OAF-CEP interface tissues. Data pooled from n = 4 biological

replicates. 4-week unloaded control total sample size N = 11; 4-week Dynamic Loaded total

sample size N=10. Data expressed as mean ± SEM.

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CHAPTER 4: DISCUSSION

4.1 Summary

The objective of the current thesis was to develop a structurally integrated, biomimetic OAF-

CEP interface model, and to determine if the mechanical strength at the interface could be

enhanced by cyclic compressive loading. In order to accomplish this, in vitro OAF-like tissues

were generated using AF cells seeded into multilamellar angle-ply nanofibrous PU-ADO

scaffolds, which were placed in static co-culture on top of in vitro formed cartilage tissues. This

resulted in integrated tissue constructs containing an interface which had mechanical strength

that increased significantly between 2 and 4 weeks.

The morphology of the cells and ECM at the interface had features that modeled the native OAF-

CEP. More specifically, the elongated OAF cells were aligned parallel to the cartilage layer at

the interface, similar to what was observed in the native fetal tissue. The localization patterns of

collagen types I and II resembled that of native OAF-CEP, and both collagen types were

intermixed at the interfacial region. Unlike native tissue, diffuse collagen type II was present in

the in vitro interfacial OAF and had a histological appearance of loose AF-like tissue; the denser

OAF tissue distant to this zone only stained for collagen type I. Aggrecan was present at the

interface and extended into the OAF region above the interface by 4 weeks of co-culture, similar

to its diffuse distribution in native interfacial OAF. OAF cells appeared to grow into the cartilage

layer at the interface, while chondrocytes demonstrated minimal, if any, migration up into the

OAF region by remaining localized to the cartilage over the 4 weeks of the culture study. Failure

of the interface during tensile testing predominately occurred between the OAF and CEP-like

tissues, with rare occurrences of mid-substance failure in either of the original two tissue states.

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Although it was possible to generate an in vitro model of the OAF-Cartilage interface, the cyclic

compressive loading studies performed to date did not demonstrate an effect on the strength,

organization or composition of the interface in the time period studied.

4.2 Discussion and Study Limitations

4.2.1 Use of in vitro-formed OAF tissue on aligned angle-ply PU-ADO nanofibrous

scaffolds and hyaline-like cartilage are suitable to generate an OAF-CEP model

During development in vivo, the organization of elongated fibroblast-like AF cells into aligned,

lattice-like concentric sheets precedes the appearance of collagen70,71. The polarized orientation

of AF cells presumably promotes the aligned deposition of collagen fibres to form the fully

developed AF tissue91. The current in vitro-generated OAF model mimicked this process,

whereby seeded AF cells first attached and oriented with aligned, angle-ply, fibronectin-coated

nanofibers, which then allowed for subsequent directed accumulation of aligned ECM174,175.

Alignment of electrospun polymer nanofibers has been shown to promote aligned AF cell

orientation, whereas randomly oriented fibres led to irregular cell orientation and less collagen

type I and proteoglycan production29. The present study confirmed that collagen type I

accumulation between scaffold layers mimicked the interlamellar distribution of collagen type I

in native OAF tissue (Supplemental Figure S7). This suggests that the in vitro-formed OAF

tissue was appropriate for use in this model.

In vitro CEP-like tissues were not formed by chondrocytes of the very thin native bovine CEP,

but rather from the more abundant supply of articular chondrocytes from metacarpal-phalangeal

joints. This was necessary in order to obtain a sufficient number of cells to perform all the

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experiments. The IVD CEP and the articulating surfaces of synovial joints are similarly

comprised of hyaline cartilage. The current study demonstrated that the in vitro cartilage formed

by articular chondrocytes was rich in collagen type II, similar to the native CEP63. The in vitro-

formed cartilage contained aggrecan, which has also previously been shown to be present in the

native CEP (Rok 2017, unpublished), and in the current study (Figure 3.9). Furthermore, native

CEP tissue was shown to increase matrix proteoglycan accumulation of in vitro-formed NP

tissue when placed in co-culture, and this anabolic effect was mimicked by in vitro cartilage

formed in 3D culture by articular chondrocytes105. From this, it was suggested that chondrocytes

of both CEP and articulating joints produce similar factors that may promote IVD cell growth

and/or ECM synthesis105. These data suggest that in vitro cartilage formed by articular

chondrocytes was an appropriate tissue to model the CEP in the current study. Incorporation of a

similar in vitro-formed cartilage layer in a triphasic NP-CEP-VB model was previously shown to

have improved shear strength when compared to a biphasic NP-VB construct, adding support to

the rationale for the need to include a cartilaginous component in an IVD-VB interface model203.

A Teflon porous membrane was chosen as the substrate on which to form the in vitro cartilage in

the current model based on previous studies that have used the same method of 3D cartilage

generation196,204. However, other studies have generated in vitro cartilage from articular

chondrocytes in 3D culture on top of calcium polyphosphate (CPP), a hard bone-like

substrate105,198,205. Waldman et al. reported that the ECM of cartilage tissues grown on CPP

contained similar amounts of collagen but more proteoglycans over eight weeks compared to

tissues grown on the more flexible Teflon membrane inserts, suggesting that ECM composition

is dependent on the underlying substrate on which it is grown204,205. The specific effect of

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culturing chondrocytes on a Teflon membrane compared to CPP is unknown from this study,

however substrate stiffness may have various effects on cell behaviour and may play a role. For

example, Schuh et al. found that articular chondrocytes seeded on soft polymer gels had a

rounder shape and showed higher collagen type II and aggrecan gene expression in comparison

to chondrocytes seeded on stiffer gels206. The effect of substrate material properties should be

investigated with respect to the current model in terms of ECM production, composition, and the

ability to integrate with multilamellated OAF tissues, as the focus will eventually shift to

integration with a hard bone substitute (i.e. CPP) as part of a complete implant model.

4.2.2 Requirements for successful co-culture of in vitro OAF and cartilage tissues

To determine the optimal age of pre-grown cartilage for integration with in vitro OAF tissues,

articular chondrocytes were cultured for 3, 5, or 7 days before being placed in co-culture. The

AF-cartilage interface formed more reliably if 3-day old cartilage was used in the OAF-CEP

construction. When placed in co-culture, OAF cells and interfacial chondrocytes of 3-day pre-

grown cartilage organized in a manner similar to what is seen in native fetal interface tissue.

Tissue constructs assembled from 5-day pre-grown cartilage developed a layer of paucicellular

ECM between the OAF and the deeper zone of the cartilage. This layer was even more

pronounced in 7-day old cartilage. Hence, the age of the developed cartilage used in this system

was important, as different outcomes were observed depending on the pre-culture period. The

zone of ECM with sparse cellularity is believed to have developed as a result of the culture

conditions, as cartilage grown alone in complete AF-optimized DMEM media accumulated a

similarly paucicellular ECM layer in the superficial region of the tissue over time (Figure 3.7).

Paucicellular ECM did not accumulate to such an extent in cartilage formed in Ham’s F12 media

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(Supplemental Figure S9), which is the standard media used to form these tissues204.

Immunohistochemical studies showed that this layer was rich in collagen type I (Supplemental

Figure S10), which healthy hyaline cartilage chondrocytes do not normally produce. The native

interface did not exhibit such a layer of paucicellular ECM, and collagen type I was absent from

the CEP except at the OAF-CEP interface where the chondrocytes interfaced with OAF cells

(Figure 3.9)63. In co-culture, this matrix-rich region may prevent integration and cross-talk by

acting as a barrier that separates the OAF cells from the chondrocytes. As there were few

chondrocytes at the interface when using the older cartilage, 3 days was selected as the time at

which cartilage would be combined with OAF tissues. It is not known what mechanism(s) were

involved to restrict the further development of this paucicellular ECM zone once the two tissues

were in co-culture. The data suggests that some type of crosstalk between AF cells and

chondrocytes must have prevented this layer from forming in the case of 3-day old cartilage. The

specific effect of the AF-optimized media formulation on articular chondrocytes has not been

investigated, however the current findings suggest evidence of a media-related agent that might

have an influence on cell-cell contact events. Unlike the AF-optimized complete high-glucose

(4.5 g/L) DMEM media, the Ham’s F12 control media has a lower glucose concentration (1.8

g/L) and is not supplemented with the same agents (ITS, dexamethasone, sodium pyruvate,

proline). The presence of dexamethasone for example may act as a stressor to chondrocytes

which may lead them to undergo autophagy, a process whereby cells digest their own

intracellular components as a self-protective response207. Autophagy in chondrocytes may

suppress the expression of matrix degrading MMPs208. In our model system, dexamethasone may

have induced autophagy thereby disrupting the balance between synthesis and degradation, and

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may have thus contributed to the accumulation of the thick ECM layer. Further study is required

to investigate the mechanism(s) involved in the formation of this paucicellular ECM zone.

An important aspect of the current model design was the perpendicular orientation of the

nanofibrous scaffold fibres towards the cartilage layer. This modeled the portion of OAF

collagen fibers that inserts vertically into the CEP at the native interface63,100. An earlier

construct assembly method prevented this perpendicular scaffold orientation. Macroscopically,

in vitro OAF tissue appeared to be contractile when placed in static culture without the internal

support of the TYGON tubing (Method 1 of co-culture, Appendix Section 6.2; Supplemental

Figure S1A). Microscopically, this led to curled and convoluted scaffold layers within the OAF

ECM (Supplemental Figure S1D). The presence of tubing within the OAF tissues during static

co-culture prevented PU-ADO scaffold convolution and collapse that was seen with Method 1 of

co-culture, as determined histologically (Supplemental Figure S1B, E). Cell-mediated

contraction has been suggested to depend on the stiffness of the scaffold209, which is in

agreement with the current finding that the provision of a stiffer physical support by the tubing

prevented visible contraction of OAF cell-seeded PU-ADO scaffolds. As OAF cells aligned with

the scaffold layers, it may be assumed that deformed scaffold lamellae led to abnormal

orientation of the OAF cells and therefore to irregular collagen fibre orientation at the interface

in this model system. This was not desired as the purpose of the model design was to recreate the

organized structure of the highly linear native OAF fibres. Maintenance of the tubing during co-

culture allowed for the ends of scaffold layers to be oriented perpendicular to the cartilage layer,

similar to the way in which collagen fibres insert into the native CEP63,100.

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In order for the TYGON tubing to be maintained within the OAF tissue during co-culture, it was

also necessary for the cartilage layer to be maintained in its insert holder. In one experiment, 3-

day old in vitro cartilage and its underlying membrane were excised from tension (the holder).

OAF tissues with a central TYGON tube were then placed on top of the excised membrane and

cartilage within a 6-well plate and co-cultured in complete media. After 24 hours, only one of

three OAF constructs formed an attachment with the underlying cartilage (data not shown). The

two OAF tissues that failed to attach to the cartilage contracted and had completely pushed out

their tubing. The tension provided by the insert holder kept the cartilage tissue flat, which

facilitated attachment with OAF tissues.

4.2.3 Characterization of the OAF-CEP interface

The in vitro formed OAF and cartilage tissues resembled their native counterparts with respect to

cell morphology, cell orientation, and ECM characteristics. OAF cells at the interface aligned

parallel to the cartilage layer, which mimicked the orientation of OAF cells adjacent to the native

fetal CEP. Chondrocytes remained localized to the proteoglycan-rich in vitro cartilage layer,

which appeared histologically distinct from the OAF tissue. Similar to the native OAF-CEP

which was characterized both previously63 and in the current study (Figures 3.9I-K),

immunostaining showed that the in vitro cartilage-like tissue was abundant in collagen type II,

and that the OAF-like tissue contained collagen type I. The intermixing of both collagen types I

and II at the interface mimicked what was observed at the native interface. The co-localization of

collagen types I and II has been linked to biomechanical properties and load transfer at the

tendon-bone insertion site. Collagen type I is localized in the aligned fibres of the tendinous

region (similar to the OAF) while collagen types II and I are localized to the fibrocartilage

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adjacent to the bony region of the insertion site210,211. Differences in viscoelastic properties found

in the bony insertion region as compared to the tendinous region suggested that the latter was

more highly resistant to tension211. Therefore, the differences in ECM collagen localization and

organization may be a result of load transfer, whereby tensile stress is transferred

unidirectionally along the aligned collagen type I fibres and distributed to the collagen type II/I-

containing ECM which is less organized therefore able to transfer multidirectional forces to the

underlying bone211. Although the tendon insertion is structurally very different than the OAF-

CEP interface, the overlap of collagen types I and II observed at the transition zone of the in

vitro and in vivo OAF-CEP interface suggests that this model has a composition that would

enable it to appropriately distribute forces in a manner similar to what occurs physiologically.

Collagen type II, normally not present in OAF tissues, was present in the OAF region at the

interface at both 2 and 4 weeks. Histological examination showed that this collage type II-

positive zone contained elongated cells within a loose ECM, suggesting that this tissue was

developing AF-like tissue. A possible reason for this may be that AF cells proliferated and

deposited new ECM at this interfacial region, which may explain why the tissue in this zone

appeared looser than the pre-grown OAF tissue surrounding the scaffold. The influence of the

adjacent cartilage may have induced these AF cells to produce collagen type II by a crosstalk

mechanism. Alternatively, collagen type II synthesized by the chondrocytes may have diffused

into this looser AF-like tissue. In situ hybridization for collagen type II mRNA expression would

be required to determine if AF cell(s) are responsible for the production of collagen type II, in

order to support either of these hypotheses.

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Aggrecan, a proteoglycan that is responsible for the compressive-resistant nature of hyaline

cartilage212, was detected in the in vitro-formed OAF at the interface by 4 weeks of co-culture,

which resembled its appearance in native interfacial OAF tissue. Aggrecan was absent in the

OAF distant from the interface region as expected, as its presence in the native IVD is lowest in

the OAF in comparison to the NP and IAF30. The distribution of aggrecan in the native disc

follows the type of stresses each region experiences; the NP predominantly experiences

compressive stress and therefore contains the highest amount of compressive-resistant aggrecan,

while the OAF experiences radial tensile stress and therefore contains high amounts of tensile-

resistant collagen type I64,213.

Although proteoglycan (aggrecan) remained mostly localized to the cartilage layer at 2 weeks

and appeared in the interfacial OAF tissue by 4 weeks, it should be noted that when biochemical

analysis of the interface region was performed, no difference was shown in the proteoglycan and

collagen contents relative to DNA over time. DNA content also remained unchanged. One

possible explanation for this discrepancy is that the proteoglycan may have diffused from the

cartilage layer into the OAF over time, thereby altering its distribution but not its overall content.

An alternative explanation is that crosstalk with the cartilage layer may have induced

proteoglycan synthesis by OAF cells at a level too low to detect as a difference by the GAG

assay, however this is purely speculative. The specific effect(s) of chondrocyte co-culture on

OAF cell phenotype and matrix production remains an area of study that, to our knowledge, has

not been investigated to date.

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It must be noted that the methods used to prepare tissues for biochemical analysis had several

limitations. Firstly, despite using a standard method of tissue dissection for all samples, it was

difficult to isolate only the interfacial region of each construct. Biopsy punches were used in

order to standardize the amount of cartilage cut away from the constructs. Cutting away excess

cartilage directly from OAF tissue edges (i.e. using a scalpel) would have resulted in different

amounts of tissue being removed from each construct due to the inherently irregular shape of

each individual OAF tissue. 3-mm and 7-mm biopsy punches were therefore chosen to trim

excess cartilage, however both left behind overhanging tissue around the OAF. A 4-mm biopsy

punch was too large to fit into the centre of the OAF, and a 6-mm biopsy punch was too small to

fit around the entire perimeter of the irregularly shaped OAF tissues. Dissection of the OAF 1

mm above the interface was performed by hand under a dissecting microscope. Precision using

this method was difficult to achieve. As a result, the tissues analysed for biochemistry contained

both interfacial and non-interfacial tissues, which likely affected the biochemical values

obtained. Secondly, the biochemical assays used for determination of overall DNA, OH-Pro and

sulfated GAG content may not have been sensitive enough to determine significant small

changes at the interface. The excess cartilage tissue around the interface and the 1 mm of OAF

tissue above the interface may have concealed any detectable OH-Pro or GAG changes that had

occurred in the interface region over time. Lastly, as the assays were only able to detect overall

hydroxyproline and GAG content, they did not provide any information as to whether different

subtypes of collagen or proteoglycan may have increased or decreased over time.

Another aim was to determine the cellular composition at the interface, more specifically

whether chondrocytes were migrating upwards into the OAF region. Notochordal cells have been

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shown to promote CEP chondrocyte migration in vitro214, however the effect of AF cells on

chondrocyte migration is unknown. CFDA has low toxicity to cells at appropriate concentrations

and has been used previously to show migration of chondrocytes from engineered cartilage into

adjacent host cartilage198. Upon diffusion into the cell, CFDA is cleaved by intracellular

esterases and reacts with intracellular amines to produce a stable fluorescent signal. The results

showed that the green-labelled cells in 2-week constructs remained within the cartilage layer and

did not migrate up into the OAF. Examination of a CFDA-labelled cells of the 4-week construct

did not show evidence of chondrocyte migration into the OAF, however the fluorescent signal

was much weaker compared to the 2-week construct. A limitation of using CFDA is that its

concentration is reduced by half with each cell division. Thus, it is not known whether the

decrease in staining reflected AF cell migration into the cartilage, or chondrocyte proliferation

causing dilution of the dye below detectable limits. Stable, long-term labelling of both

chondrocytes and OAF cells by transfection with green or red fluorescent protein would confirm

the exact distribution of OAF cells and chondrocytes at the interface.

4.2.4 Failure testing and apparent mechanical strength of the OAF-CEP interface

Axial tensile testing to failure revealed that co-culture time had a significant effect on the

mechanical strength of OAF-CEP interface constructs. 2-week co-cultured constructs possessed

measurable mechanical strength, which was significantly enhanced by 4 weeks of co-culture.

Collagen type I and collagen type II appeared to accumulate at the interface from 2 to 4 weeks of

co-culture as shown by immunostaining. This suggested that while attachment between OAF and

CEP-like tissues occurred early (within 1 week of co-culture, visible by histology), there was a

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period of development and ECM deposition that functioned to further enhance the integration

and mechanical strength between the two tissues.

It was of importance to note that although there was no difference in the quantified ECM

proteoglycan or collagen contents between 2 and 4 weeks of co-culture, there was an increase in

the apparent mechanical strength of the interface. There are several possible explanations for

this. It is possible that small changes in matrix accumulation at particular points along the

interface may have been insufficient to be detected by the relatively insensitive biochemical

assays. Despite using a standard method of tissue dissection for all samples, it was difficult to

isolate only the interfacial region of each construct for subsequent biochemical analysis. Due to

this lack of precision, excess non-interfacial OAF and cartilage tissues were included in the

biochemical analyses which may have masked any small ECM changes that occurred specifically

at the narrow interface region. Inherent variability within interface tissues, manifesting for

example in the form of discontinuities within integrated constructs (Supplemental Figure S8) and

non-uniform protein distribution along the length of the interface due to the presence of a

scaffold layer (seen in Figure 3.13F), also limited the accuracy of biochemical quantification,

thereby limiting our ability to isolate the specific points of integration that may have been

responsible for the enhanced mechanical strength.

While the increase in mechanical strength therefore cannot be directly explained by any

significant increases in overall collagen or proteoglycan content per mg of DNA, the apparent

accumulation of collagen types I and II at the interface, as well as the increase in aggrecan across

the interface boundary as detected by immunostaining, may provide an indirect explanation for

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the increased mechanical strength that was observed in 4-week as compared to 2-week tissues.

The proteoglycan seen in the interfacial OAF tissue at 4 weeks may be an indication of tissue

remodelling, which may have contributed to the increase in apparent interfacial strength.

Hedlund et al. provided evidence that aggrecan monomers interact directly with collagen fibrils

through their keratin sulfate-rich domains215. Increasing such aggrecan-collagen interactions in

the interfacial OAF over time may have enhanced the mechanical properties at this region. Due

to its osmotic swelling properties, the main role of aggrecan in cartilaginous tissues is to resist

compressive forces while collagen fibres resist tensile forces64,216. The increased spreading of

aggrecan through the OAF therefore does not necessarily explain the higher tensile strength

observed at the interface. However, it does indicate that some level of ECM remodelling might

have occurred in the tissue and that this may have contributed to the observed increase in tensile

strength.

Similar to the presently observed increase in tensile strength as a result of tissue maturation,

previous work by Williamson et al. found that increases in both compressive and tensile

properties from fetal to mature cartilage were not related to changes in quantified proteoglycan

content, but to an increase in the amount of pyridinoline, a chemical cross-link of collagen217,218.

The proportion of mature compared to immature collagen cross-links has also been shown to

increase with cartilage tissue maturity219. Lysyl oxidase is the enzyme that is responsible for the

formation of pyridinoline cross-links in both collagen and elastin220, and has been suggested to

improve integrity and integration strength between TE and native cartilage221. Thus, the observed

increase in apparent mechanical strength may reflect increased activity of this enzyme resulting

in the catalyzation of collagen cross-links. Alternatively, just as fibrils of the native AF extend

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towards adjacent CEPs to form interconnections with them in the later stages of development

(seventh embryonic month in humans)74, it may be that increased collagen fibril intermingling

occurred with maturity of the model interface. Cartilage collagen fibril diameter and alignment

have also been shown to increase with maturity, which was related to an increase in mechanical

properties219. The OH-Pro assay used in the present study only quantified the amount of

collagen, and not the level of fibril cross-linking, intermingling, size, or organization. Scanning

electron microscopy imaging may provide insight into the level of collagen fibril intermingling,

organization, and diameter increase at the interface over time. Quantification of lysyl oxidase

gene and protein expression, as well as pyridinoline cross-link content at the in vitro OAF-CEP

interface should be performed to determine if a cross-linking mechanism is responsible for the

enhanced mechanical strength observed at the interface over time, in the absence of changes in

collagen content.

The weaker interface at 2 weeks raises the possibility that tissue remodelling may have been

more dynamic at this early time point, with the possibly of more protease activity leading to less

stable ECM. Matrix metalloproteinases (MMPs) are enzymes that mediate ECM remodelling by

regulating ECM protein degradation, cell migration, and molecular signalling222. Their specific

roles with respect to IVD development have not been extensively investigated. However, one

study found that MMP-1, MMP-3, MMP-14, and active MMP-2 were more highly present in the

NP, AF and CEP of human fetal IVDs than in non-fetal IVDs, suggesting that MMP-mediated

degradative processes may contribute to IVD ECM remodelling during early tissue development

and may be less active as maturity is approached223. In this regard, the development of ECM and

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mechanical strength over time in the current model may have been mimicking the temporal

progression of physiological developmental remodelling processes.

In vivo, the mineralized layer of the CEP serves to reinforce the anchorage of native AF fibrils

into the CEP and the underlying bone100. In the current model, OAF scaffold fibres did not

extend through the entire length of CEP and there was no calcified cartilage present to provide

the ability for integration with a mineralized layer. This is likely one reason why construct failure

predominantly occurred at the attachment site between the OAF and cartilage rather than within

each respective tissue. The native interface is mechanically stable under tension. Uniaxial tensile

failure testing of native bovine VB-OAF-VB samples found that failure never occurred at the

OAF-CEP interface; it rather occurred in the adjacent VB growth plate or OAF tissue, with

failure stress values between 1.2 and 1.4 MPa depending on the number of lamellae present in

each sample (Rok 2017, unpublished). In comparison, the apparent mechanical strength of the

current engineered model was 21 kPa, approximately 57 times weaker than the mechanical

strength of native VB-OAF-VB segments containing 6-10 lamellae (Rok 2017, unpublished).

Importantly, it was noted that failure within our in vitro interface occurred between the OAF and

cartilage tissues, and rarely occurred within the OAF or cartilage phases themselves. Since the

current developed model focused only on the AF-hyaline cartilage interface and did not have

features to mimic native AF-VB fibril insertion or CEP mineralization, it is not unexpected that

the tensile failure strength of the generated interface was not yet comparable to that of the native

interface. The native OAF-CEP interface normally possesses a high degree of mechanical

strength due to the extension of the AF collagen fibers through the CEP and mineralized

cartilage, and into the VB1,102. Sapiee et al. demonstrated that failure across the AF-CEP-VB

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interface occurred either in the soft AF tissue or in the vertebral bone itself, where AF fibres

pulled out the VB osteons with them, indicating a high degree of integration with the VB102.

Thus, mimicking the anchorage of the AF through a cartilage layer, a mineralized cartilage layer,

and then into a subchondral bone region should be a goal of further experiments in order to

strengthen the tissue-engineered AF-cartilage interface. Despite its limitations, this model system

will still allow for studies on the first critical steps of integration that will facilitate the

recapitulation of mechanical stability following implantation.

The apparent mechanical strength of the interface model was estimated using maximum force

measurements obtained at failure during axial tensile testing, which was dependant on

determining the apparent surface area of the OAF in contact with the cartilage. The major

limitation here was that OAF-cartilage contact area was assumed to be the surface area

measured. However, this may not necessarily be an accurate assumption as it is unknown if all of

the OAF tissues were in contact and fully integrated with the underlying cartilage tissue. OAF

contact surfaces were painted with black ink in order to provide enough contrast for tissue edge

detection during imaging and surface area determination. Tracing the surface area in ImageJ was

performed manually due to irregularities in the OAF tissue images (i.e. holes and discontinuities

within the tissues, ambient light reflections). It is acknowledged that this method of surface area

determination was crude, as ink was applied by hand, and the tracing of surface area using

ImageJ software was performed manually. The methods were however kept consistent

throughout, and the obtained mechanical strength results are reported as apparent values.

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4.2.5 Dynamic mechanical loading of the in vitro OAF-CEP interface

Dynamic compression did not have a significant effect on the tensile strength or composition of

OAF-CEP constructs. It may appear from the images that loading enhanced proteoglycan

accumulation in comparison to the unloaded construct (Figures 3.13B, E), however this was not

supported by aggrecan immunostaining which showed no change in response to loading (Figures

3.13D, G), nor by any change in quantified GAG content (Figure 3.13L).

Previous studies have reported anabolic effects of dynamic mechanical compressive and tensile

loading on IVD and cartilage native tissue, engineered tissue, and cells (Chapter 2, Table 1).

However to our knowledge, there has been no study of the effect of loading on the AF-CEP

interface. We conducted a literature review to attempt to identify potential conditions for

mechanical loading. It was apparent that application of loading magnitudes and frequencies

outside of a “physiological range” may lead to catabolic changes such as decreased ECM

macromolecule gene and protein expression and increased matrix degrading enzyme expression

(examples of physiologic levels of intradiscal pressure in the human spine can range from 0.1

MPa while lying down, to 1 MPa while jogging, to over 2 MPa while lifting a heavy weight with

improper back posture224. 1 Hz is an example of a physiologic walking frequency51)33,35,42,45.

Based on a review of the literature, a 10% radial displacement of the OAF during agarose plug

compression was chosen as the standard for which the applied force was not to exceed. A force

of 0.4 N was determined to be large enough in magnitude to cause sufficient radial displacement

of the OAF tissue, yet small enough as to not crush the tissue construct or the agarose plug.

Although tissue constructs remained intact upon harvesting and handling at the end of the culture

period, histologic examination revealed that dynamic loading led to destruction of the interface

in 2 of the 3 samples examined. Biochemical analysis of the destroyed interface tissue might

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have masked any anabolic effects that may have occurred in other intact regions. DNA content

analyses did not however indicate a loss of cellularity, as DNA content between unloaded and

loaded constructs remained unchanged. In the case that anabolic changes did occur, DNA

content may have appeared to remained unchanged due to being offset by the loss of tissue after

destruction, leading to an apparent net change of zero.

The application of dynamic mechanical load to OAF-CEP constructs via compression of a

centrally-placed agarose plug did not lead to significant changes in interfacial mechanical

strength, ECM composition as determined by evaluation of histology and

immunohistochemistry, or proteoglycan or collagen content as quantified by biochemical

analysis. Similar to these findings, Tsai et al. found that the application of dynamic compression

to TE-IVD constructs seeded with passaged human AF/MSC or NP/MSC cells did not result in

detectable changes in GAG content per unit of DNA, however it was suggested by scanning

electron microscopy that ECM was more uniformly deposited between scaffold nanofibers as

compared to unloaded controls40. The authors suggested that dynamic loading promoted cell

infiltration into the nanofibers, which led to more homogenous deposition of ECM across the

scaffold and between its layers. Our study did not find such evidence of enhanced ECM

deposition in response to dynamic loading at the interface. It may be that this reflects differences

in species (human versus bovine), cell type (co-cultured MSC and AF cells versus AF cells

only), cell passage (passaged versus primary), or scaffolds (PCL/PLA versus PU-ADO) between

the two studies. Although the load frequency and strain for both studies were similar (1 Hz, 10%

strain), Tsai et al. adhered to a more rigorous loading schedule (1 hour every day for 3 weeks)40

as compared to the current study (30 minutes every other day for 1 week). The previous study

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also differed in the timepoints selected for construct pre-culture time. Tsai et al. seeded scaffold

strips with cells overnight before assembling them into IVD constructs, which were then only

allowed to incubate again overnight before being placed under dynamic load40. This contrasts

with the protocol of the current study, whereby OAF cell-seeded constructs were allowed to

grow tissue for 2 weeks and integrate with cartilage for an additional 2 weeks before being

subjected to dynamic loading. Recent work by Fotticchia et al. in fact found a detrimental effect

of dynamic loading to an AF construct which was similar in design to the current model,

however the tissue was only pre-cultured for 2 days before dynamic compression was

introduced53. A two-layered MSC-seeded electrospun polymer AF construct with a centrally-

placed acellular agarose NP-like plug was subjected to dynamic compression (1 Hz), and showed

decreased cell viability at both 5% and 10% vertical displacement as compared to low static

compression controls53. It has been suggested that MSCs may be induced to express an AF-like

phenotype under appropriate media cocktail conditions40, however 2 days of cell seeding prior to

compressive stimulation may have only been enough time for cell attachment and was likely not

sufficient time to generate tissue, let alone AF-like tissue. Additionally, constructs were

subjected to dynamic load for 8 hours per day (every other day for 1 week)53, which is a duration

that has specifically been suggested to have a catabolic effect on AF ECM accumulation in

comparison to shorter durations between 1 and 4 hours41. The detrimental effects that were

observed in response to loading as shown by Fotticchia et al. provide further support for the

design of our study, where seeded scaffolds were allowed to grow robust tissue for 2 weeks

before being placed under dynamic load for a relatively short duration. The combined effects of

construct pre-culture time and duration of dynamic load on ECM accumulation and integration

require further investigation.

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To the best of our knowledge, the generation of a cell-based TE IVD-CEP interface has not been

previously investigated, and the response of such an interface model to mechanical loading is

therefore a novel area of study. Previous studies have investigated the effects of loading on

single tissue types alone and did not need to consider the risk of physical disruption between a

newly formed two-tissue interface. Difficulties are introduced when another tissue type is

incorporated into the engineered model to form an interface, which is vulnerable to separation

(specifically by shearing) upon mechanical loading. Any mechanical strength that an in vitro

generated interface possesses might not be enough to resist an applied mechanical load that has

previously been found to be an appropriate magnitude for the stimulation of single in vitro

tissues. In the current study, a 10% strain value was selected from previous literature as the

standard for which to aim for with respect to OAF distension under load. However, this value

was likely too high with respect to the generated OAF-CEP interface model, as it led to interface

separation in some samples, likely due to shearing of the OAF from the cartilage. It can be seen

from Figure 3.12D that there is heterogeneity in the percent diameter change between individual

OAF samples under the same load of 0.4 N. Although the mean diameter change pooled from all

samples did not exceed 10%, it can be seen from the scatter plot that one sample tested had a

mean diameter change that did exceed 10% and had individual measurements that approached a

15% change. This is due to the inherent heterogeneity of the OAF tissues generated, as

irregularities in tissue thickness within and between samples are unavoidable under the current

method of generation, thus contributing to the variability observed. At this magnitude during

dynamic loading, some samples therefore experienced larger diameter changes and were

therefore more likely to shear off from the cartilage than others. The average maximum force

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values from the axial tensile testing of 2-week and 4-week OAF-CEP constructs were lower than

the 0.4 N force selected for dynamic compression of the central agarose plug (~0.1 N for 2-week

and ~0.2 N for 4-week, data not shown). Although axial tensile failure results do not translate

into values for shear strength at the interface, it might be of interest to use these failure values in

the determination of optimal loading conditions. Specifically, load values of no more than 0.1 N

to 0.2 N should be tested in order to minimize the risk of ECM disruption at the interface.

A centrally-placed agarose plug has been used in a previous study to model the NP within an

electrospun multilamellar AF construct155 and has been suggested by mathematical modeling to

be an appropriate material for dynamic compression within the current AF model system (Rok

2017, unpublished). The agarose plug used in our study was larger in height than the OAF by

approximately 2 mm, so as to only investigate the effects of OAF lateral expansion caused by

agarose (NP) compression. In generating the constructs, OAF tissues were not always attached in

the centre of the cartilage, and were sometimes skewed closer to the edge of the membrane

insert. The jig indenter surface area used in our study was therefore designed to be small enough

to fit within the plastic insert and large enough to cover the entire surface of the OAF construct,

which ensured that the entire surface area of the agarose plug would be contacted by the indenter

regardless of the OAF’s position atop the cartilage. Thus, the application of force to the agarose

plug across all samples was assumed to be uniform. However, small differences between

individual agarose plug height dimensions as a result of variability between batches may have

led to variable applications of outward force to OAF tissues. Additionally, the forces applied in

the current model system did not fully model what occurs in the native IVD. Axial compression

applied to the agarose plug led to radial tension and shear at the in vitro OAF-CEP interface,

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whereas the native tissue also experiences compression and torsion. The current loading

conditions were selected in attempt to tease out the potential roles of different forces (i.e. tension

and shear may have had different effects on the interface compared to radial tension, shear and

compression combined). The unexpected outcome of interface disruption prevented any potential

effects from being found. It is possible that this focus on only tension and shear may have

favoured interface disruption. Further optimization of loading conditions is required when it

comes to the mechanical stimulation of a two-tissue interface.

4.3 Recommendations for Future Work

Despite modeling particular structural and compositional features, this model did not fully mimic

native interface characteristics. The interface region of the native AF and CEP demonstrates

anchorage of the relatively soft AF collagen fibres to the hard VB100,102. The CEP is the

intermediate tissue through which this occurs. AF fibril interconnections with the CEP

cartilaginous phase are strengthened by the CEP mineralized zone. Specifically, these

characteristics contribute to the mechanical strength at this interface100,102, and provide a clear

direction for subsequent future work. As the major structural difference between the current

OAF-CEP model and the native interface is the absence of AF fibril insertion through a

mineralized cartilage layer that is integrated with the underlying bone63, it is anticipated that an

enhancement of mechanical properties could be achieved in future iterations of the model that

include these components.

The current model utilized OAF cells only for the generation of in vitro tissue on PU-ADO

scaffolds, which was confirmed by immunostaining to compositionally resemble native OAF

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tissue with respect to the presence of collagen type I and the lack of collagen type II

(Supplemental Figure S7). The native AF is composed of distinct inner and outer regions, each

differing in their respective ratios of collagen type I to collagen type II. Previous work on the

current AF culture system has shown that DMEM supplemented with dexamethasone, pyruvate,

proline, and ITS maintained the distinct phenotypes of both OAF and IAF cells31. Future work

should use this system to determine if in vitro formed AF tissue containing both IAF and OAF

components will integrate to form an interface with CEP-like tissues as well.

The interface that was successfully formed between OAF and CEP-like tissues only represents

one amongst five other interfaces within the IVD space: IAF-NP, IAF-OAF, NP-CEP, IAF-CEP,

and CEP-VB. Previous work has generated a biphasic in vitro bovine IAF-NP interface that

accumulated mechanical strength over time in co-culture194, as well as a biphasic cartilage-CPP

construct that showed improved mechanical properties upon in vivo implantation225. These along

with the current study demonstrate promise for successful generation of the other necessary

biphasic IVD interface models that must be developed in order to build a complete multiphasic

implant model consisting of distinct and integrated parts. Determining if cartilage grown on CPP

can integrate similarly with multilamellated OAF tissues will be of particular interest, as a hard

bone material will be part of an eventual implant model. Incorporation of a mineralized zone

within the cartilage layer will also be necessary. Mineralization of in vitro cartilage tissues may

be induced in a similar manner to what occurs in vivo by the addition of β-glycerophosphate into

the media cocktail226. The presence of both NP and AF cell-seeded components in a future IVD-

cartilage model may also provide insight into the effect of dynamic loading at the interface(s), as

crosstalk between cells might influence outcomes with respect to mechanical loading.

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The current study only examined the presence of three major ECM molecules known to be

present in native AF and CEP tissue: collagen type I, collagen type II, and aggrecan. Other ECM

molecules such as collagen type VI (a potential marker of fibrocartilage96) and small leucine-rich

proteoglycans (SLRPs; e.g. decorin, biglycan, and fibromodulin; involved in collagen

fibrillogenesis) have been reported in the developing and mature native AF, CEP, and interface

regions72,73. Fibromodulin may be of particular interest, as it was reported to occur in the

developing AF-CEP region during embryogenesis and hence might play a specific role in tissue

integration and organization at the interface73. Elastin fibres are also present at the IVD-CEP

interface104. Future work should therefore further characterize the in vitro OAF-CEP model with

a wider panel of ECM components in mind.

Based on the outcomes of the current study it is also recommended that dynamic loading be

applied to an in vitro OAF-CEP construct at a later timepoint, perhaps after 3 or 4 weeks of static

co-culture. The current protocol applied dynamic compression to a 2-week old construct, before

there was visible proteoglycan (aggrecan) diffusion past the interface and into the OAF region.

This led to the destruction of the interface in some samples. Two weeks may not have been

enough time to allow for sufficient ECM synthesis and subsequent remodeling to occur between

the in vitro tissues, and too early to introduce a mechanical stimulus to see any enhancing effect.

Aggrecan has been found to protect collagen against proteolytic degradation227 and therefore

might have a protective effect against disruption of collagen fibers at the interface during

mechanical loading. This further supports the rationale for application of dynamic loading at a

later timepoint, by which time aggrecan may be more abundant in the interfacial OAF region. A

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prolonged period of overall ECM deposition (i.e. 4 weeks) may also be required before

mechanical loading would be able to induce anabolic rather catabolic, necrotic and/or apoptotic

responses. Establishing optimized loading conditions will be required, specifically for an

interface model which has little mechanical strength, along with the optimization of sufficient

time for remodelling to occur post-loading. Other parameters such as duration of cyclic

compression and timing of application may need to be considered. Future dynamic compression

may also consider applying dynamic load to the construct in a fashion that more closely mimics

what occurs in vivo (i.e. simultaneous compression, radial tension and shear). When loading

conditions have been optimized, other forces such as torsion and axial tension may be introduced

to the model.

Although generation of a small-scale model is required as a first step towards generating a

clinically relevant implant, it is necessary to consider the feasibility of scaling up. The outer

diameter of the OAF in the current interface model was approximately 7 mm. Gullbrand et al.

recently investigated rabbit (“medium”) and human-sized (“large”) IVD constructs generated

from IVD cell-seeded angle-ply nanofibrous polymer scaffolds228. Although the large constructs

matured compositionally in vitro, the medium AF constructs showed higher GAG and collagen

content compared to large constructs. Therefore, effective scale-up of implant constructs that

structurally mimic the native IVD remains challenging and should be considered in future work.

4.4 Conclusions

In summary, this is the first study, to our knowledge, that has conceived, established and

characterized a tissue-engineered AF-CEP interface that simultaneously 1) models the AF multi-

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lamellated angle-ply structure and 2) has an integrated in vitro cartilage layer to yield an

interface capable of reproducing important ECM features within an IVD model. Increased

interfacial mechanical strength can be achieved by increasing in vitro culture time. The ability to

generate integration between in vitro OAF and cartilage tissues is one step towards the

development of a fully integrated, multiphasic, biomimetic biological IVD implant. This model

can be used to investigate approaches towards enhancing integration of such tissue for in vivo

applications. Future work on this model will focus on inducing mineralization in the deep zone

of the cartilage layer in order to further mimic the native OAF-CEP-VB region. Further

investigation is required to produce enhanced strength at the construct interface.

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149

APPENDIX A: METHOD DEVELOPMENT FOR IN VITRO

OAF-CEP CONSTRUCT ASSEMBLY

Method 1: 3, 5, or 7-day old cartilage tissues were excised from their membrane insert holders

using a scalpel (see methods section for tissue generation in Chapter 3). TYGON tubing was

removed from 2-week old OAF tissues (see methods section for tissue generation in Chapter 3).

An OAF was placed on top of each circular cartilage tissue in a 6-well plate. The tissues were

submerged in 2 mL of complete media, which was sufficient volume to prevent the OAF from

floating and to ensure that it remained in physical contact with the cartilage. On the third day of

co-culture, media volume in each well was doubled to 4 mL. Complete media and 100 µg/mL

ascorbic acid were replaced every 2-3 days. The OAF-CEP tissues were harvested after 1 and 2

weeks of static co-culture. This method of OAF-CEP tissue assembly led to the collapse of

scaffold layers, visible contraction of OAF tissue, and curling of cartilage tissue (Supplemental

Figure S1A). A second method was developed (Method 2, Methods Section 3.2.5) whereby the

TYGON tubing was kept inside the OAF during co-culture in order to maintain tissue shape.

150

APPENDIX B: SUPPLEMENTARY FIGURES

151

Figure S1. Culture of OAF-CEP constructs with centrally-placed tubing prevented PU-ADO

scaffold convolution (i.e. scaffold folding onto itself; see Figure 3.7C, Chapter 3). (A) Co-culture

of OAF-CEP using Method 1 resulted in contracted, irregularly shaped 3-layered OAF tissue and

curling of the cartilage. (B) Co-culture using Method 2 (with centrally-placed TYGON tubing)

allowed OAF to maintain a regular concentric shape, and resulted in flat cartilage. (C)

Histological cross sections of one wall of a 2-week old in vitro OAF tissue harvested

immediately after dynamic culture showed that scaffold lamellae were relatively straight and not

curled. (D) A 2-week old OAF cultured statically by itself for an additional week in an upright

position showed downward collapse of the scaffold (black arrows). OAF-CEP constructs

assembled with physical support from tubing showed ends of scaffold lamellae inserting

perpendicularly towards the cartilage (E). H&E stained. Scale bars represent 100 µm.

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Figure S2. After curing at 45° C overnight, the 6 cm-wide electrospun PU-ADO scaffold and

excess aluminum foil (A) were cut down using scissors (B) to a width of 5 cm (C). (D) Scaffold

was peeled from the foil, with care being taken to not stretch the scaffold too much or cause tears

along the fibres. (E) A paper template was taped to the underside of a single-well plate lid to be

used as a guide to cut 2.7 mm-wide scaffold strips at a 30° angle relative to nanofiber direction.

(F) The scaffold was aligned with the template, and the foil that underlaid the scaffold was taped

to the surface of the lab bench to hold it in place. (G) Scaffold strips were cut using a scalpel

along the template lines. (H) Scaffold strips were peeled carefully from the template using thin

tweezers, and (I) placed in a 50 mL conical tube for subsequent sterilization by gamma

irradiation.

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Figure S3. DIC and fluorescent microscopy confirmed that 96% of cells were labelled with

green fluorescent CFDA dye.

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Figure S4. (A) Schematic diagram of plate set up for dynamic loading. 3 pieces in addition to

the jig were required in order to keep the insert holder in the centre of the well, to prevent it from

moving up and down with the indenter during loading, and to ensure compression of the

agarose/OAF-CEP construct during dynamic loading: A glass disc (1), having a height equal to

the height of the insert holder feet (1 mm). Placement of the glass disc under the membrane

insert holder (2) ensured that the membrane was in contact with a hard surface and did not

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stretch downwards during compressive loading; A plastic ring (3), made by cutting off the

bottom portion of an insert holder, provided a continuous surface with the wall of the insert

holder and thus an unobstructed path for the indenter to move through; A Teflon ring (4) with

dimensions to fit exactly around the circumference of the insert holder, and inside of the plate

well. (B) Assembly of all pieces as they appeared inside the 24-well plate. The Teflon ring rested

above the lip of the insert holder and ensured that it was locked down upon placement of the jig

plate cover and during loading. (C) Pieces were assembled inside a standard 24-well plate under

sterile conditions. The glass disc was the first to be placed inside, followed by the insert holder

(containing the OAF-CEP, not shown), the plastic ring, and lastly the Teflon ring. (D) When

assembled, the top surfaces of the outer Teflon ring and inner plastic ring were flush with the top

of the plate. The indenter contacted the surface of the agarose plug during loading.

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Figure S5. (A) During cryostat sectioning, cryotape (kept cold inside the cryostat before use)

was cut to size with scissors and placed sticky side down (gold side of tab down, silver side up)

onto the surface of frozen blocks. (B) Air bubbles were removed using a small paint brush. (C) 7

μm-thick sections were cut. (D) Sections were handled using thin tweezers and (E) transferred

temporarily onto room-temperature glass slides (gold side up). (F) A thin layer of optical glue

was spread onto clean glass slides. (G) The metallic tab of the cryotape was trimmed off from

each section with scissors. (H) Trimmed sections were placed on top of the layer of glue,

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ensuring that the tissue side of the sections faced up and that air bubbles between the glue and

crytotape were minimal. (I) Sections were placed under the UV lamp of a biosafety cabinet for 1

hour to allow the glue to cure.

Figure S6. Scanning electron microscopy images of the top side (A) and underside (B) of a

scaffold sample and their corresponding fibre alignment.

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Figure S7. (A) Collagen type I (green) was positive and collagen type II (red) was negative in

the OAF region of 2-week and (B) 4-week OAF-CEP constructs, away from the interface. (C)

Collagen type I distribution between scaffold layers mimicked the interlamellar distribution seen

in native OAF tissue. S = PU-ADO scaffold. L = native OAF lamellae. Fluorescent microscopy.

Scale bars represent 100 µm.

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Figure S8. (A) Site of discontinuity (arrowheads) between the in vitro OAF and cartilage layer

between surrounding interfacial tissue. (B) Continuous integration between the in vitro OAF and

cartilage layer. # = OAF, * = cartilage, S = PU-ADO scaffold. Toluidine blue stained. Light

microscopy. Scale bars represent 100 μm.

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Figure S9. The paucicellular (sparse cellularity) layer of ECM that developed above 5 and 7-day

pre-grown cartilage in OAF-CEP constructs also developed in in vitro cartilage grown alone in

the AF-optimized complete media (DMEM, dexamethasone, sodium pyruvate, L-proline, ITS,

10% FBS). 5 and 7-day pre-grown cartilage cultured in control media (Ham’s F12, 10% FBS)

developed this ECM layer to a much lesser extent compared to cartilage in complete media.

Cartilage tissues in complete media accumulated more paucicellular ECM as the culture time

increased from 7 to 21 days.

Figure S10. Immunostaining of full-thickness primary chondrocytes pre-grown for 3 days in

Ham’s F12 and subsequently cultured alone for 2 weeks in (A) control media (Ham’s F12 + 10%

FBS), or in (B) complete AF-optimized media (DMEM + 10% FBS). Compared to the control,

AF media resulted in accumulation of an acellular collagen type I-rich ECM layer above the

chondrocytes (white arrows). (C) IgG antibodies were used as negative controls. Scale bars

represent 100 µm. Fluorescent microscopy.