Design and fabrication of scaffolds for cartilage ...

Design and fabrication of scaffolds for cartilage

applications through photopolymerization of hydrogels

A thesis submitted to the University of Manchester for the Degree of

Doctor of Philosophy in the Faculty of Science & Engineering

September 2019

Hussein H. M. Mishbak

Supervisor:

Professor Paulo Jorge Bartolo

Dr. Glen Cooper

School of Mechanical, Aerospace and Civil Engineering

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Table of Contents Table of Contents ....................................................................................................... 2

List of Figures ............................................................................................................. 6

List of Tables ............................................................................................................ 13

List of Abbreviations ............................................................................................... 14

List of Notations ....................................................................................................... 16

Abstract ..................................................................................................................... 17

Declaration ................................................................................................................ 19

Copyright .................................................................................................................. 20

Acknowledgements ................................................................................................... 21

Chapter 1 Introduction ............................................................................................ 25

1.1 Background .................................................................................................... 26

1.2 Research aims and objectives ....................................................................... 27

1.3 Thesis outlines ............................................................................................... 30

Chapter 2 State-of-art of Cartilage Tissue Engineering ....................................... 33

Biological perspectives and current strategies in osteochondral tissue

engineering ............................................................................................................ 34

Introduction .......................................................................................... 35

Composition of articular cartilage ........................................................ 36

Structure of osteochondral tissue ......................................................... 38

Zonal structure of articular cartilage .................................................... 39

Current osteochondral treatments ........................................................ 41

Tissue engineering: requirements and challenges ................................ 47

Conclusion and future perspectives ..................................................... 54

Engineering natural-based photocrosslinkable hydrogels for cartilage

applications ............................................................................................................ 56

Introduction .......................................................................................... 56

Articular cartilage................................................................................. 59

Engineering hydrogels for cartilage tissue engineering ....................... 64

Crosslinking mechanisms of hydrogels ............................................... 68

Biofunctionalization of hydrogels ........................................................ 69

Conclusions and future perspectives .................................................... 76

Chapter 3 Development and characterisation of a photocurable alginate bioink

for 3D bioprinting .................................................................................................... 78

Introduction ................................................................................................. 79

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Materials and methods ................................................................................. 81

Synthesis of methacrylate alginate ....................................................... 81

Characterisation of pre-polymerised methacrylate alginate ................. 81

Hydrogel formation .............................................................................. 82

Characterisation of photocrosslinked hydrogels .................................. 83

Results and discussion ................................................................................. 84

Alginate functionalization .................................................................... 84

Rheological behaviour of pre-polymerised alginate methacrylate

systems .............................................................................................................. 87

Rheological changes during the photopolymerization process ............ 92

Viscoelastic properties of formed hydrogels........................................ 93

Internal morphology of alginate hydrogels .......................................... 95

Mechanical characterisation ................................................................. 96

Swelling and degradation kinetics ....................................................... 97

Conclusion ................................................................................................... 98

Chapter 4 Photocurable alginate bioink development for cartilage replacement

bioprinting ................................................................................................................ 99

Introduction ............................................................................................... 100

Materials and methods ............................................................................... 101

Synthesis of methacrylated alginate ................................................... 101

HNMR ................................................................................................ 101

Rheological characterization .............................................................. 102

Cell viability and proliferation ........................................................... 102

Result and discussion ................................................................................ 102

Chemical modification through 1HNMR ........................................... 102

Rheological characterization of alginate methacrylate ...................... 104

Cell viability assessment .................................................................... 105

Conclusions ............................................................................................... 106

Chapter 5 Photocrosslinkable gelatin methacrylate (GelMA) hydrogels:

Towards 3D cell culture and bioprinting for cartilage applications ................. 107

Introduction ............................................................................................... 108

Experimental ............................................................................................. 109

Synthesis of methacrylated gelatin .................................................... 109

Photocurable hydrogel solutions ........................................................ 110

HNMR characterization ..................................................................... 110

Rheological characterization .............................................................. 110

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Scanning Electron Microscopy .......................................................... 111

Mechanical characterization............................................................... 111

Swelling and degradation characterization ........................................ 111

hADSCs Cell encapsulation ............................................................... 112

Data analysis ...................................................................................... 113

Results and discussion ............................................................................... 113

Characterization of gelatin and GelMA by HNMR ........................... 113

Rheological characterization ..................................................................... 115

Prepolymer rheology .......................................................................... 115

Morphological characterization ......................................................... 119

Mechanical characterization............................................................... 121

Swelling and degradation characterization ........................................ 121

Cell viability and proliferation ........................................................... 122

Conclusion ................................................................................................. 124

Chapter 6 Development and characterisation of a photocurable alginate-gelatin

for cartilage applications ....................................................................................... 125

Introduction ............................................................................................... 126

Materials and methods ............................................................................... 128

Photocrosslinkable polymer preparation ............................................ 128

Alginate methacrylate preparation (AlgMA) ..................................... 128

Hydrogel formation ............................................................................ 128

Structural conformation of the hybrid hydrogels ............................... 130

Rheological characterisation .............................................................. 130

Structure morphology and porosity .................................................... 131

Swelling and degradation kinetics of the hybrid hydrogels ............... 132

Mechanical assessment ...................................................................... 132

Biological assessment ........................................................................ 132

Histological analysis .......................................................................... 133

Glycosaminoglycan quantification..................................................... 135

Results and discussion ............................................................................... 136

Functionalisation confirmation and rheological characterisation ...... 136

Rheological Characterization ............................................................. 137

Structure morphology and porosity .................................................... 141

Swelling and degradation kinetics and hydrogel ............................... 142

Mechanical characterisation ............................................................... 143

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Hydrogel encapsulated chondrocyte viability and proliferation ........ 144

Histological assessment of hydrogel ECM formation ....................... 146

GAGs quantification .......................................................................... 151

Conclusion ................................................................................................. 152

Chapter 7 Hybrid PCL/Hydrogel scaffold fabrication and in-process plasma

treatment using PABS ........................................................................................... 153

Introduction ............................................................................................... 154

Materials and Methods .............................................................................. 157

Polycaprolactone (PCL) ..................................................................... 157

Hybrid hydrogel) methacrylate anhydride (Alg-Gel)MA preparation .....

............................................................................................................ 157

Plasma-assisted Bioextrusion System set up...................................... 158

Scaffolds printing strategies ............................................................... 160

Nuclear Magnetic Resonance characterisation .................................. 161

Rheological characterisation .............................................................. 161

Morphology Characterization ............................................................ 161

Wettability measurement ................................................................... 162

Biological tests ................................................................................... 162

Data analysis ...................................................................................... 163

Results ....................................................................................................... 163

Hybrid Scaffold Fabrication ............................................................... 163

Full-layer treated PCL scaffold .......................................................... 163

Wettability assessment of full-layer plasma modified scaffold ......... 164

Biological assessment of full-layer plasma modified scaffold .......... 166

Discussion ................................................................................................. 166

Conclusion ................................................................................................. 167

Chapter 8 Conclusion and future work ............................................................... 169

Conclusion ................................................................................................. 170

Future works .............................................................................................. 173

References ............................................................................................................... 175

6

List of Figures

Figure 1.1: Framework structure of this thesis. ........................................................ 32

Figure 2.1: Comparison between the physiology and healing capacity of bone and

cartilage. Cartilage is avascular tissue so lacks ready access to a supply of circulating

stem cells and nutrients so relies on the synovial fluid for nourishment. This combined

with its largely acellular composition and low metabolic activity results in a nearly

complete lack of innate regenerative capacity. Consequentially, defects fail to heal,

and clinical interventions typically result in the formation of inferior fibrocartilage. In

contrast, bone has a greater ability for self-repair due to the constant tissue remodelling

by osteoblasts and osteoclasts. Bone is also a highly vascularised tissue which allows

a ready supply of nutrients and proteins that stimulate bone repair. Furthermore, there

is a large source of stem cells in the bone marrow and periosteum which are able to

differentiate into bone producing cells. This allows bone to heal defects up to a certain

critical size after which vascularisation becomes an issue. Image from reference. ... 36

Figure 2.2: Hierarchical and graded composition, structure, and properties of

osteochondral tissue. a) Procollagen polypeptides bind together to form collagen that

assemble into microfibrils and then fibrils which can be crosslinked together by

collagen type IX. Collagen fibrils exhibit a characteristic banding pattern of ~67 nm

due to the staggered packing arrangement of collagen. Image from reference . b)

Molecular composition and arrangement of the chondron depicting the pericellular,

territorial, and interterritorial matrix with increasing distance from the chondrocyte.

Image from reference. c) Zonal structure and properties of osteochondral tissue. Image

adapted from .............................................................................................................. 39

Figure 2.3: Current treatment options for cartilage regeneration. a) A full-thickness

chondral lesion (Grade III). b) Debridement of the lesion is used as precursor to more

invasive procedures as it removes damaged cartilage and bone to create a healthy

border which enables improved integration of neotissue with native tissue. In smaller

cartilage defects debridement may be used without further procedures as a minimally

invasive treatment option. c) Microfracture is treatment that drills into the subchondral

bone to create channels that allow MSCs to migrate into the defect. d) Autologous

chondrocyte implantation (ACI) uses a population of 12-48 million autologous

chondrocytes to which are inserted into the defect which is then covered with either a

periosteal patch or a collagen membrane. e) Matrix-induced autologous chondrocyte

implantation (MACI) is a development of the ACI technique. The autologous

chondrocytes are cultured in vitro and seeded onto an absorbable 3D scaffold (collagen

or hyaluronic acid) before being implanted into the defect and fixed to the defect with

fibrin glue. Image from reference. ............................................................................. 42

Figure 2.4: a) Discrete cartilage defect; b) End stage arthritis; c) Full thickness area

of cartilage loss. ......................................................................................................... 45

Figure 2.5: a) Normal cartilage; b) Fibrocartilage next to hyaline after microfracture.

.................................................................................................................................... 46

Figure 2.6: Tissue engineering strategies for articular cartilage regeneration. Image

from reference . .......................................................................................................... 48

Figure 2.7: 3D bioprinting technologies; a) Inkjet bioprinting; b) Laser-assisted

bioprinting; c) Extrusion bioprinting . ....................................................................... 50

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Figure 2.8: Bioink properties for successful 3D bioprinting require a suitable a)

biofabrication window which balances printability and biocompatibility whilst

providing a variety of b) suitable characteristics. ...................................................... 51

Figure 2.9: Schematic diagram of photocrosslinkable 3D hydrogel constructs and

approaches for use in cartilage tissue engineering. .................................................... 59

Figure 2.10: a) Articular cartilage covers the surfaces of bones in diarthrodial joints

(e.g. knee and elbow); b) Zonal structure and composition of articular cartilage .... 60

Figure 2.11: Stages of articular cartilage damage. 0) Healthy cartilage. I) Softening,

swelling and fissuring of the cartilage. II) Lesions <50% of cartilage depth. III)

Lesions >50% of cartilage depth but not through subchondral bone layer. IV) Severely

abnormal with penetration through the subchondral plate. Reproduced with permission

from the International Cartilage Regeneration & Joint Preservation Society (ICRS).

.................................................................................................................................... 62

Figure 2.12: Current techniques for articular cartilage repair and future scaffold-based

approaches. a) Debridement and microfracture; 1-4 debridement of damaged tissue, 5

microfracture into the subchondral bone, 6 blood clot formation containing

mesenchymal stem cells (MSCs), and 7 formation of inferior fibrocartilage. b) ACI; 1

cartilage biopsy from a non-load bearing site, 2 in vitro expansion of chondrocytes to

obtain sufficient numbers for implantation, 3 a second operation cleans the damaged

area and is covered with the periosteal flap and the expanded chondrocytes are injected

under this or the expanded chondrocytes are seeded into a matrix (e.g. collagen

membrane) and fixed with fibrin glue into the defect area (MACI/MACT). c) AMIC;

a microfracture is created and a membrane placed over the top and fixed with a fibrin

glue or sutures. The membrane protects the clot formation and allows chondrogenic

differentiation of the MSCs in the blood to aid the regeneration process. d)

Osteochondral auto/allograft transplantation; cylindrical plug of fresh cartilage and

subchondral bone is taken from the patient or a cadaver donor, 1 a guide pin is placed

perpendicular to the joint surface in the defect, 2 implantation of the graft plug, and 3

securing the plugs with the insertion of screws. e) Acellular scaffold only requiring a

single surgery and in vivo recruitment and stimulation of cells for tissue regeneration.

f) Cellular-based scaffold involving two surgeries of initial cartilage biopsy,

chondrocyte expansion, and seeding into the scaffold before implantation into the

defect. Adapted with permission from the ICRS ....................................................... 63

Figure 2.13: Engineering 3D hydrogels for cartilage application through a) cell-

encapsulation, b) optimizing crosslinking density, and c) functionalizing with

biomolecules (e.g. RGD) to guide cell behavior. ....................................................... 64

Figure 2.14: Hybrid hydrogel network systems. Multi-material bioink networks are

two different component polymers or more crosslinked together. Interpenetrating

networks (IPN) is a 3D structure of two or more polymeric networks, which are

partially or fully chemically bonded but not covalently bonded, or a network of

embedded linear polymers entrapped within the original hydrogel (semi-IPN).

Nanocomposites bioink network can be produced by adding nanoparticles to the

polymeric hydrogels. Supramolecular bioink network is composed of short repeating

units with functional groups that can interact non-covalently with other functional

units, forming large, polymer-like entanglements . ................................................... 66

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Figure 2.15: Crosslinking mechanisms using in physical and chemical gelation of

hydrogels. ................................................................................................................... 68

Figure 2.16: Influence of hydrogel stress relaxation on cartilage matrix production

and formation. Immunohistochemical staining of chondrocytes embedded within 3

kPa alginate hydrogels for 21 days (scale bar: 25 µm) . ............................................ 72

Figure 2.17: Effect of chondrocytes embedded within cell-degradable hydrogels. A)

Schematic illustration of cell-laden hydrogel formation with tethered MMP

fluorescent sensor (Dab‐ GGPQG↓IWGQK‐Fl‐AhxC), and TGF‐β1; hydrogels are

crosslinked using either an MMP‐degradable peptide sequence

(KCGPQG↓IWGQCK) or non‐degradable (3.5 kDa PEG dithiol) linker. B)

Determination of in situ cleavage of fluorescent sensor by chondrocytes. C) GAG

staining of sections from cell-laden hydrogels after 28 days of culture: nuclei stained

black and GAGs stained red. D) Collagen staining of sections obtained from cell-laden

hydrogels at day 28: nuclei stained black and collagen stained blue (scale bars: 100

μm). ............................................................................................................................ 76

Figure 3.1: Alginate showing a linkage between the mannuronic and guluronic acid.

.................................................................................................................................... 80

Figure 3.2: Schematic representation of the photopolymerization process of alginate-

methacrylate. a) After exposing the polymer solution to UV radiation, the

photoinitiators generated free radicals that react with the vinyl methylene starting the

crosslinking reaction; b) The reaction propagates with macro-radicals reacting with

unreacted carbon-carbon double bonds; c) At the end through a bimolecular

termination mechanism a 3D network of crosslinked hydrogel is formed. ............... 83

Figure 3.3: a) Non-functionalized alginate 1% w/v; b) Functionalized alginate 1% w/v

after 8 hours of reaction; c) Functionalized alginate 1% w/v after 24 hours of reaction.

The functionalization is confirmed by the presence of new peaks in the spectra at 5.63

ppm and 6.09 ppm attributed to the methylene group and a peak at 1.82 ppm that

corresponds to the methyl group. ............................................................................... 85

Figure 3.4: a) Non-functionalized alginate 2% w/v, b) Functionalized alginate (2%

w/v) after 8 hours of reaction, c) Functionalized alginate 2% w/v after 24 hours of

reaction. The functionalization is confirmed by the presence of new peaks in the

spectra at 5.63 ppm and 6.09 ppm attributed to the methylene group and a peak at 1.82

ppm that corresponds to the methyl group. ................................................................ 86





corresponds to the methyl group. ............................................................................... 87

Figure 3.6: Stress versus shear rate profiles for alginate-methacrylate solutions

containing different alginate-methacrylate concentrations, reacted for 8 hours. a) 1%

w/v of alginate; b) 2% w/v of alginate; c) 3% w/v of alginate, and d) comparison of

all compositions. ........................................................................................................ 88

Figure 3.7: Stress versus shear rate for solutions containing different alginate-

methacrylate concentrations reacted for 24 hours. a) 1% w/v of alginate, b) 2% w/v of

alginate, c) 3% w/v of alginate, and d) comparison of all compositions. ................. 89

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Figure 3.8: Viscosity versus shear rate for solutions containing different alginate-

methacrylate concentrations reacted for 8 hours. a) 1% w/v; b) 2% w/v; c) 3% w/v d)

comparison of all compositions. ................................................................................ 89


methacrylate concentrations reacted for 24 hours. a) 1% w/v; b) 2% w/v; c) 3% w/v

d) comparison of all compositions. ............................................................................ 90

Figure 3.10: 2% wt. Alginate solution reacted for 8 hours. a) Storage and loss modulus

vs strain b) Storage and Loss modulus vs frequency; c) Complex modulus vs

frequency; d) tan δ vs frequency. ............................................................................... 91

Figure 3.11: 2% wt. Alginate solution reacted for 24 hours. a) Storage and loss

modulus vs strain; b) Storage and Loss modulus vs frequency; c) Complex modulus

vs frequency; d) tan δ vs frequency. .......................................................................... 91

Figure 3.12: 2% wt. methacrylate alginate with different concentration of VA-086

photoinitiators functionalized for 8 hours a) 0.5% w/v of PI; b) 1% w/v of PI; c) 1.5%

w/v of PI; d) 0.05% w/v of PI. ................................................................................... 92


photoinitiators functionalized for 8 hours a) 0.5% w/v of PI ; b) 1% w/v of PI; c) 1.5%

w/v of PI. .................................................................................................................... 93

Figure 3.14: 2% wt methacrylate alginate hydrogel with different concentration of

VA-086 functionalized for 8 hours a) 0.5% w/v of PI; b) 1% w/v of PI; C) 1.5% w/v

of PI. ........................................................................................................................... 94

Figure 3.15: 2% wt methacrylate alginate with different concentration of VA-086

functionalized for 24 hours. a) 0.5% w/v of PI; b) 1% w/v of PI; C) 1.5% w/v of PI.

.................................................................................................................................... 95

Figure 3.16: SEM images of crosslinked methacrylate-alginate hydrogel structures

obtained from alginate-methacrylate at different reaction times: 8 and 24 hours. ..... 96

Figure 3.17: Compression tests for alginate methacrylate samples (8h) with different

photoinitiator concentrations a) 0.5% w/v; b) 1% w/v; and c) 1.5 % w/v. ................ 96


photoinitiator concentrations a) 0.5% w/v; b) 1% w/v; and c) 1.5 % w/v. ............... 97

Figure 3.19: Swelling and degradation rate for 2% wt. functionalized alginate (24 and

8) hours reaction time in a) distilled water b); degradation rate in distilled water; c)

DMEM; d) degradation rate in DMEM. ................................................................. 98

Figure 4.1: 1H NMR spectra of a) unmodified alginate and; b) 2% wt methacrylate

alginate (15 mL MA) after 8 hrs of reaction; c) 2% wt methacrylated alginate (25 mL

MA) after 8 hrs of reaction; d) 2% wt methacrylate alginate (15 mL MA) after 24 hrs

of reaction................................................................................................................. 103

Figure 4.2: Rheological characterization of 2 % wt methacrylate alginate samples. a)

Stress vs shear rate profile for samples functionalized for 8 hrs reaction time with

different methacrylate concentrations; b) Stress vs shear rate profile for samples

functionalized for 24 hrs reaction time with different methacrylate concentrations; c)

Viscoelastic behaviour low methacrylate concentration (15 mL); d) Viscoelastic

behaviour high methacrylate concentration (25 mL). .............................................. 105

Figure 4.3: Live/dead results at a) day 3, in AlgMA hydrogels; b) day 3, in the control;

c) day 7, in AlgMA hydrogels; d) day 7, in the control; and Alamar blue results e)

10

Fluorescence intensity as a function of time for different time point (1, 3, 5 and 7 days).

Note: in the live/dead results green corresponds to live cells and red to dead cells.106

Figure 5.1: Schematic crosslinking mechanism of gelatin and methacrylate. ........ 110

Figure 5.2: 1H NMR spectra for non-functionalized and functionalized gelatin. (a)

Non-functionalized gelatin (b) GelMA functionalized for 2 hrs; c) GelMA

functionalized for 4 hrs; d) GelMA functionalized for 6 hrs. .................................. 114

Figure 5.3: Rheological behaviour of 10 and 15% wt of GelMA samples considering

2 and 6 hrs of functionalization time. (a)10% wt of GelMA functionalized for 2 hrs;

(b) 10% wt of GelMA functionalized for 6 hrs; (c) 15% wt of GelMA functionalized

for 2 hrs; (d) 15% wt of GelMA functionalized for 6 hrs. ...................................... 115

Figure 5.4: Oscillatory tests: storage and loss modulus of 10% wt GelMA for 2 and 6

hrs; (a) Amplitude sweep for 10 and 15% wt GelMA functionalized for 2hrs; (b)

Frequency sweep for 10 and 15% wt GelMA functionalized for 2hrs; (c) Amplitude

sweep for 10 and 15% wt GelMA functionalized for 6hrs (d) Frequency sweep for 10

and 15% wt GelMA functionalized for 6hrs. ........................................................... 117

Figure 5.5: Oscillatory results for GelMA hydrogel samples functionalized for 2 hrs

(a) Storage and loss modulus as a function of strain for 10 and 15% wt GelMA and

1% wt of PI; (b) Storage and loss modulus as a function of frequency for 10 and 15%

wt GelMA and 1% wt of PI; c) Storage and loss modulus as a function of strain for

10 and 15% wt GelMA and 1.5% wt of PI; d) Storage and loss modulus as a function

of frequency for 10 and 15% wt GelMA and 1.5% wt of PI. ................................ 118






of frequency for 10 and 15% wt GelMA and 1.5% wt of PI. ................................ 119

Figure 5.7: Cross-section images of GelMA crosslinked structures considering

different GelMA concentrations, reaction times and PI concentration; (a) 10% of

GelMA, 2 hrs of functionalization time and 1% wt of PI; (b) 10% of GelMA, 6 hrs of

functionalization time and 1% wt of PI; (c) 10% of GelMA, 2 hrs of reaction time and

1.5% wt of PI; (d) 10% GelMA, 6 hrs of reaction time and 1.5% wt of PI; (e) 15% of

GelMA, 2 hrs of reaction time and 1% wt of PI; (f) 15% of GelMA, 6 hrs of reaction

time and 1% wt of PI; g) 15% of GelMA, 2 hrs of reaction time and 1.5% wt of PI; (h)

15% of GelMA, 6 hrs of reaction time and 1.5% wt of PI. ...................................... 120

Figure 5.8: Unconfined compression tests for gelatin methacrylate hydrogels samples

with different photoinitiator concentrations a) 10% wt GelMA functionalized for 2h

containing 1% wt of PI; (b) 10% wt GelMA functionalized for 2h containing 1.5% wt

of PI; (c) 10% wt GelMA functionalized for 6h containing 1% wt of PI; (d) 10% wt

GelMA functionalized for 6h containing 1.5% wt of PI. ......................................... 121

Figure 5.9: Swelling and degradation profiles as a function of functionalization time

for different functionalized gelatin methacrylate systems; a) Swelling in DMEM ; b)

degradation rate in DMEM. .................................................................................... 122

Figure 5.10: Live/dead images of adipose-derived mesenchymal stem cells

encapsulated in samples based on 10% wt GelMA dissolved at 1% wt of PI solution

and exposed to UV light (Light intensity of 8 mW/cm2) for 8 min a) After 3 days of

11

encapsulation; b) After 7 days of encapsulation; c) After 19 days of encapsulation.

Note: Green corresponds to live cells and red to dead cells. ................................... 123

Figure 5.11: Alamar blue fluorescence intensity at different time points (1, 3, 5 and 7

days) considering samples functionalized at two different reaction times (2 and 6 hrs)

.................................................................................................................................. 124

Figure 6.1: Photocrosslinking mechanism to form a) AlgMA and b) GelMA hydrogels

using UV irradiation. ................................................................................................ 129

Figure 6.2: Photocrosslinking of hybrid alginate-gelatin systems when exposed to UV

irradiation. ................................................................................................................ 130

Figure 6.3: Representative 1H NMR spectra of a) non-modified alginate, b) non-

modified gelatin, c) AlgMA, d) GelMA, and the hybrid Alg-Gel(MA) systems e)

75:25, f) 50:50, and g) 25:75. ................................................................................... 137

Figure 6.4: Viscosity vs share rate profiles of AlgMA, GelMA, and Alg-Gel(MA)

(25:75, 50:50, and 75:25) samples. .......................................................................... 138

Figure 6.5: Oscillatory results for pre-polymerized systems a) Storage and loss

modulus as a function of strain for AlgMA systems; b) Storage and loss modulus as a

function of frequency for AlgMA systems; c) Storage and loss modulus as a function

of strain for GelMA systems; d) Storage and loss modulus as a function of strain for

GelMA systems; e) Storage and loss modulus as a function of strain for hybrid

systems; f) Storage and loss modulus as a function of frequency for hybrid systems.

.................................................................................................................................. 139

Figure 6.6: Oscillatory results for polymerized systems a) Storage and loss modulus

as a function of strain for AlgMA systems; b) Storage and loss modulus as a function

of frequency for AlgMA systems; c) Storage and loss modulus as a function of strain

for GelMA systems; d) Storage and loss modulus as a function of strain for GelMA

systems; e) Storage and loss modulus as a function of strain for hybrid systems; f)

Storage and loss modulus as a function of frequency for hybrid systems. .............. 140

Figure 6.7: Cross-section images of different hydrogel systems a) AlgMA, b)

GelMA, c) Alg-Gel(MA) 75:25, d) Alg-Gel(MA) 50:50, and e) Alg-Gel (MA) 25:75.

Scale bar: 200 µm. ................................................................................................... 141

Figure 6.8: Swelling and degradation profiles of all hydrogel groups. a) Swelling ratio

profile and b) degradation (mass loss profile).......................................................... 143

Figure 6.9: Mechanical properties of different photocrosslinked hydrogel, a) stress-

strain curve; b) Young modulus. Legend: A-alginate; G-gelatin; H-Hybrid system.

.................................................................................................................................. 143

Figure 6.10: Cell viability of human chondrocytes encapsulated in AlgMA, GelMA,

Alg-Gel(MA) 50:50, and seeded on a TCP control at day 3, 7, and 16. Scale bar: 200

µm. ........................................................................................................................... 145

Figure 6.11: Fluorescence intensity indicating the metabolic activity of human

Chondrocytes in AlgMA, GelMA, Alg-Gel (MA) 50:50 and TCP control at days 1, 3,

5, 7 and 14. ............................................................................................................... 146

Figure 6.12: Histological analysis. a) Gomori; b) Aggrecan; c) Safranin-O production

at days 14, 21, 28 and 35 in Alginate methacrylate hydrogels, Gelatin methacrylate

hydrogels and hybrid system (50:50 Alg-Gel)MA. ................................................. 148

12

Figure 6.13: Semi-quantification of the relative area of histological staining of the

AlgMA, GelMA, and Alg-Gel(MA) hydrogels at day 14, 21, 28, and 35 for a) Gomori

trichrome and b) aggrecan. ....................................................................................... 150

Figure 6.14: Assessment of GAG secretion in the photocrosslinked AlgMA, GelMA,

and Alg-GelMA hydrogels at day 7, 14, 21, 28, and 35. ......................................... 151

Figure 7.1: scaffold-based approach for tissue engineering. .................................. 155

Figure 7.2: a) Set-up of plasma-assisted bio-extrusion system; b) system diagram:

multi-extrusion system, plasma modification system and control system. .............. 159

Figure 7.3: a) Sequence of operations to fabricate hybrid PCL/hydrogel scaffolds; b)

Sequence of operations to fabricate full-layer plasma treated scaffolds. ................. 161

Figure 7.4: (a) Hybrid polycaprolactone (PCL)/hydrogel scaffold; (b) Scanning

electron microscopy (SEM) image of top view of PCL/hydrogel scaffold; (c) SEM

image of side view of PCL/ hydrogel scaffold; (d) photo of a full-layer N2 plasma-

treated PCL scaffold; (E) SEM image of top view of full-layer N2 plasma treated PCL

scaffold; (f) SEM image of filament surface of full-layer N2 plasma-treated PCL

scaffold. .................................................................................................................... 164

Figure 7.5: a) Temporal variation of fluorescence intensity of cell-seeded PCL

scaffolds with and without N2 plasma treatment; b) Confocal microscope images of

untreated and full-layer treated of cell seeded scaffolds, 1 day and 14 days after cell

culture. Scale bar 250 µm. ....................................................................................... 165

13

List of Tables

Table 2.1: Photocrosslinkable natural polymers used in cartilage tissue engineering.

.................................................................................................................................... 65

Table 3.1: Effect of reaction time on the modification degree of alginate methacrylate.

.................................................................................................................................... 85

Table 3.2: Rheological constants for solutions containing 15 mL of methacrylate and

different alginate concentrations, 24 h. ...................................................................... 88

Table 3.3: Rheological constants for solutions containing 15 mL of methacrylate and

different alginate concentrations, 8h. ......................................................................... 88

Table 4.1: Effect of reaction time on the degree of modification. .......................... 104

Table 5.1: Effect of reaction time on the degree of functionalization..................... 114

Table 5.2: Power law coefficients for different reaction time and GelMA

concentration. ........................................................................................................... 116

Table 6.1: Degree of functionalization for all functionalized group polymers. ...... 136

Table 6.2: Showing the hydrogel composition, experimental density, porosity, and the

average pore size. ..................................................................................................... 142

Table 7.1: Temporal variation of water contact angles for treated and untreated

scaffolds. .................................................................................................................. 165

14

List of Abbreviations

1D One dimension

2D Two dimensions

2PP Two-photon polymerisation

3D Three dimensions

3DP Three-dimensional printing

ASTM American Society for Testing and Materials

BASCs Brown adipose-derived stem cells

CAD Computer aided design

CaP Calcium phosphates

CGCaP Collagen–glycosaminoglycan

DAPI 4,6-diamidino-2-phenylindole

DEF Diethyl fumarate

ECM Extracellular matrix

FBS Fetal bovine serum

GelMA Gelatin methacrylate

hADSCs Human adipose-derived stem cells

hBMSCs Human bone marrow mesenchymal stem

cells

HCAECs Human coronary artery endothelial cells

hMSCs Human mesenchymal stem cells

hUVECs Human umbilical vein endothelial cells

hUVSMCs Human umbilical vein smooth muscle cells

IPN Interpenetrating Polymer Networks

MC Methylcellulose

MeCol Methacrylate collagen

MRI Magnetic resonance imaging

MSCs Mesenchymal stem cells

PBS Phosphate buffered saline

PBT Poly(butylenes terephthalate)

PCL Poly-ε-caprolactone

TGA Thermal gravimetric analysis

TGF Transforming growth factor

TIPS Thermally induced phase separation

15

TPU Thermoplastic polyurethane

UV Ultraviolet Light

VEGF Vascular endothelial growth factor

ACAN Aggrecan gene

ACI Autologous chondrocyte implantation

AlgMA Alginate methacrylate

DMEM Dulbello’s modified eagle medium

DMMB DMMB Dimethylmethylene blue

DNA Deoxyribonucleic acid

ECM Extracellular matrix

GAG Glycosaminoglycan

GelMA Gelatin-methacrylate

NMR (1H NMR) Nuclear magnetic resonance (Proton NMR)

FGF Fibroblast growth factor

HA Hyaluronic acid

HA-MA Hyaluronic acid methacrylate

Hep-MA Heparin methacrylate

KS Keratan sulfate

MAAh Methacrylic anhydride

MMP Matrix metalloproteinase

MSC Mesenchymal stromal cell

OA Osteoarthritis/osteoarthritic

PABS Plasma-Assisted Bio-extrusion System

PCL Polycaprolactone

16

List of Notations

�� Shear rate (s-1)

G' Storage modulus (Pa)

G'' Loss modulus (Pa)

ɳ Flow coefficient or consistency (Pa sn)

tanδ Loss factor

η Apparent viscosity (Pa∙s)

τ Shear stress (Pa)

17

Abstract

Articular cartilage is a load-bearing tissue that covers the ends of bone joints, acting

as a low-friction bearing surface and a mechanical damper for the bones. Articular

cartilage is a matrix-rich tissue with specialized cells (chondrocytes) that maintain the

structural and functional integrity of the extracellular matrix (ECM). An important

feature of cartilage is lacking self-repair ability and as a consequence, the zonal

structure and cartilage functions are often irreversibly lost following trauma and

disease. As a result, cartilage defects are prone to develop into predominant cartilage

disease osteoarthritis (OA), characterized by loss of cartilage, pain and debilitation.

The most common treatment for OA is joint replacement surgery, in which damaged

joint components are replaced by artificial prosthesis. This typically reduces the pain

and increases the mobility. However, over time such implants often fail and require

revision surgeries.

Tissue engineering has emerged as an alternative strategy to overcome the limitations

of traditional therapies by replacing the damaged part of cartilage tissue with a

biofabricated cartilage tissue for long-lasting periods and holds great promise as an

effective treatment for cartilage repair. However, most of the approaches in this field

focus on the use of single material structures, not able to mimic the real structure of

cartilage. This research project focus on the design and fabrication of novel multi-

material cartilage replacement by using two different hydrogels (alginate, gelatin) and

a hybrid system based on the combination of both alginate and gelatin. These

materials were successfully functionalized with methacrylate anhydride allowing them

to be processed through UV photopolymerization. The effect of polymer

concentration, methacrylated anhydride concentration and functionalization reaction

time was investigated allowing to tailor rheological and viscoelastic properties. Pre-

polymerized polymeric solutions were also prepared considering different

concentrations of photoinitiators. Mechanical, swelling and degradation

characteristics of these different systems were determined allowing to identify the

most suitable compositions for each single polymer system. Additionally, the ability

of each system to be used as a cartilage bioink was further investigated through

live/dead assay using human chondrocyte and mesenchymal stem cells. Finally, the

novel hybrid system proposed by this research was further investigated in terms of its

18

composition (ratio between gelatin and alginate) and a detailed biological study was

preformed not only to investigate the ability of this system to support cell growth and

cell-cell network formation but also the ECM formation through the quantification of

GAGs, Collagen and aggrecan. Results show that the developed systems can be used

as bioinks, their properties can be easily tailored and that the novel hybrid system,

which overcomes the main limitations of each individual system, is a promising bioink

material to support cartilage formation.

19

Declaration

I declare that no portion of the work referred to in the dissertation has been submitted

in support of an application for another degree or qualification of this university or

any other university or other institute of learning.

Hussein Mishbak

2019

20

Copyright

i. The author of this thesis (including any appendices and/or schedules to this

thesis) owns certain copyright or related rights in it (the “Copyright") and

she/he has given The University of Manchester certain rights to use such

Copyright, including for administrative purposes.

ii. Copies of this thesis, either in full or in extracts and whether in hard or

electronic copy, may be made only in accordance with the Copyright.

Designs and Patents Act 1988 (as amend) and regulations issued under it

or, where appropriate, in accordance with licensing agreements which the

University has from time to tie. This age must form part of such copies

made.

iii. The owner of certain copyright, patents, designs trademarks and other

intellectual property (the \Intellectual Property") and any reproductions of

copyright works in the thesis, for example graphs and tables

(\Reproductions"), which may be described in this thesis, may not be

owned by the author and may be owned by third parties. Such Intellectual

Property and Reproductions cannot and must not be made available for use

without the prior written permission of the owner(s) of the relevant

Intellectual Property and/or Reproductions.

iv. Further information on the conditions under which disclosure, publication

and commercialisation of this thesis, the Copyright and any Intellectual

Property and /or Reproductions described in it may take place is available

in the University IP Policy (see

http://documents.manchester.ac.uk/DocuInfo.aspx?DocID=487 ), in any

relevant Thesis restriction declarations deposited in the University Library.

The University Library’s regulations (see

http://www.manchester,ac,uk/library/aboutus/regulations) and in the

University’s policy on presentation of Theses.

http://documents.manchester.ac.uk/DocuInfo.aspx?DocID=487

21

Acknowledgements

First and foremost, I would like to express my gratitude to my advisor and supervisor

Prof. Paulo Jorge Bartolo. He has guide me through this PhD journey with lots of

support and patience. Taught me, both consciously and unconsciously, how great

experimental work is done and most importantly, how to take advantages of any

challenge that I experienced. In brief, I would consider myself very lucky to have such

a supervisor, not only for his tremendous academic support, but also for giving me so

many wonderful opportunities. Secondly, Special thanks goes to my co-supervisor,

Dr. Glen Cooper for his great support and encouragement. I was so fortunate to receive

their guidance, endless support and encouragement for all these years.

I’m also grateful to Dr. Carl Diver, Dr Liu, Dr Bo, Dr David and Ms. Svetlana

Kuimova for their support throughout these years

I would like to thank all members of staff at the University of Manchester for providing

me with a perfect working environment all these years. Special thanks to the

Manufacturing Centre in the School of MACE and those fantastic technicians: Eddie,

Kevin, Bill, Ross, Tony, Rus, Marie and Jack, and to the managers Stuart and Phil.

Special thanks goes to Stuart Mores, Maria and Louise from Materials School.

I am grateful to all colleagues in my research group: Cian, Salam, Enes, Mohammed,

Vag, Alex Pinar and Tom. Their daily support, friendship and our discussion helped

me to move forward. My time at Manchester was made enjoyable in large part because

of you, thanks for being part of my life. Big thanks to all my friends, old and new, who

have made the last few years so enjoyable.

I would like to acknowledge the support of the Government of Iraq for supporting this

PhD through a grant provided by Higher Committee for Development Education Iraq

(HCED). This work was also partially performed within the framework of the

SKELGEN project – Establishment of a cross continent consortium for enhancing

regenerative medicine in skeletal tissues (Marie Curie Action, International Research

Staff Exchange Scheme (IRSES) – Project reference: 318553).

Last but not least, I would like to acknowledge from the bottom of my heart the endless

love and support of small family, my lovely supportive wife (Najlaa) who was always

in my side in the good and bad days, my children (Bilal and Layaan). Special thanks

to my parents who supported me through my whole life during this journey since the

22

primary school till now. My brothers and Sisters and all my big family. They all kept

me moving forward and without them it would be very difficult to successfully

complete my PhD.

23

وثمرة إلحبيبة ، زوجتيإلغاليين وإلدي لى إ

بلال و ليان ولاديأ فؤإدي

24

Graphical abstract illustrating this research work focusing on material preparation and

characterization, and bioprinting of multi-material cell laden constructs for cartilage

tissue engineering.

25

Chapter 1 Introduction

26

1.1 Background

Millions of people around the world have experienced cartilage disease through

accidents, such as a tear to the anterior cruciate ligament (ACL), trauma, joint injuries

and abnormal joint loading, or due to ageing such as, degenerative joint diseases, that

can develop into osteoarthritis. Osteoarthritis (OA) is the most common disease

affecting cartilage tissue (1). When the hyaline cartilage is completely damaged, this

means that the joint surface may no longer be smooth to facilitate bone movement.

Despite new therapies have been developed to treat cartilage diseases, articular

cartilage repair is still a major clinical challenge. This is particularly critical since

cartilage presents very limited capacity to naturally heal due to its avascular nature

and relatively low cellular mitotic activity. Moreover, there are no available therapies

for long term treatment of articular cartilage damages. This is a very active field of

research and new therapeutic techniques are being developed not only to treat

damaged or diseased joint cartilage tissue but also to achieve and restore functional

normal cartilage that will give long-lasting improvements and allow patients to return

to a fully active lifestyle. Particularly relevant to this field is the concept of tissue

engineering (TE), combining the knowledge of engineers, biologists, material

scientists, and clinicians towards the development of biological substitutes that restore,

maintain, or improve the function of tissues (2). For cartilage applications, the

combined used of biodegradable and biocompatible materials, cells and growth factors

and the use of additive biomanufacturing systems, present high potential to create

customised cartilage replacements. In this case, three-dimensional biocompatible and

biodegradable structures (hydrogel-based scaffolds) are normally used and designed

considering the following requirements (3):

• The designed structure must allow cells to accommodate, proliferate and

differentiate.

• Support the metabolic activity of cells serving as a template for cells to

produce their own extra-cellular matrix (ECM);

• Provide an adequate mechanical and biological environment promoting tissue

regeneration.

In this context, the aim of hydrogel-based 3D constructs is to direct cartilage cells

(chondrocytes) to grow, mature, and build their own ECM mimicking the native

cartilage tissue (4, 5). Besides chondrocytes, non-differentiated stem cells are also

27

used together with specific growth factors (transforming growth factor beta

superfamily) to promote chondrogenesis (6).

Two approaches have been considered for cartilage tissue engineering, the use of

scaffolds seeded with cells or the use of bioinks (hydrogels encapsulating cells) (7).

Both approaches strongly depend on both the characteristics of the biomaterials being

considered and the fabrication processes. The scaffold-based approach is the most

commonly used but requires a pre-culture phase in a bioreactor before implantation,

while the fabrication of cell laden constructs allow the development of in situ printing

strategies (5).

The use of additive manufacturing for cartilage applications is a relatively new

approach. Most of the studies focused on the use of extrusion-based or pressure-

assisted additive manufacturing systems with or without cells (8, 9). Interesting results

were obtained in terms of the design of scaffolds (it is clear from the literature that

porous scaffolds for cartilage applications require small pores in comparison to bone

scaffolds) and cell differentiation (positive effect of hypoxia conditions) (9).

Additionally, most of these studies used hard materials (e.g. polycaprolactone) or very

soft materials (e.g. collagen), not able to completely simulate the mechanical

behaviour of cartilage and compromising the stability of any constructs. Besides this,

up till now, very few studies investigated the combined used of different materials to

create a structure mimicking cartilage. Naturally derived polymers such as alginate,

agarose, gelatin and collagen have been investigated as a single material system. Most

of these polymers are non-covalently crosslinked or ionically crosslinked, which could

be considered a major challenge since the produced hydrogels are weak and the gelling

conditions complex. Hybrid systems combining naturally derived and synthetic

polymers have also been investigated to overcome some mechanical and degradation

limitations of natural polymers (10, 11). Hybrid systems based on natural polymers

with tailored properties represents a novel area of research.

1.2 Research aim and objectives

The common problem with all available clinical strategies for cartilage repair is their

limitation to fully repair damaged cartilage providing long-term treatment. Usually,

they are expensive strategies and not appropriated to be applied in children. Thus,

tissue engineering is an alternative approach to overcome the problems of traditional

28

therapies by replacing the damaged part of cartilage tissue with a biofabricated

engineered cartilage tissue for long-lasting periods of time.

The aim of this research project is to design and evaluate a novel cell-laden multi-

hydrogel able to provide a temporospatial environment that could form

functional cartilage. This cell-laden, produced using additive manufacturing, is based

on biodegradable and biocompatible hydrogels resembling the structure of cartilage,

with tuneable mechanical and biodegradability properties able to support cells,

promote cell-cell interactions and the formation of ECM and processed under mild

and physiological conditions.

This construct is based on two different polymer materials (alginate and gelatin), one

for the superficial zone and the other for the deeper region of cartilage) and a hybrid

system obtained by mixing both alginate and gelatin for the transitional zone. Alginate

and gelatin were selected due to their biocompatibility, biodegradability properties,

easy to process under mild conditions and low cost.

Both alginate and gelatin were functionalized with methacrylate anhydride (MA) to

introduce photoreactive sites into the polymer backbone to produce

photocrosslinkable constructs. Ultraviolet (UV) photopolymerization was selected to

induce rapid gelation, allowing to control the polymeric network properties such as

crosslinking density, matrix stiffness and cell encapsulation by tuning the

photopolymerization parameters (e.g. light intensity, wavelength, photinitiator and

polymer concentration, and exposure time). Key research challenges considered to

design the cartilage constructs are related to the mechanical properties of produced

constructs, photoinitiator concentration, biological effect of the unsaturated degree,

cell exposure to UV radiation and cytotoxicity of both photoinitiators and unreacted

polymers.

To achieve the research aim, the following research objectives were considered:

✓ Objective 1: understand cartilage structure and its role, main clinical problems

and treatment strategies; investigate current biofabrication technologies,

suitable materials and design requirements to produce a bioink for cartilage

regeneration;

✓ Objective 2: select and functionalize polymeric materials and develop a

suitable approach for cartilage regeneration by investigating two different

hydrogels and the combination of both polymers into a hybrid system able to

functionally mimic the articular cartilage structure, composition and

29

properties. Variations in polymer concentration, methacrylate anhydride

concentration and degree of substitution can provide hydrogels with tunable

and controllable properties;

✓ Objective 3: establish a correlation between process conditions and

morphology, mechanical and biological properties of 3D polymeric constructs

and the capacity of producing a functional engineered cartilage replacement;

✓ Objective 4: quantify the optimal composition, conditions, biomechanical,

biophysical and biochemical properties of tailored polymeric constructs for

better ECM matrix production.

✓ Objective 5: develop a new tissue engineering approach to enable the

neocartilage formation fully anchored to the underlying subchondral bone.

In order to achieve the research aims and objectives four main research phases were

considered:

❖ Material design (phase 1) – Selected polymeric systems must be

functionalized in order to be photopolymerizable. Functionalization

corresponds to the introduction of unsaturations (C=C sites) into the hydrogel

structure through the use of methacrylate. Solutions containing different

concentrations of alginate or gelatin mixed with different concentrations of

methacrylate during different reaction times were prepared. Proton nuclear

magnetic resonance 1HNMR is used to assess the equivalent unsaturation level

and to confirm the functionalization process. The level of unsaturation has an

impact on the processability of these materials and crosslinked density, which

determines the mechanical properties and free volume available for cell

encapsulation, viability and proliferation.

❖ Photopolymerization process (phase 3) – Low light intensity values and low

concentrations of free-radical photoinitiators are considered. Photocurable

polymers prepared in phase 1 are mixed with different concentrations of

photoinitiator and polymerised at different light intensity values, by changing

the distance between the material and the light source. The curing kinetics and

mechanical properties are assessed as a function of light intensity and

photoinitiator concentration.

❖ Printing optimisation (phase 2) – The rheological and viscoelastic properties

of the prepared solutions are critical. High viscous materials and Newtonian

formulations require high pressures during the printing process, which might

30

induce significant shear stresses on encapsulating cells. During this phase a

general understanding on the effect of the material design characteristics on

the rheological properties is obtained.

❖ Biological assessment (phase 4) – Human chondrocytes and/or human

mesenchymal stem cells (hMSC) were encapsulated on those three different

bioink systems (alginate, gelatin and hybrid systems), before

photopolymerization. The production of glycosaminoglycan, Collagen

secretion, and the formation of cell networks were studied. Histological

analysis is used to semi quantify the gene expression in the deposited ECM in

the hydrogel constructs.

1.3 Thesis outlines

This thesis, based on published and submitted papers and book chapters, comprises

eight Chapters, which progress in accordance with the identified research objectives.

Besides this introductory Chapter, which presents the current research context,

research aims and objectives, the remaining chapters can be summarized as follows:

Chapter II: State-of-art in cartilage tissue engineering;

This Chapter, focusing on objective 1, comprises two sections. First, it overviews the

problem being addressed by this research, describing the cartilage structure and

composition, main clinical problems, and most commonly used clinical approaches. It

discusses also the concept of tissue engineering, presenting the key requirements of

cartilage tissue engineering constructs. The second Section provides an overview on

recent advances in the design and engineering of naturally derived photocrosslinkable

hydrogels for cartilage tissue engineering. Hydrogel properties, synthesis routes,

crosslinking methods, and strategies for hydrogel bio-functionalization are described

and future research perspectives discussed.

Chapter III: Development and characterisation of a photocurable alginate bioink

for 3D bioprinting

This Chapter, focusing on the research objectives 2 and 3, investigates the preparation

of alginate-based systems for UV bioprinting applications. The material is

characterised before and after the polymerisation (curing) process and the effect of

both functionalization time and photoinitiator concentration on the rheological,

mechanical, morphological, swelling and degradation properties are investigated.

31

Chapter IV: Photocurable alginate bioink development for cartilage replacement

bioprinting

This Chapter, focusing on the research objectives 2 and 3, investigates the use of the

photocurable alginate presented in Chapter 3 as a carrier of chondrocyte cells for

cartilage repair. Different methacrylate concentration and reaction times are

considered and their effects on the degree of substitution, rheological and biological

properties investigated. Alginate methacrylate cell laden constructs were produced,

and preliminary cytotoxicity and cell proliferation tests performed using the live/dead

assay.

Chapter V: Photocrosslinkable gelatin methacrylate (GelMA) hydrogels:

towards 3D cell culture and bioprinting for cartilage applications

This Chapter, focusing on the research objectives 2 and 3, investigates the use of

photocurable gelatin to encapsulate human adipose-derived stem cells. Gelatin is

functionalised and the level of functionalisation is investigated. Material preparation

and characterisation is discussed. The influence of the functionalisation time, polymer

concentration, degree of modification and viscoelastic behaviour on the mechanical

strength and cell viability is also discussed.

Chapter VI: Development and characterisation of a photocurable alginate-

gelatin bioink for cartilage applications

This Chapter, addressing the research objectives 2 and 4, investigate a novel hybrid

alginate/gelatin photocrosslinking hydrogel system. The synthesis, morphological,

rheological, mechanical and biological characterisation is detailed.

Chapter VII: Hybrid PCL/Hydrogel scaffold fabrication and in-process plasma

treatment using PABS;

This Chapter, addressing the research objectives 4 and 5, investigates the use of a novel

plasma-assisted bio-extrusion system (PABS), to produce multi-material scaffolds for

the interface between the subchondral bone and cartilage, consisting in a synthetic

biopolymer (polycaprolactone) and a hybrid natural hydrogel system (previously

described in Chapter 6). The Plasma-Assisted Bio-extrusion System (PABS),

developed at the University of Manchester, is also used for the scaffold surface

modification (N2 plasma modification).

Chapter VIII: Conclusions and future work.

32

This Chapter provides an overall summary of the thesis and the main conclusions that

can be drawn from the research work carried out. Possible future directions for

research following on from this thesis are also presented.

The structure of this thesis is illustrated in Figure 1.1

Figure 1.1: Framework structure of this thesis.

33

Chapter 2

State-of-the- art of

Cartilage Tissue

Engineering

34

This Chapter, which provides the current state-of-the-art on cartilage tissue

engineering comprises two main sections. The first section overviews the problem

being addressed by this research, describing the cartilage structure and composition,

and the most commonly used clinical approaches presenting their main advantages

and limitations. It discusses also the concept of tissue engineering, presenting the key

requirements of cartilage tissue engineering constructs. The second section provides

an overview on recent advances in the design and engineering of naturally derived

photocrosslinkable hydrogels for cartilage tissue engineering. Hydrogel properties,

synthesis routes, crosslinking methods, and strategies for hydrogel

biofunctionalisation are described and future research perspectives discussed.

Therefore, these two Sections covered the two main cartilage tissue engineering

approaches, scaffold-based and cell laden (bioprinting), covered by this research.

Biological perspectives and current strategies in osteochondral

tissue engineering †

Articular cartilage and the underlying subchondral bone are crucial in human

movement and when damaged through disease or trauma results in a severe impact on

quality of life. Cartilage has a limited capacity to regenerate due to the avascular

composition of the tissue and limited efficacy of current therapeutic interventions.

Worldwide the numbers of patients requiring therapy for osteochondral indications is

rising due to an ageing population, thus increasing the pressure on healthcare systems.

Subsequently, research into new therapies especially using the principles of tissue

engineering is increasing. However, a basic understanding of the biological

composition, structure, and mechanical properties is required. Furthermore,

consideration of material design, architecture, biomanufacturing strategies, and the

role of experimental standardisation is needed to help develop tissue engineering

therapies that can be successfully translated into the clinic.

This Section provides an overview of the biological properties of osteochondral tissue

and the role of tissue engineering principles in developing new therapies is provided.

† This Section is based on the paper: Hussein Mishbak, Cian Vyas, Glen Cooper, Ruben Pereira, Chris

Peach, Paulo Bartolo, “Osteochondral tissue: a biological and mechanical review for tissue

engineering”, Biomanufacturing Reviews, Springer, in Press.

35

Introduction

Osteochondral tissue is composed of articular cartilage, a specialized tissue that covers

the distal ends of the bones in articulating joints, and the subchondral bone which

anchors the cartilage to the underlying bone (12-16). Articular cartilage has a highly

flexible and lubricated surface to reduce frictional forces during movement and

facilitate smooth articulation. The tissue enables the transmission of mechanical loads

from movement to the skeleton (17-20). Osteochondral tissue is composed of distinct

regions with articular cartilage, comprising the majority of the structure, and an

underlying subchondral bone phase.

Articular cartilage is avascular and aneural with low metabolic activity and thus

trauma or diseases (e.g. osteoarthritis and rheumatoid arthritis) are difficult to treat

due to the inherent inability of articular cartilage to self-regenerate in comparison to

the greater healing capacity of bone (Figure 2.1) (20). A thorough understanding of

the composition and structure of the tissue will enable superior tissue engineering

approaches to be explored to solve the clinical challenges of osteochondral defects.

Current clinical approaches, typically palliative, are ineffective and the tissue either

continues to degrade, resulting in total replacement with an implant, or fibrocartilage

is formed. Tissue engineering approaches have been widely explored to develop new

approaches to repair and regenerate the tissue. However, to date these have been

unsuccessful in producing mature articular cartilage. The current fundamental

approaches based on scaffolds (acellular and cellular), biomaterials, and cells may

need to be revaluated to understand why these approaches consistently fail in

producing a relatively simple and thin tissue, although exhibiting a highly complex

hierarchical organisation (21). Adult articular cartilage tissue takes over 20 years to

mature and the tissue generated once produced does not turnover or regenerate unlike

other tissues in the body. The cartilage tissue produced during childhood is the same

throughout a person’s lifetime. Malda et al. suggest that new approaches must

appreciate this fundamental aspect of cartilage tissue physiology which may require

the incorporation of permeant non-degradable structures to support the major

mechanical properties required of the tissue and the restoration of the

microenvironment (cytokines, growth factors, and mechanical loading profiles) in the

early stages of development (fetal and childhood) to recapitulate the developmental

processes which generate the correct articular cartilage tissue (21). Furthermore,

elucidating the underlying processes of cartilage development will provide a

36

mechanistic understanding that can be utilised in tissue engineering approaches for

cartilage regeneration.

Figure 2.1: Comparison between the physiology and healing capacity of bone and

cartilage. Cartilage is avascular tissue so lacks ready access to a supply of circulating

stem cells and nutrients so relies on the synovial fluid for nourishment. This combined

with its largely acellular composition and low metabolic activity results in a nearly

complete lack of innate regenerative capacity. Consequentially, defects fail to heal,

and clinical interventions typically result in the formation of inferior fibrocartilage. In

contrast, bone has a greater ability for self-repair due to the constant tissue remodelling

by osteoblasts and osteoclasts. Bone is also a highly vascularised tissue which allows

a ready supply of nutrients and proteins that stimulate bone repair. Furthermore, there

is a large source of stem cells in the bone marrow and periosteum which are able to

differentiate into bone producing cells. This allows bone to heal defects up to a certain

critical size after which vascularisation becomes an issue. Image from reference (20).

Composition of articular cartilage

There are three forms of cartilage in the human body which are fibrocartilage (22),

elastic cartilage, and hyaline cartilage each with their own specific biological,

mechanical, and structural properties (23). Hyaline cartilage is a thin connective tissue

present at synovial joints such as the knee, elbow, shoulder, and hip where it covers

the bearing surface of the underlying bone and is termed articular cartilage (24). Thus,

37

hyaline cartilage provides an efficient load bearing surface that has a low friction

coefficient thus lubricating the movement of joints and can support load transfer of up

to six times the human body weight in the knees (17). The complex, nonlinear,

viscoelastic, anisotropic, and heterogeneous structure and composition of cartilage

enable these vital properties (25, 26).

Adult articular cartilage consists of predominately extracellular matrix (ECM),

approximately 95–99% of the total volume, which itself consists of 80% water and

20% solid contents (27). The solid content is mostly comprised of collagens (50-75%),

proteoglycans (15-30%), and a small amount of non-collagen proteins (24, 28). Cells

account for a small percentage of the total volume (1-5%) and consist of only a single

cell type, chondrocytes (29). The composition of articular cartilage changes as the

tissue matures from initial formation during embryogenesis to final maturation (18-21

years old) (30).

The mature tissue has a low density of chondrocytes in a low proliferative and

metabolic state, and are isolated from each other, so lack cell-cell interactions due to

being encapsulated in the pericellular matrix (19). This is partly responsible for the

low healing capacity of mature articular cartilage. However, the mature chondrocyte

has an important role in the tissue homeostasis by coordinating and producing the

ECM components. The morphology, orientation, and phenotypic expression of

chondrocytes are depth and biomechanically dependent as the cells are influenced

through mechanotransduction (31, 32). This results in the wide range of morphologies

observed, which range from rounded, elongated, flattened, and hypertrophic, all within

the same tissue (33).

The collagen network in articular cartilage is highly organised and primarily

composed of collagen type II (~50% dry weight of articular cartilage) (25). The

organisation is depth dependent in the tissue and is partly responsible for the

biomechanics, especially the tensile and compressive properties (26). The fibre

diameter increases from the articular surface through the depth of the tissue

(superficial zone ~ 55 nm, middle zone ~ 87 nm, and deep zone ~ 108 nm) (34).

Although collagen type II is the main collagen (95% of total collagen), there are other

collagens present such as type I, II, VI, IX, X, and XI. Collagen type I is present in

small amounts only in the superficial zone but can be found abundantly in

fibrocartilage so can be used as a useful indicator of fibrocartilage formation (35-37).

Subsequently, the ratio between collagen I and collagen II can be used as a marker to

38

assess the status of the cartilage tissue as chondrocytes cultured in vitro monolayer

express higher levels of collagen type I indicating that the chondrocytes have

undergone dedifferentiation (32). Collagen IX and XI are found throughout the tissue

in small amounts and are involved in crosslinking between fibrils, regulation of fibril

size, and interactions with other biomolecules (25). Collagen X is found in the deep

and calcified zones and is believed to have a role in the mineralisation between the

cartilage and the subchondral bone (38, 39). A major non-collagenous component of

the ECM is proteoglycans and glycosaminoglycan’s (GAGs) (17). Proteoglycans

consist of a core protein that is heavily bound with covalently attached polysaccharide

chains, GAGs. Aggrecan is a main proteoglycan within articular cartilage and is

covalently attached to negatively charged GAGs, keratan sulphate and chondroitin

sulphate. This is able to bind multiple times to a hyaluronic acid backbone to form

aggrecan-hyaluronan aggregates which are highly negatively charged (40, 41). The

negative charge on this proteoglycan aggregate, and other proteoglycans, causes an

osmotic pressure to be generated as water is taken in and entrapped which causes

swelling (38). This turgidity produced by the network of proteoglycans and combined

with the structural confinement caused by the organisation of collagen results in a high

compressive modulus (14, 32). Subsequently, as the concentration of proteoglycans

increases with depth in articular cartilage towards the subchondral bone region, the

water content and swelling pressure rises, thus the compressive modulus of the tissue

increases (42-45). This enables articular cartilage to have a high mechanical load

bearing capability which can transfer and distribute loads effectively (46).

Structure of osteochondral tissue

Articular cartilage has a hierarchical organisation from the nanoscale to the macroscale

with a distinct zonal structure. Each zone has its own specific ECM composition,

biomolecule orientation, chondrocyte shape and organisation (47), imparting specific

biomechanical characteristics (Figures 2.2 a, b and c). These zones are termed,

descending from the articulating surface, the superficial or tangential zone, the middle

or transitional zone, the deep or radial zone, and the calcified zone, before the articular

cartilage gives way to the subchondral bone region. Furthermore, there is a microscale

radial organisation surrounding the chondrocyte termed the chondron which has a

unique composition depending on the distance from the chondrocyte (Figure. 2.2b).

39

Zonal structure of articular cartilage

On the articular surface and above the superficial zone is a thin layer, ranging from a

few hundred nanometres to a micrometre, termed the lamina splendens which is

acellular and composed of proteins (48). Although the structural role is unclear it is

thought that the gradual build-up of proteins from the synovial fluid acts as a protective

and low friction interface for the articular cartilage surface (35).

Immediately below the lamina splendens is the superficial zone (10-20% of cartilage

thickness) which comprises small diameter collagen fibres (predominately type II and

IX) that are organised parallel to the articular surface and densely packed. This allows

a low coefficient of friction, which enables smooth movement of the joint and imparts

the ability to withstand both the high tensile and shear stresses that the articular

cartilage encounters under loading. The chondrocytes are densely packed and arrange

themselves along the collagen fibres parallel to the surface, displaying a flattened

morphology. Chondrocytes also secrete proteins such as superficial zone protein (SPZ,

also known as lubricin or PGR 4) and collagen I which act as lubricants (25, 49-51).

Figure 2.2: Hierarchical and graded composition, structure, and properties of

osteochondral tissue. a) Procollagen polypeptides bind together to form collagen that

assemble into microfibrils and then fibrils which can be crosslinked together by

collagen type IX. Collagen fibrils exhibit a characteristic banding pattern of ~67 nm

due to the staggered packing arrangement of collagen. Image from reference (36). b)

Molecular composition and arrangement of the chondron depicting the pericellular,

territorial, and interterritorial matrix with increasing distance from the chondrocyte.

Image from reference (37). c) Zonal structure and properties of osteochondral tissue.

Image adapted from (52).

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SPZ is a potential marker to identify this zone. The amount of proteoglycans is low

compared to the other zones which increases the permeability, thus resulting in

compressive strains of up to 50% and high fluid flow which influences the

compressive properties of the entire cartilage tissue (25, 26, 40, 52).

Below the superficial zone is the middle zone (40–60% of cartilage thickness), which

is predominately composed of collagen II fibres randomly arranged and displaying a

larger diameter than the superficial zone. Chondrocytes are present at a lower density,

have a rounded morphology, and also express large quantities of collagen II and

aggrecan. A marker for this zone is the cartilage intermediate layer protein which is

expressed throughout this zone (39). This zone also has the highest concentration of

proteoglycans especially aggrecan (52, 53).

The deep zone (20-50% of cartilage thickness) is comprised of the largest diameter

collagen fibres that are oriented perpendicular to the subchondral bone region with the

chondrocytes organised along the collagen fibres in columns with an elongated

morphology. The cells themselves are present in lower density compared to the other

zones and express lower levels of collagen II (53, 54).

The final zone before the subchondral bone region is the calcified zone which is

distinguished from the deep zone by the presence of a tidemark which demarks the

boundary between calcified and non-calcified regions (55, 56). The zone anchors the

collagen fibres of the deep zone to the subchondral bone thus integrating the cartilage

to the underlying bone. This also provides an interface between the hard phase of bone

and the soft phase of cartilage since the presence of hydroxyapatite reduces the

mechanical gradient between the phases (25, 41, 57). The calcified zone also has the

highest number of chondrocytes which are in a hypertrophic state.

The final zone of the osteochondral tissue is the subchondral bone which lies directly

below the calcified zone and separates the articular cartilage from the bone marrow.

This zone consists of the bony lamella (cortical endplate) and the subarticular

spongiosa, the supporting trabeculae and bone components, which is separated from

the calcified zone of articular cartilage by a cement line (58, 59). The subchondral

bone differs markedly in composition and structure to the articular cartilage. The

subchondral trabeculae are highly vascularised which enables transportation of

nutrients, gases, and waste through channels which cross the subchondral bone plate

and enter the calcified zone, which is entirely reliant on the synovial fluid as a source

of nutrients if these channels are absent (58, 59). The main collagen is type I due to

41

the tissue being mineralised bone. These collagen fibres do not cross between the

calcified cartilage and subchondral bone so does not act as an anchor as occurs with

collagen fibres that cross the tidemark and connect the non-calcified and calcified

cartilage. Furthermore, the subchondral trabecular structure and mechanical properties

are anisotropic and can dynamically remodel itself to respond to applied forces (60-

62). The main function of the subchondral bone is to maintain joint shape and provide

mechanical support since it has a high compressive modulus and is impermeable so is

able to stabilise the tissue and distribute the applied mechanical forces (25, 40, 58, 63,

64). Furthermore, the subchondral bone acts as a nutrient source for articular cartilage

which is not vascularised itself (43).

Current osteochondral treatments

The zonal structure of osteochondral tissue and the innate inability of articular

cartilage for self-regeneration poses a problem for clinical interventions.

Subsequently, a variety of both clinical treatments are available with different degrees

of success and tissue engineering approaches are currently in clinical trials (65, 66).

The type of treatment used will depend on the defect category, stage, size, and

location. Osteochondral defects are commonly classified by the Outerbridge

classification system which indicates the severity of a lesion (67). This system

classifies defects from grade I-IV, where a grade 0 defect is normal healthy articular

cartilage; grade I indicates swelling and softening of the tissue; grade II indicates a

partial thickness defect with a diameter less than 1.5 cm; grade III defect has a

diameter greater than 1.5 cm and presents as a full thickness lesion up to the

subchondral bone; and grade IV is a full thickness defect that exposes the subchondral

bone.

The main treatments currently used clinically are classified into (1) microfracture, (2)

autologous chondrocyte implantation (ACI), (3) matrix-induced autologous

chondrocyte implantation (MACI), and (4) osteochondral auto- and allografts as

shown in Figure 2.3 (18, 44, 68). These approaches mainly utilise in vivo methods to

repair and regenerate the tissue in situ.

Microfracture is a technique that has been used clinically since the 1980s to produce

fibrocartilage in articular cartilage defects (69, 70). The technique recruits MSCs from

the bone marrow by drilling into the subchondral bone. The defect area is debrided to

ensure a clean and stable margin before drilling into the subchondral bone. This

42

induces bleeding into the defect area which forms a fibrin clot that contains MSCs

which subsequently remodels into fibrocartilaginous tissue over a period of 12-16

months. This long period of remodelling requires a lengthy postoperative

rehabilitation period with limited mechanical loading (51, 71).

Figure 2.3: Current treatment options for cartilage regeneration. a) A full-thickness

chondral lesion (Grade III). b) Debridement of the lesion is used as precursor to more

invasive procedures as it removes damaged cartilage and bone to create a healthy

border which enables improved integration of neotissue with native tissue. In smaller

cartilage defects debridement may be used without further procedures as a minimally

invasive treatment option. c) Microfracture is treatment that drills into the subchondral

bone to create channels that allow MSCs to migrate into the defect. d) Autologous

chondrocyte implantation (ACI) uses a population of 12-48 million autologous

chondrocytes to which are inserted into the defect which is then covered with either a

periosteal patch or a collagen membrane. e) Matrix-induced autologous chondrocyte

implantation (MACI) is a development of the ACI technique. The autologous

chondrocytes are cultured in vitro and seeded onto an absorbable 3D scaffold (collagen

or hyaluronic acid) before being implanted into the defect and fixed to the defect with

fibrin glue. Image from reference (37).

Although considered the gold standard by the FDA and some clinicians, the long-term

outcomes for joint functionality using microfracture technique show limited

improvement (49, 50). This is due to the inferior biomechanical and biochemical

properties of fibrocartilage compared to hyaline cartilage, which creates a mismatch

between the native tissue and the neotissue (52, 72). The treatment provides a short-

term benefit to the patient but only postpones cartilage degeneration as the repair tissue

typically deteriorates approximately 18-24 months after surgery (52). Subsequently,

five years after surgery the likelihood of treatment failure is high irrespective of the

cartilage defect (49, 50). This technique can also be combined with other treatment

43

methods and an advancement on the technique utilises a collagen matrix that is

inserted into the defect to promote MSC differentiation into chondrocytes (53).

In the early 1990s a new technique called autologous chondrocytes implantation (ACI)

was developed as an improvement on the microfracture technique (18, 44, 54, 73).

The technique is a two-stage surgical procedure that first involves harvesting

autologous chondrocytes from a minimal load bearing region through a biopsy punch

and expanding these cells in vitro to obtain a population of approximately 12-48

million cells. In the second surgery, the defect area is debrided, and the cell suspension

is seeded into the defect area and confined to the defect by membrane coverage. This

membrane is typically a periosteal patch; however, a major cause of treatment failure

is hypertrophy of the patch (56). Subsequently, synthetic collagen or hyaluronic acid

patches have been utilised, showing reduced failure rates (5 to 26%) due to patch

hypertrophy (57). However, these synthetic patches are considered as off-label in the

USA due to the sourcing of the material from allogenic sources which may increase

the chance of an immune response (44). ACI has been shown to be effective through

clinical trials which demonstrated positive long-term functional and clinical outcomes.

This is potentially due to the use of autologous chondrocytes which are having a

greater inherent ability to form hyaline cartilage than MSCs. However, the technique

has limitations as fibrocartilaginous tissue was shown to develop in the majority of

patients. This may be a result of the in vitro culturing stage as studies have

demonstrated that chondrocytes dedifferentiate into fibro chondrocytes in 2D culture

however other studies have shown that by culturing the cells in a 3D and hypoxic

environment this can be reversed (58, 59). There are further limitations to ACI which

include the need for two invasive surgical procedures combined with in vitro culturing

and the subsequent long recovery period (6-12 months) needed to ensure successful

neotissue formation.

A derivative of the ACI technique is MACI, which can be considered as a tissue

engineering approach as it utilises a scaffold to assist in cell attachment, distribution,

proliferation, and to guide matrix formation. MACI is similar to ACI in that it requires

the initial isolation of autologous chondrocytes from the patient, in vitro cell expansion

and seeding onto the scaffold, which can be made of collagen or hyaluronic acid. The

cell-seeded scaffold is then cultured in vitro before implantation into the debrided

defect and fixation with fibrin glue. The clinical advantage of MACI over other

techniques remains to be confirmed as current clinical trial data has shown that MACI

44

either has similar or better functional results. MACI however has the same limitations

as ACI in that two surgical procedures are required, slow tissue maturation, and long

recovery periods. However, the long recovery periods may happen with any treatment

due to the nature of the limited regenerative capacity of articular cartilage. There are

a number of benefits to using a scaffold based treatment such as easier fitting of the

graft into the defect, improved graft stability, and better control in preventing

dedifferentiation of the chondrocytes due to fact that cells are cultured in a 3D matrix

which reduces the formation of fibrocartilaginous tissue (44, 60).

A further treatment option is the use of osteochondral auto- and allografts,

mosaicplasty, which involves the transplantation of osteochondral tissue from either

patient (autograft) or a cadaveric donor (allograft) (61). A biopsy is taken from a non-

loading or minimal loading region, in an autograft transplant, which also includes a

portion of the subchondral bone. This is then transplanted into the defect, which has

been prepared with only healthy tissue remaining, and the graft is then aligned with

the native tissue. Mosaicplasty has been shown to have better results than

microfracture however there are a number of limitations to the technique (49, 62, 63).

The use of autografts is limited due to the need to restrict donor site morbidity which

results in only small defect sizes (<4 cm2) being treated. Furthermore, the use of

allografts raises the issue of disease transmission and immunogenicity which is

potentially an issue, however the size of defect treated can be larger as the tissue is

derived from cadavers. Finally, graft failure has been observed in 15-55% of patients

after 10 years due to poor integration with the host tissue which is also a major

limitation with the other techniques as well (64).

Although there are a number of clinical approaches none as of yet have fully repaired

either an articular cartilage defect or a full osteochondral defect over the long-term.

All the approaches have significant limitations which fail to recapitulate the native

structure and result in mechanical mismatch with failure over the long-term.

Subsequently, new tissue engineering approaches are required that enable rapid weight

bearing neotissue that develops into hyaline cartilage and that is fully anchored to the

underlying subchondral bone to enable vertical integration.

Patients with clinical conditions causing cartilage loss present either discrete cartilage

loss in a joint surface or full thickness cartilage degeneration as shown in Figure 2.4

(74). Discrete areas of cartilage loss e.g. osteochondritis dissecans, often caused by

trauma affects the younger population, whereas full thickness cartilage loss in the

45

whole joint affect more elderly patients and is a result of a systemic disease such as

rheumatoid or osteoarthritis (75).

When considering discrete small areas of cartilage loss, patient present with pain due

to exposed bone and mechanical symptoms (e.g. joint locking), due to the disruption

in the smooth low friction joint surface.

Figure 2.4: a) Discrete cartilage defect; b) End stage arthritis; c) Full thickness area

of cartilage loss.

Current solutions haven’t addressed the goal of restoration of the hyaline cartilage

surfaces of the joint (76). The most common surgical procedure to address small areas

of cartilage loss is arthroscopic (keyhole) surgery, called debridement, which uses

mechanical shavers to remove debris from the joint and loose cartilage material which

solves some of the mechanical symptoms (77). The aim of the surgery is to remove all

loose material and roughen the surface of the exposed bone enough to allow new tissue

to adhere and form in the base.

Due to the unpredictable results from debridement techniques were then developed to

create more biological healing in the cartilage defect. Microfracture, popularised by

Steadman (70), involves making 3mm perforations in the subchondral bone, ensuring

that the structural integrity of that bone is not compromised. This allows a super clot

to form in the space creating an environment for pluripotential marrow cells to

differentiate into stable tissue. In one study biopsies after microfracture treatment

noted that 11% had formed predominantly hyaline cartilage and 17% a mixture of

fibrocartilage and hyaline cartilage within them (Figure 2.5). The remaining patients

formed either predominantly fibrocartilage or no tissue at all (78, 79).

46

Figure 2.5: a) Normal cartilage; b) Fibrocartilage next to hyaline after microfracture.

With continued poor results from microfracture due to poor differentiation of new

tissue, allografting techniques were developed. Mosaicplasty involves tacking small

osteochondral allograft plugs from the periphery of weightbearing joints and

transplanted into the defect. There are issues of donor site morbidity with this

technique and the lack of integration of the periphery of the plugs with each other or

the native hyaline cartilage (80).

Lastly, autologous chondrocyte implantation techniques have been in use since the

first reported series in 1987 (81). Cartilage is harvested and at a second surgery,

culture-expanded autologous chondrocyte cells are injected into the cartilage defect

underneath a patch of periosteum that prevents cell migration from the site. Clinical

results are promising in the early years following this treatment and hyaline-like tissue

has been identified in some specimens. However, the tissue is not morphologically or

histochemical identical to normal hyaline cartilage and fibrocartilage is often found in

samples.

These are promising techniques however, rarely last more than 5 years and results

often gradually deteriorate over this period. However, as no other superior alternatives

are available to clinicians, these procedures continue to represent the mainstay of

management despite their failure to recreate the hyaline cartilage that has been

damaged.

Whole joint osteoarthritis represents a different clinical challenge. Not only is the

articular cartilage damaged, but the subchondral bone becomes deformed and the

periarticular capsule, muscles and tendons become affected. Current treatment

strategies are dominated by joint replacement surgery (82). Since the development of

the successful low friction arthroplasty by the late Sir John Charnley, the metal on

47

high molecular weight polyethylene bearing surfaces has been widely used in total

joint prostheses. In the UK over 160,000 hip and knee replacements are carried out

each year (83). However, the challenge continues to overcome the mechanical

loosening of these prostheses over time. In addition, cellular regeneration strategies

are unlikely to overcome the structural changes in all of the tissues types that are

involved (84). Revision surgery to replace failed prostheses causes not only significant

morbidity to the patient but also the results of the second or subsequent joint

replacement being significantly inferior to primary procedures. There is an associated

substantial financial burden to the healthcare system due to increased cost of peri-

operative investigations, blood transfusions, surgical instrumentation, implants and

operating time, as well as an increased length of stay in hospital which accounts for

most of the actual costs associated with surgery. Innovation has been able to reduce

the incidence of prosthetic failure by improving the techniques employed to ensure

good integration with the host bone, e.g. coating implants with hydroxyapatite and

using uncemented implants.

Tissue engineering: requirements and challenges

Tissue engineering is a multidisciplinary field of research conducted to meet clear

clinical requirements of therapies to promote the regeneration and repair of diseased

and damaged tissues. Tissue engineering approaches in cartilage regeneration and

repair has great potential and provides an alternative to current available therapies

which are inadequate, however, challenges remain (18, 20, 44, 85, 86). Engineered

cartilage constructs are comprised of biomaterials, cells, and stimulatory factors (e.g.

growth factors and biomechanical stimulation), which are considered key to the design

of functional cartilage tissue (Figure 2.6) (87, 88). However, the current paradigm still

lacks in the development of long-term phenotypically stable articular cartilage tissue

which exhibits integration with the surrounding tissue, mechanical stability, and

withstands inflammatory factors, especially in a diseased environment such as

osteoarthritis.

48

Figure 2.6: Tissue engineering strategies for articular cartilage regeneration. Image

from reference (85).

2.1.6.1 Biomaterials design and selection

Biomaterials are the backbone of 3D engineered constructs and support tissue growth

and formation by providing a similar environment to the native tissue and structural

integrity during maturation to allow cell proliferation, cell to cell communication, and

ECM formation. The ideal design specifications of 3D tissue engineered scaffolds

from the biomaterials perspective include: i) biocompatibility, cell viability with a

desired cellular behaviour; ii) biodegradability, the scaffold degrades at a controlled

rate which matches tissue formation; iii) provides mechanical and biochemical cues to

promote a desired cellular response. An alternative strategy to the use of biomaterials

is cell self-assembly approaches such as spheroid formation which do not require

supporting materials.

Biomaterial scaffolds can be produced from different sources such as naturally derived

polymers, synthetic polymers and ECM derived materials.

Natural derived biopolymers have been explored extensively for cartilage tissue

engineering applications (86, 89, 90). The most associated advantages with naturally

derived polymer are their capacity of supporting cell attachment, viability,

proliferation, attachment, and differentiation, and, in some cases, maintenance of cell

phenotype (91). This is mediated in protein derived biopolymers through binding

motifs present in the polymer. Despite these key advantages, naturally derived

polymeric materials present significant limitations such as poor degradation kinetics

49

and mechanical properties, however, these limitations may be improved through the

modification of the backbone of the polymer chain or via crosslinking mechanisms

(91).

Naturally derived polymeric materials can be classified into two main categories:

1- Polysaccharides: gel forming polysaccharides such as alginic acid and

mucopolysaccharides (glycosaminoglycans), storage polysaccharides which include

starch and glycogen, and structural polysaccharides such as cellulose and chitin.

2- Protein based polymeric material composed of amino acid groups such as

collagen, gelatin, and silk fibroin. Proteins can be classified by their shape, size,

solubility, composition, and function.

Synthetic polymers have shown potential in tissue engineering due to their improved

mechanical and degradation properties with the capacity to be more easily chemically

modified or engineered to tune their properties (92-94). The hydrolytic and enzymatic

degradation of the polymer can be controlled through modification of the polymer

(95). However, due to lack of biologically functional domains, which can reduce the

risk of immune response, synthetic polymers may not facilitate cell phenotype

expression or cell attachment as occurs in naturally derived protein-based polymers.

Decellularized extracellular matrix (dECM) based biomaterials have also been

explored to create 3D constructs. The native ECM is ideal for tissue engineering as it

is identical to the desired matrix structure required and helps controls cell behaviour

(96-100). Thus, the use of dECM is suitable as it is biodegradable, does not produce

antagonistic immune responses, provides cues for cell differentiation, and presents

bioactive molecules that determine tissue homeostasis and tissue regeneration (101,

102). The decellularisation process comprises mechanical and chemical manipulation

to remove the cellular components (103). This protocol from harvesting,

decellularisation, and sterilisation to creating the dECM based scaffolds affects the

hydration status and 3D configuration of the proteins and ECM, and hence strongly

influences biomechanical and biological behaviour properties which may not be

suitable anymore (104, 105). The replication of the ECM microenvironment has

provided inspiration to use the ECM from articular cartilage as matrix for tissue

regeneration (101, 106). However, concerns remain about potential immunogenicity

and poor biomechanical and biological performance.

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2.1.6.2 Bioprinting cartilage tissue constructs

The advancement of 3D printing or additive manufacturing in tissue engineering has

allowed the fabrication of scaffolds and biological models that more accurately reflect

the complex organisational structure and material properties of tissues and organs

(107, 108). 3D bioprinting uses biomaterials, cells (encapsulated or seeded), and

biomolecules, typically referred to as a bioink, which are precisely deposited in a

layer-by-layer process to build-up a 3D structure. The ability to print multiple

biocompatible materials with greater design freedom compared with conventional

fabrication techniques has enabled the development of 3D structures that resemble the

complex 3D biophysical and biochemical environment in tissues. The use of 3D

bioprinting within cartilage tissue engineering is becoming widespread as an enabler

technology to fabricate complex multi-material structures that mimic the biological

and mechanical properties of cartilage tissue (109, 110). Currently, 3D bioprinting

predominately uses inkjet, extrusion, and laser-assisted systems to fabricate 3D

structures (Figure 2.7). However, stereolithography based systems are gaining

attention due to novel visible light photo crosslinkers with improved cytocompatibility

and advancements in the technology which promises faster fabrication times and

increased structure complexity (111, 112).

Figure 2.7: 3D bioprinting technologies; a) Inkjet bioprinting; b) Laser-assisted

bioprinting; c) Extrusion bioprinting (113).

The development of bioinks has become a key factor in 3D bioprinting in particular

biomaterials with controllable mechanical, biological, and biophysical characteristics

which can modulate cell behaviour combined with printability (Figure 2.8) (114, 115).

Printing resolution, structure fidelity, material viscoelasticity are crucial parameters in

determining the printability of bioinks and its relationship to the final mechanical and

biological properties of the structure. Developing advanced bioinks requires

51

consideration of pre-functionalisation processes to incorporate biological functional

groups and crosslinking moieties, the rheological behaviour of the bioink to ensure

printability and fidelity, and the crosslinking method to ensure rapid gelation of the

hydrogel (116-119). Depending on the biofabrication process and material properties,

the bioink polymers will have various chemical and physical characteristics that will

determine the corresponding application (120). These properties can be determined by

rheological characterization, mechanical assessment and crosslinking properties of the

hydrogel (120, 121). Shear-thinning behaviour, a decrease in the viscosity as a

function of increasing shear rate, is crucial for bioprinting applications, since the

material will flow with an applied force during printing and the lower the applied force

the higher the cell viability (119, 122, 123). The viscoelastic behaviour characterised

by the material response during printing needs to be optimised as low viscous materials

will deform and collapse during printing, unless a rapid crosslinking process can be

initiated. On the contrary high viscosity materials can be difficult to print as they can

block the printing nozzle, require high deposition force, and restrict cell attachment

and spreading which can negatively impact cell viability (124, 125).

Figure 2.8: Bioink properties for successful 3D bioprinting require a suitable a)

biofabrication window which balances printability and biocompatibility whilst

providing a variety of b) suitable characteristics (126).

2.1.6.3 Biochemical, biophysical and biomechanical stimulation

Biochemical, biophysical, and biomechanical stimulation of 3D bioengineered

constructs is crucial in the formation of functional neocartilage tissue. The stimulation

of cells can be achieved at specific stages: cell expansion, differentiation, and the

maturation of the tissue construct in vitro. This can be attained through

supplementation of the cell culture media, incorporation of biomolecules within the

52

3D structure, engineering the physical ECM environment, and mechanical stimulation

of the construct.

A key approach to direct cell behavior and facilitation of neocartilage tissue formation

is the use of biological signaling molecules (e.g. growth and transcription factors)

during cell culture, tissue maturation, and utilization via direct inclusion,

encapsulation or binding to the biomaterial matrix of the construct (127-132). A range

of growth and transcription factors have been identified and investigated including

transforming growth factors (TGFs), bone morphogenetic proteins (BMPs), insulin

growth factors (IGFs), fibroblastic growth factors (FGFs), platelet-derived growth

factors (PDGFs), and sex determining region Y (SRY)-box (SOXs).

Growth factors are proteins that have a key role in cell behavior and regulate cellular

growth, proliferation, differentiation, and migration and are grouped into families with

shared amino acid sequences and superfamilies with shared structural folds (130, 133).

In articular cartilage tissue engineering, 3D engineered constructs have been used to

deliver these biological factors (130). For instance, TGF-β1 stimulates the synthesis

of cartilage ECM, maintain chondrocytes phenotypes, synthesis of proteoglycans,

aggrecan and type II collagen and can enhance the repair of cartilage defects (134,

135). IGF-1 has been reported showing high anabolic effects and decrease in catabolic

responses in articular cartilage metabolism in vitro (135, 136). Other studies have

reported using different growth factors which were successful in producing several

features that resembled typical articular cartilage (137-139). Despite of the capacity to

promote cartilage matrix formation, growth factors have shown some drawbacks with

IGF-1 associated with a loss of chondrocyte phenotype and extracellular matrix

breakdown (136). Furthermore, IGF-1 in human MSCs inhibited collagen II

expression and overexpression can induce hypertrophic differentiation and

mineralization (140).

Biomechanical stimulation is key in the development and homeostasis of functional

cartilage tissue (141-145). The importance of mechanical loading and physical

movement on embryonic chondrogenesis has demonstrated in chicken embryos which

when physically impaired exhibited poor development of cartilage tissue (146-148).

Mechanical loading is required for healthy tissue; however, excessive loading can lead

to trauma and disease progression (141). Thus, mechanical stimulation of cells and

tissue constructs via compression, shear and hydrostatic pressure is important to

promote chondrogenic differentiation, maintain a chondrogenic phenotype, and

53

generation of functional tissue in vivo. Mechanical stimulation of MSCs under varied

loading regimes have been shown to increase the deposition and expression of

collagen II, GAG, TGF-1/3, and SOX9 and modulated secretory factors such as

stromal-derived factor-1 (SDF-1), matrix metalloproteinase-2 (MMP-2), FGF,

vascular endothelial growth factor (VEGF), activated leukocyte cell adhesion

molecule (ALCAM), nitric oxide, urokinase receptor (uPAR), macrophage

inflammatory protein 3 (MIP3). The advancement of bioreactors in tissue

engineering has enabled the use of mechanical stimulation as a key capability in

engineering cartilage tissue formation (149-151). Dual compressive and shear

mechanical stimulation of human articular chondrocytes encapsulated in gelatin

methacrylate and hyaluronic acid methacrylate hydrogel have been demonstrated by

Meinert et al (152). Cartilage specific marker genes and ECM were upregulated with

significant increase in collagen II synthesis. Combining compressive and shear

stimulation has been investigated using multi-axial loading bioreactor which mimics

the movement of articulating joint (153, 154). Vainieri et al investigated a chondrocyte

seeded hybrid fibrin-polyurethane scaffold implanted in an osteochondral defect

model that was mechanically stimulated using a joint mimicking bioreactor (154). The

results showed increased production of chondrogenic specific markers, proteoglycan

4 and cartilage oligomeric matrix protein, and the improved collagen II to I ratio.

Alternatively, tensile stimulation of a self-assembled scaffold-free neocartilage

construct has shown to increase the tensile strength and modulus of the construct and

once implanted in vivo in a mice model had similar mechanical properties and collagen

content of native tissue (155). The development of improved mechanically stimulating

bioreactors and osteochondral models will provide a valuable tool in understanding

cartilage development and will aid the screening of biomaterials and tissue

engineering.

Challenges remain in the use of stimulatory processes to guide cell behavior and tissue

development. A complete understanding of chondrogenic development is still

developing so the entire milieu of factors that influence tissue formation incomplete.

However, tissue engineering strategies will most likely, to be successful, use a

combination of growth factors, a controlled biophysical environment, and a

mechanical stimulation to promote tissue formation with phenotypic stability. As the

use of soluble factors alone in the differentiation of chondrocytes and MSCs in vitro

54

typically results in the expression of hypertrophic chondrocyte phenotype. The design

of these stimulatory environments will aim to recapitulate the native environment

during early stage development of the tissue. A successful strategy will need to

determine the combination, dosage, and delivery profile of growth factors. The design

of the physical matrix surrounding the cells by controlling parameters such as

crosslinking density, ECM protein selection, and oxygen tension. Finally, the timing,

type, and loading conditions of mechanical stimulation will be key in promoting an

ECM which is mechanically compliant. However, the complexity of this environment

and the actual implementation of a multi-stimulatory strategy is a serious challenge

for researchers.

Conclusion and future perspectives

This Section discusses articular cartilage structure, composition, and current clinical

therapies. Tissue engineering strategies are discussed with a focus on selection of

biomaterials, stimulatory factors, and bioprinting to provide an overview of

approaches to generate osteochondral tissue.

The clinical size and market of cartilage and osteochondral problems is expanding

due to the ageing worldwide population and the most common treatment approaches

are ineffective at stopping the progression of degeneration of the tissue. Thus, tissue

engineering strategies are key in solving this pressing clinical problem. The design

specification of any biomaterial-based 3D construct must fulfil a stringent criterion

requiring suitable biomaterial selection, scaffold architecture, fabrication technique,

stimulatory factors, and maturation of the tissue.

Key research challenges are the maintenance of phenotype in the engineered tissue

construct and the prevention of hypertrophic or fibrocartilage phenotypes being

expressed. The tissue constructs also require strategies to promote integration with

the surrounding healthy tissue in an osteochondral defect, although in total

replacement the strategy would primarily to be to anchor the neocartilage to

underlying bone. Furthermore, the underlying biological behaviour of the tissue, the

early-stage developmental biology, and haemostatic processes in adult tissue need

further understanding. This knowledge may unlock key aspects of the tissue which

may guide tissue engineering strategies. This could require a strategy of multiple

stimulatory factors (e.g. growth factors and mechanical stimulation) over an

extended maturation time of up to many years to mimic the underlying biological

55

development of the tissue and even then, the incorporation of permanent mechanical

structures may be necessary. Subsequently, future studies should focus on a multi-

stimulatory environment, long-term studies to determine phenotypic alterations and

tissue formation, and the development of novel bioreactor systems that can more

accurately resemble the in vivo environment. Finally, a clear and considered route

in the development process of the materials, structures, and strategy should be

evaluated prior and during the research phase to expedite clinical and regulatory

approval. This will allow faster and more successful access to animal trials and

eventually human clinical trials with the prospect of a successful therapy being

developed.

56

Engineering natural-based photocrosslinkable hydrogels for

cartilage applications †

Articular tissue is an avascular tissue at the ends of articulating joints which provides

lubrication and transmission of compressive forces during movement. The tissue has

poor regenerative capacity and low cellular metabolic activity thus disease or trauma

can result in significant clinical issues characterised by pain in the joints and restriction

of movement. This can affect patient’s quality of life and impose substantial

socioeconomic costs on society. Currently no treatment can prevent the long-term

degradation of the tissue, so new and improved therapies are required. Tissue

engineering based strategies are a promising approach to repair and regenerate

damaged tissue. Natural hydrogel-based scaffolds have been widely developed to

promote cartilage regeneration due to their biocompatibility, resemblance to the native

extracellular matrix, and capacity for cell encapsulation. However, a key challenge is

the engineering of biomimetic hydrogels to promote the development of articular

cartilage rather than fibrocartilage. Hydrogels can be designed through the selection

of appropriate materials, crosslinking mechanisms, mechanical properties and the

incorporation of biofunctional moieties that can regulate biodegradation, cell

attachment and differentiation, and the delivery of biomolecules. The range of

parameters to engineer provides the opportunity to fabricate biomimetic structures that

can facilitate cartilage regeneration. However, the complexity of the challenge is

daunting and requires an interdisciplinary approach to successfully fabricate hydrogel-

based scaffolds that can promote long-term regeneration of articular cartilage.

Introduction

The desire to regenerate and replace damaged or dysfunctional human tissues to

improve quality of life is as old as human history. This aspiration has been expressed

in numerous cultures throughout history such as the myth of Prometheus with his

eternally regenerating liver and Mary Shelley’s ‘Frankenstein’ in which the scientist

created new life from the rejuvenation of dead tissue. This human fascination to

regenerate the body has in recent decades developed into the rapidly expanding

scientific field of tissue engineering (TE). TE holds the promise to offer a paradigm

† This section is based on the book Chapter: Hussein Mishbak, Cian Vyas, Glen Cooper, Chris Peach,

Rúben F. Pereira, Paulo Bartolo, “Engineering natural-based photocrosslinkable hydrogels for cartilage

applications’’, in Bio-materials and Prototyping Applications in Medicine, Edited by B. Bidanda and

P.J. Bartolo, Springer, 2020 (ISBN: 978-3-030-35875-4).

57

shift in clinical treatment and understanding of a host of disease states. This shift will

enable a science fiction like prospect of infinite replacement human body parts which

will have tremendous positive impact on human health. The concept of TE has

developed into the currently accepted definition put forward by Langer and Vacanti in

1993 that states:

“Tissue engineering is an interdisciplinary field that applies the

principles of engineering and the life sciences toward the development

of biological substitutes that restore, maintain, or improve tissue

function.” (156).

Due to an ageing population worldwide and increasing numbers of age-related and

lifestyle linked diseases such as osteoarthritis, diabetes, and cardiovascular disease,

the importance of TE is enormous. TE will enable the development of superior clinical

treatments to the benefit of patients but also relieve the strain on healthcare systems.

Furthermore, TE has applications beyond tissue regeneration and replacement such as

the development of three-dimensional (3D) culture and tissue constructs that allow

pharmaceutical testing, disease modelling, toxicology testing, and improved

understanding of cell biology (157).

There are multiple strategies employed in TE to enable tissue regeneration and

replacement. A scaffold based or top-down approach, typically utilises a scaffold

which is a supporting construct that outlines the spatial requirement of the tissue,

enables cell attachment, and provides environmental cues to promote a desired cellular

response (158-160). This approach can be typically split into either acellular or cellular

techniques. Cellular approaches using a traditional in vitro TE strategy require cells to

be isolated, expanded in culture, and then seeded onto a scaffold in vitro to enable

maturation and remodelling of the construct before subsequent implantation into a

patient. An alternative method is to seed the scaffold with cells isolated at the point of

surgical implantation which avoids the in vitro maturation phase and requires a single

surgical intervention. The acellular in vivo TE scaffold approach bypasses the need for

cell seeding and maturation completely as a bio-functional scaffold is directly

implanted into the patient. This approach relies on the recruitment of cells in vivo to

initiate tissue repair and remodelling of the scaffold. An alternative approach is the

scaffold-free strategy which is a bottom-up process utilising cell spheroids, cell sheets,

and self-assembly to generate new tissue and typically does not require exogenous

scaffolds or cell seeding (161). These approaches utilise the ability of cells to fuse into

58

larger constructs and in self-assembly necessitates the recapitulation of embryonic and

developmental pathways of tissue and organ genesis (162-164). The scaffold-free

approach using cells, their own extracellular matrix (ECM) and signalling molecules

aims to generate new tissues and organs with high level of biomimicry, thus requiring

a deep understanding of the underlying biological and physical mechanisms of these

pathways. Finally, a new hybrid approach aims to combine both scaffold and scaffold-

free TE in a synergistic approach to benefit from the inherent advantages of both

techniques (165).

Hydrogels are a specific class of materials highly relevant for both the fabrication of

acellular scaffolds and cell-laden constructs. They are biocompatible and can be

designed to be biodegradable with a structure that closely mimics the native in vivo

ECM (166-168). This is a result of hydrogels typically being composed of a three-

dimensional (3D) polymer network which swells on contact with water (169). This

enables storage of large quantities of water which resembles soft biological tissues and

allows for encapsulation of cells and bioactive molecules. Furthermore, hydrogels are

typically biocompatible and allow for diffusion of oxygen, nutrients, and water-

soluble compounds. Hydrogel properties can be tuned through chemical, physical, and

biological modification which allows a variety of cell behaviours such as adhesion,

proliferation, and differentiation to be optimised. Moreover, hydrogels can be tuned

to respond to stimuli such as temperature, light, pH, and biomolecules which enables

hydrogels to have a variety of properties such as self-healing, controlled degradation,

and drug delivery (170). A pioneering example of a cell-responsive hydrogel is

Hubbell et al. demonstration that inclusion of specific peptides in a hydrogel could

enhance matrix metalloproteinases (MMP) activity and thus control the rate of

scaffold degradation (171). The ability to fabricate hydrogels with specific properties

has enabled precise control of cell behaviour which is a benefit in creating scaffolds

that will allow tissue engineering to be successfully translated.

Photocrosslinkable or photocurable hydrogels are particularly relevant for TE

applications due to the rapid gelation, control over spatiotemporal formation, and the

ability to tune polymer network properties such as crosslinking density and matrix

stiffness (172). This is often achieved by controlling operating parameters such as light

intensity, exposure time, and illumination area. When used in combination with

bioprinting techniques, photopolymerization enables the fast and automated

crosslinking of complex scaffold architectures during and post-printing. However,

59

photocrosslinking conditions need to be optimized to overcome concerns regarding

cell exposure to UV radiation and potential cytotoxicity from photoinitiators, by-

products, and unreacted reagents.

This Section outline the recent advances in the design of naturally derived

photocrosslinkable hydrogels for cartilage tissue engineering. A brief description of

the structure and composition of cartilage tissue will be firstly provided, followed by

an overview of hydrogel properties, synthesis routes, and strategies for hydrogel

biofunctionalisation within cartilage tissue engineering (Figure 2.9).

Figure 2.9: Schematic diagram of photocrosslinkable 3D hydrogel constructs and

approaches for use in cartilage tissue engineering.

Articular cartilage

Articular cartilage is a load-bearing and non-vascularised tissue that covers the ends

of bone joints (Figure 2.10a), acting as a low-friction bearing surface and a mechanical

damper for the bones (173, 174). It is a matrix-rich tissue with specialized cells

(chondrocytes) that maintain the structural and functional integrity of the ECM (175-

178). However, cartilage lacks inherent self-repair ability and remains a significant

clinical and economic challenge throughout the world, hence the zonal structure and

function are often irreversibly lost following trauma and disease. Thus, cartilage

defects are prone to develop into osteoarthritis (OA), which is a major disease

60

associated with cartilage tissue and is characterized by a loss of healthy and functional

cartilage tissue, pain, and debilitation thus is a significant clinical and economic

challenge (179, 180).

2.2.2.1 Composition and zonal organization

Cartilage is characterized by a distinct zonal structure comprising, in descending

order, the articular or superficial surface, the middle (or transitional) zone, the deep

zone, a tidemark (separates the non-calcified and calcified regions) (18, 181), and the

calcified zone (Figure 2.10b). These zones differ with respect to the molecular

composition and organization of the cartilage ECM, the shape, density, alignment of

the resident chondrocytes and the mechanical properties of each zone. The

subchondral bone acts as a shock absorber, the loading forces experienced by the joint

are transmitted to this region, and the bone tissue contains blood vessels that supply

nutrients, oxygen, and removes waste from the lower regions of the avascular cartilage

tissue (182). Surrounding each individual chondrocyte are three distinct ECM regions

which depending on the distance from the cells are termed the pericellular, territorial,

and interterritorial region. The highly organised and hierarchical properties of the

tissue are responsible for its remarkable mechanical properties (183-187).

Figure 2.10: a) Articular cartilage covers the surfaces of bones in diarthrodial joints

(e.g. knee and elbow); b) Zonal structure and composition of articular cartilage (40)

Articular cartilage is typically composed of:

Chondrocytes: responsible for production and maintenance of the cartilage ECM and

essential for tissue formation and functionality. They represent about 5-10% of the

total volume of cartilage (17, 187). However, they have limited proliferation capability

and low metabolic activity, which is partially responsible for the limited capacity of

cartilage to recover from trauma or disease (188, 189).

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Collagen: the main component of cartilage with up to 60% of the dry weight being

collagen with type II the most common type representing 90-95% of total collagen in

the ECM which can form fibrils and fibres entwined with proteoglycan aggregates

(65).

Proteoglycans: molecules responsible for the resistance to compression of cartilage

tissue (190). Proteoglycans (PG), corresponding to around 35% of the cartilage dry

weight, are made up of repeating disaccharide units, glycosaminoglycans (GAGs),

with two main types; chondroitin sulphate and keratin sulphate (190). Proteoglycans

have a turnover rate of three months and weave between the collagen fibres creating a

mesh responsible for retaining water and providing elasticity to the tissue (191).

Water: represents 65-80% of the total cartilage mass (192). This content reduces with

aging and is linked to a reduction of strength and elasticity of the tissue (193).

Synovial fluid: a highly viscous liquid secreted by synovial lining cells, they are

responsible for secreting hyaluronic acid (the intrinsic component of synovial fluid)

(194, 195). The main function of the synovial fluid is to distribute nutrients and gases

to the upper regions of the cartilage tissue whilst providing lubrication at the interface

between articulating surfaces in the joint to allow facilitation of movement (196).

2.2.2.2 Clinical treatments and challenges

Cartilage disease and traumas are a serious clinical problem that can lead to

debilitating pain, swelling, inflammation, and can result in drastically impaired

mobility. An ageing worldwide population is resulting in increased number of patients

suffering from osteoarthritis, a leading cause of cartilage damage (197). Additionally,

tissue degeneration through disease or significant trauma results in steady degradation

of the tissue with limited successful clinical interventions to prevent this long-term

decline which eventually results in a total joint replacement (Figure 2.11). The limited

regenerative capacity of cartilage is due to the avascular nature of the tissue and low

metabolic activity of the resident chondrocytes and is in marked contrast to other

tissues such as bone and the liver which have remarkable healing properties in

comparison (163, 198). Current clinical interventions have limited success in

regenerating or replacing articular cartilage with most therapies producing inferior

fibrocartilage or only delaying the progression of tissue degeneration (199, 200).

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Figure 2.11: Stages of articular cartilage damage. 0) Healthy cartilage. I) Softening,

swelling and fissuring of the cartilage. II) Lesions <50% of cartilage depth. III)

Lesions >50% of cartilage depth but not through subchondral bone layer. IV) Severely

abnormal with penetration through the subchondral plate. Reproduced with permission

from the International Cartilage Regeneration & Joint Preservation Society (ICRS)

(200).

The most commonly used cartilage repair strategies are debridement and

microfracture, autologous chondrocyte implantation (ACI), matrix-assisted

autologous chondrocyte transplantation/implantation (MACT/MACI), autologous

matrix-induced chondrogenesis (AMIC), and osteochondral auto- and allografts

(Figure 2.12) (201, 202). Each of these approaches has major disadvantages such as

the formation of inferior fibrocartilage, requirement of two surgeries, donor site

morbidity, and limited prevention of continued tissue degeneration. Therefore, new

approaches are required to address these unmet clinical challenges. TE has the

potential for a breakthrough approach by synergistically incorporating

chondrocytes/stem cells, biomaterials, 3D scaffolds, stimulatory growth factors,

mechanical stimulation, and bioreactors for the fabrication of tissue constructs. If this

approach is successful, then we will approach our final goal of the predictable

regeneration of articular cartilage (110, 203-207).

63

Figure 2.12: Current techniques for articular cartilage repair and future scaffold-based

approaches. a) Debridement and microfracture; 1-4 debridement of damaged tissue, 5

microfracture into the subchondral bone, 6 blood clot formation containing

mesenchymal stem cells (MSCs), and 7 formation of inferior fibrocartilage. b) ACI; 1

cartilage biopsy from a non-load bearing site, 2 in vitro expansion of chondrocytes to

obtain sufficient numbers for implantation, 3 a second operation cleans the damaged

area and is covered with the periosteal flap and the expanded chondrocytes are injected

under this or the expanded chondrocytes are seeded into a matrix (e.g. collagen

membrane) and fixed with fibrin glue into the defect area (MACI/MACT). c) AMIC;

a microfracture is created and a membrane placed over the top and fixed with a fibrin

glue or sutures. The membrane protects the clot formation and allows chondrogenic

differentiation of the MSCs in the blood to aid the regeneration process. d)

Osteochondral auto/allograft transplantation; cylindrical plug of fresh cartilage and

subchondral bone is taken from the patient or a cadaver donor, 1 a guide pin is placed

perpendicular to the joint surface in the defect, 2 implantation of the graft plug, and 3

securing the plugs with the insertion of screws. e) Acellular scaffold only requiring a

single surgery and in vivo recruitment and stimulation of cells for tissue regeneration.

f) Cellular-based scaffold involving two surgeries of initial cartilage biopsy,

chondrocyte expansion, and seeding into the scaffold before implantation into the

defect. Adapted with permission from the ICRS (200)

64

Engineering hydrogels for cartilage tissue engineering

Hydrogels are three-dimensional (3D) insoluble crosslinked polymer networks able to

retain a large amount of water and biological fluids (between 10-200%) in their

swollen state (208-211). The 3D network of the hydrogel (Figure 2.13) is maintained

by polymer crosslinking among the polymer chains, with water taken up and contained

within the polymeric structure (212, 213).

Figure 2.13: Engineering 3D hydrogels for cartilage application through a) cell-

encapsulation, b) optimizing crosslinking density, and c) functionalizing with

biomolecules (e.g. RGD) to guide cell behavior.

Hydrogels have a range of useful characteristics such as their hydrophilic capacity to

absorb a volume of water that is significantly beyond their weight when dehydrated

(214), potential printability and ability to create a self-supporting 3D structure to suit

various biological engineering needs (cell encapsulation, cell support and

biodegradability) (215), uniform cell distribution, as well as providing both chemical

and biological signalling (216). The ability of the 3D network to resist dissolution in

wate r is related to the presence of the crosslinked polymer molecules which will

determine the quality and stability of the hydrogel. Due to these properties and

characteristics, a wide range of different hydrogels have been used for ECM formation

and cartilage regeneration applications. Currently, natural based hydrogels represent

the most relevant group of materials to produce constructs for cartilage applications

(Table 2.1). Depending on the source, they can mimic the ECM structure and

composition of articular cartilage presenting relevant swelling, degradation,

mechanical and lubrication characteristics. Natural based hydrogel materials such as

alginate (217), agarose (218), gelatin (219), collagen (220-222), hyaluronic acid (HA)

(223), chitosan (224), and interpenetrating network/hybrid systems have been

extensively explored for cartilage applications (225).

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Table 2.1: Photocrosslinkable natural polymers used in cartilage tissue engineering.

Polymer Crosslinking conditions Biological properties Refs.

Alginate

Methacrylate anhydride, photoinitiators (VA-

086, Irgacure 2959) and 254-365 nm exposure wavelength. Simple preparation.

Chondrocytes showed ECM formation with ~80%

similarity to native cartilage. >85% viability

Stem cells/ MSCs/ hASCs: could be regulated via the

introduction of soluble factors and biophysical cues in 3D cell culture systems

(226-228)

Agarose

Typically, physical crosslinked (thermal),

however, can be functionalized with

methacrylate group,

Chondrocytes: no results reported (for chondrocyte

cells encapsulated in photocurable agarose).

hBMSCs: cell viability in agarose scaffolds decreased

from 75% at day 3 to 40% at day 90.

Fibrochondrocytes: 80% cell viability and maintains

chondrocyte phenotype

(229-231)

Gelatin

Methacrylate and acrylic anhydride,

photoinitiators (VA-086, Irgacure 2959) and 254-365 nm exposure wavelength. Simple

preparation.

Chondrocytes showed ECM formation similar to

native cartilage . >85-97% viability, cell morphology could be controlled by the stiffness of hydrogels.

Stem cells/ MSCs/ hASCs : could be regulated via the introduction of soluble factors and biophysical cues

in 3D cell culture systems

(232-235)

Hyaluronic

acid

Glucuronic acid carboxylic acid, the primary

and secondary hydroxyl groups, and the N -

acetyl group (following deamidation) thiols, methacrylate, tyramines

Photoinitiators (VA-086, Irgacure 2959) and

254-365 nm exposure wavelength

Chondrocytes, epithelial cells and murine fibroblasts

cells

For cartilage engineered tissue, ECM formation reached about 80% of those found in native cartilage.

>85% viability

Promoted the retention of chondrocyte phenotype and

matrix synthesis of encapsulated porcine

chondrocytes and enhanced cell infiltration and tissue repair in rabbit osteochondral defect model at 2-week

post-surgery

Stem cells/ MSCs/ hASCs: could be regulated via the

introduction of soluble factors and biophysical cues in 3D cell culture systems

(233)

Chitosan

(CS)

Vinylated CS Macromers, Azido-Functionalized CS azido, and (methyl)acrylated CS

Structurally analogous to cartilage glycosaminoglycans, Biocompatible, showed low

cytotoxicity at various concentrations for

chondrocytes, cell viability between 82-96%

(236, 237)

Dextran

Methacrylate groups (glycidyl methacrylate) introduced in dextran to prepare

photocrosslinkable dextran (photoinitiator

Irgacure 2959 to UV light. 254-365 nm

wavelength

Chondrocytes, endothelial cells and stem cells, the dextran-based IPN hydrogel provides cell-adhesive

and enzymatically degradable properties

Fibroblasts ~90% cell viability.

(238, 239)

Gellan Gum

Methacrylic anhydride, the aqueous solution of methacrylated gellan is added with 0.08 mg/mL

calcium chloride and 0.5% (w/v) Irgacure 2959

Human nasal chondrocytes: The gellan gum supports is their use as cells encapsulating agents to be used in

cartilage regeneration approaches

Exhibits good cell viability and chondrocyte morphology.

(240, 241)

Collagen

Collagen I: dual crosslinking with visible light,

Rose Bengal, and chemical crosslinking using 1-

ethyl-3-(3-dimethylaminopropyl) carbodiimide and N-hydroxysuccinimide.

Chondrocytes had ≥ 95% viability. Dual-crosslinked

gels showed greater resistance to collagenase

digestion.

(222)

Collagen II: amidation reaction between amino groups on the collagen lysine and methacrylic

anhydride resulting in collagen methacrylamide.

Irgracure 2959 used as photoinitiator and exposed to UV (8 W cm−2) for 30 s. Triple

helical structure of collagen was preserved after

functionalization.

BMSCs differentiated into a chondrogenic phenotype with lacuna-like cartilage formation. High cell

viability. Increase in chondrogenic gene expression

(AGG, SOX9, COL II, COL X). In vivo subcutaneous implantation in a mouse model

demonstrated the secretion and deposition of cartilage

specific matrix.

(220)

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Although pure natural hydrogels have been widely explored in cartilage applications,

they often have significant limitations relating to the suitability of their biomechanics,

degradation profiles, and cell-material biological interactions. Subsequently,

significant research into new hydrogel systems that have tuneable mechanical and

degradation properties, provide specific control over cellular behaviour, allow

incorporation of novel materials, and display suitable material processing properties

for different fabrication routes such as 3D printing have been developed (126, 242-

245). The incorporation of multiple materials, hybrid hydrogels, to produce a system

with a mixture of two or more of distinct classes of molecules (e.g. a blend of natural

and synthetic hydrogels) or different structures (e.g. blend of organic polymer

hydrogels and inorganic particles/metals or synthetic nanoparticles) have been

developed to improve single material-based systems (Figure 2.14) (246-249).

Figure 2.14: Hybrid hydrogel network systems (96). Multi-material bioink networks

are two different component polymers or more crosslinked together. Interpenetrating

networks (IPN) is a 3D structure of two or more polymeric networks, which are

partially or fully chemically bonded but not covalently bonded, or a network of

embedded linear polymers entrapped within the original hydrogel (semi-IPN) (97, 98).

Nanocomposites bioink network can be produced by adding nanoparticles to the

polymeric hydrogels (99). Supramolecular bioink network is composed of short

repeating units with functional groups that can interact non-covalently with other

functional units, forming large, polymer-like entanglements (100).

67

The use of photocrosslinkable hybrid hydrogels is a burgeoning research field

especially within tissue engineering. Photocurable hydrogels are particularly relevant

for biomedical applications due to the advantage of being able to encapsulate cells,

fast gelation time, tuneable mechanical properties and crosslinking density, and the

cytocompatibility processing conditions that allow their in situ crosslinking into the

defect site during a surgical procedure (250-252). The use of vat photopolymerization

and light curing processes in additive manufacturing also enables the fabrication of

high resolution and complex structures that are not possible through conventional

methods (111, 172). This approach facilities the development of more biomimetic

structures that can recapitulate the complexities of native tissues.

Typically, polymers are unable to initiate a photopolymerization reaction alone and

can require both functionalisation with photocrosslinkable groups and the addition of

light sensitive molecules, called photoinitiators. The photoinitiators typically used

within tissue engineering are normally excitable within the ultraviolet (UV) (e.g. 1-4-

(2-hydroxyethoxy)-phenyl-2-hydroxy-2-methyl-1-propane-1-one, Irgacure 2959, and

2,2'-azobis-2-methyl-N-(2-hydroxyethyl)propionamide, VA-086), UV-visible (e.g.

lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), and diphenyl(2,4,6-

trimethylbenzoyl) phosphine oxide, TPO), and visible light range (e.g. ruthenium,

Bengal rose, riboflavin, eosin-Y, and camphorquinone). Upon irradiation, an initiation

step begins and the photoinitiator undergoes an excitation mechanism which results in

the formation of highly reactive species (free radicals). The next step is propagation,

free radicals attach at specific sites on the monomer/oligomer structure (unsaturation

sites) starting a chain reaction that creates a permanent crosslinked network. Finally,

a termination step results in the completion of the polymer chain and crosslinking with

free radical extinction. The degree of photopolymerization (curing percentage) of the

resulted hydrogels is highly affected by the functionalisation process and polymer

solution concentration, light wavelength and intensity, photoinitiator type and

concentration, and curing time (253). The number of parameters to control also

provides an opportunity to highly engineer the resulting hydrogels properties and

when used in conjunction with bioprinting processes can result in tuneable and

biomimetic structures.

Photocrosslinking is an advantageous process, however, specific issues need to be

addressed especially in biomedical applications. The chemical processes used can

generate toxic by-products that can result in cytocompatibility residues in the hydrogel

68

(254). Thus, the hydrogel is usually subjected to a purification step prior to its

application, for example, the modified macromer may require dialysis prior to use.

Furthermore, the photoinitiators themselves can be toxic, Thus, specific care is

required when designing the photocuring system such as using a low concentration of

photoinitiator. Finally, the wavelength and intensity of light can induce genetic

damage within the cells resulting in reduced viability and function, thus this requires

careful consideration (255-259).

Crosslinking mechanism, biofunctionalization, controlled delivery of biochemical

factors and hydrogel physical characteristics are the most crucial parameters that

determine hydrogel suitability for the target application.

Crosslinking mechanisms of hydrogels

The crosslinking mechanisms of hydrogels are broadly classified into two categories

(Figure 2.15): chemically crosslinking (permanent crosslinking) and physical

crosslinking (reversible crosslinking) (260). Chemically crosslinking hydrogels are

formed by covalent bonding that creates a permanent 3D polymer network (261).

Physical crosslinking hydrogels are produced through the entanglement mechanism of

the polymer chains (262). For the fabrication of both scaffolds and cell-laden

constructs, chemical crosslinking is the most suitable mechanism (254, 263, 264).

Figure 2.15: Crosslinking mechanisms using in physical and chemical gelation of

hydrogels (265).

Crosslinking reactions can be induced by using a variety of stimuli such as light, pH

stimulation, ionic exchange, enzymes, and temperature (254). The crosslinking

density strongly determines the characteristics of the hydrogels, such as mechanical,

degradation, porosity, pore size, swelling and biological properties (266, 267).

69

Light initiated crosslinking reactions correspond to the most common approach,

allowing the fabrication of complex structures under relatively mild conditions. Two

different reaction mechanisms are being considered: free radical polymerisation and

click chemistry.

Free radical photopolymerization is a chain growth reaction started by the absorption

of light by photoinitiators. The photoinitiator concentration determines the

polymerisation kinetics and consequently the properties of the polymerised structures.

Radical photopolymerisable hydrogels can be functionalised with cell adhesive

moieties and degradation sites in a relatively easy and reproducible manner (268).

Major limitations are associated to the relatively poor control over the crosslinking

kinetics, the presence of unreacted double bonds which can react with biological

substances, and the generation of heterogeneities due to the random chain reaction

process. As an alternative, orthogonal click reactions (e.g. thiol-norbornene) proceed

under mild reaction conditions with higher efficiency and faster curing kinetics

compared to free radical photopolymerization (269). This step-growth reaction allows

high spatio-temporal control and the fabrication of 3D networks with minimal defects

(270).

Biofunctionalization of hydrogels

Hydrogels for cartilage tissue engineering must be designed to closely capture the

composition, architecture, biophysical properties and biochemical cues of the native

tissue. This is essential to provide a cell-instructive 3D microenvironment that guides

the fate of embedded cells and morphogenesis. In the case of acellular networks,

hydrogel characteristics such as mechanical properties and porosity must ensure

adequate load transfer and eventually promote cell recruitment. Multiple approaches

have been explored to develop photocrosslinkable hydrogels that perform fundamental

functions of the ECM, including the (i) material selection and hydrogel structure, (ii)

modulation of hydrogel viscoelasticity, (iii) bioconjugation of bioactive peptides, and

(iv) tethering of biochemical factors (271).

2.2.5.1 Hydrogel composition and structure

Important design criteria for hydrogels include appropriate mechanical properties,

ensuring nutrient/waste transport and maintenance of cell viability. Several studies

have modulated the mechanical properties of hydrogels by varying the macromer

concentration or crosslinker type/content using a single polymer network. These

70

works demonstrated that hydrogel stiffness has a tremendous impact on matrix

deposition with highly crosslinked networks restricting new tissue formation and

eventually leading to matrix calcification (272, 273). In addition to the matrix stiffness,

it was also reported that the molecular weight (50 to 1100 kDa) and macromer

concentration (2 to 20 % wt ) of methacrylated hyaluronic acid hydrogels determine

the amount and distribution of new ECM via modulation of the gel structure (e.g.,

compressive modulus, swelling ratio) (274). Another important consideration of

hydrogels is to balance the mechanical integrity and the degradation rate. Gel networks

should support tissue growth as the hydrogel degrades, maintaining the structural

integrity of the new ECM. One interesting approach to control the degradation rate of

hydrogels for cartilage applications involves the synthesis of photocrosslinkable

macromers bearing hydrolytically labile linkages. This strategy was explored to design

hydrolytically degradable hydrogels via thiol-ene photopolymerization of norbornene-

modified Poly(ethylene glycol) (PEG) containing caprolactone segments with

hydrolytically labile ester linkages and PEG dithiol crosslinker in the presence of a

photoinitiator (275). Although hydrogels based on a single polymer are relatively

simple to design, their limited mechanical properties and potential reduced cell

viability due to high crosslink density constitute important concerns. To tackle these

limitations, more complex hydrogels have been developed based on the combination

of multiple polymers. Depending on the selected materials and crosslinking

chemistries, different hydrogel structures have been designed, including

interpenetrating networks (IPNs) (276), semi-IPNs (277), double networks (278) and

guest-host networks (279). Wei et al (280) synthetized a two-component guest-host

hydrogel of adamantane-functionalized hyaluronan (HA) as guest polymers and

monoacrylated β-cyclodextrin as host monomers, rapidly crosslinked through UV

photopolymerization. Hydrogels displayed self-healing properties, supported robust

chondrogenesis of embedded human MSCs and stimulated cartilage formation in a rat

model. In addition to the dynamic nature of the network and/or superior mechanical

properties often achieved by more complex hydrogel systems, these gels can also

address key compositional and biological properties of the native tissue by the

incorporation of major components of the cartilage ECM (281). One strategy consists

of the chemical incorporation of ECM components bearing a photocrosslinkable

moiety into the hydrogel network via photopolymerization. This strategy was explored

by Wang et al (282) to evaluate the influence of methacrylated ECM molecules

71

(chondroitin sulfate, heparan sulfate and hyaluronic acid) on the in vivo cartilage

formation using ECM-containing hydrogels with tuneable stiffness and controllable

biochemical composition. Methacrylated ECM molecules were incorporated into the

network of PEG dimethacrylate hydrogels through photopolymerization in the

presence of adipose derived stem cells and neonatal chondrocytes. Cellular hydrogels

with varying matrix stiffness and biochemical cues showed increased collagen type II

deposition in both chondroitin sulfate- and hyaluronic acid-containing hydrogels,

while heparan sulfate-containing hydrogels promoted a fibrocartilage phenotype and

failed to retain newly deposited matrix. This study revealed the importance of

biochemical composition of hydrogels and matrix stiffness on the phenotype of

neocartilage formation.

2.2.5.2 Hydrogel viscoelasticity

Traditional hydrogels obtained by photopolymerization reactions are made of strong

covalent bonds, leading to almost purely elastic gel networks. In contrast to native

cartilage tissue, which is viscoelastic and exhibit partial stress relaxation when a

constant strain is applied, chemical hydrogels store cellular forces and resist

deformation. Purely elastic hydrogels with embedded chondrocytes were found to

limit cell proliferation and restrict cartilage matrix deposition to the pericellular space,

probably due to the elastic nature of hydrogels (283). Recent findings have shown that

viscoelastic hydrogels, in which the stress is relaxed over time, support the spreading,

proliferation and differentiation of embedded mesenchymal stem cells in 3D culture

without requiring matrix degradation (284). It has also been reported that cell-laden

hydrogels exhibiting faster stress relaxation support significantly more new bone

growth in vivo when compared to slow stress-relaxing gel networks (285). As stress

relaxation is suggested as a key design parameter of hydrogels to both better

recapitulate the viscoelastic behaviour of natural ECM and perform important

biological behaviours, efforts have been focused on the design of viscoelastic

hydrogels with tuneable stress relaxation for cartilage tissue engineering. In a recent

work, Lee et al, (286), developed a series of viscoelastic alginate hydrogels exhibiting

tuneable stress relaxation to evaluate the effects of viscoelasticity on embedded

chondrocytes in 3D culture. Stress relaxation of calcium-crosslinked hydrogels was

controlled by varying the molecular weight of alginate and covalent coupling of short

PEG spacers. It was found that hydrogels with faster stress relaxation supported the

72

formation of increased cartilage matrix with deposition of both collagen and GAGs

(Figure 2.16), while slower relaxing gel networks led to the up-regulation of genes

associated to cartilage degradation and cell death. An alternative approach to engineer

hydrogels with tuneable viscoelastic properties consists on the formation of covalent

adaptable networks with the ability to reorganize the network connectivity to dissipate

local stress. As an example, hydrazone crosslinked PEG hydrogels were synthesised

as viscoelastic gels for cartilage applications (287). After 4 weeks of culture, porcine

chondrocytes embedded within hydrogels exhibiting average relaxation times of 3

days secreted an interconnected articular cartilage-specific matrix with increased

deposition of collagen (e.g., collagen type II) and sulfated glycosaminoglycans (e.g.,

aggrecan) compared to predominantly elastic hydrogels with slow average relaxation

times (~ 1 month). The authors suggested that a balance between slow and fast relaxing

crosslinks is essential to preserve gel network integrity and support the formation of

high-quality neocartilaginous tissue.

Figure 2.16: Influence of hydrogel stress relaxation on cartilage matrix production

and formation. Immunohistochemical staining of chondrocytes embedded within 3

kPa alginate hydrogels for 21 days (scale bar: 25 µm) (170).

73

2.2.5.3 Bioconjugation of bioactive peptides

Articular cartilage is an intricate meshwork rich in collagen and proteoglycans. Rather

than a static milieu, the ECM is a highly dynamic microenvironment that establishes

reciprocal interactions with neighbouring cells. Cartilage ECM components provide

biochemical motifs for cell adhesion and degradation sites for cell-mediated matrix

degradation, which are essential to allow cell adhesion, proliferation, migration and

tissue formation. As most cell types require adhesion sites to survive and perform their

functions, several bioconjugation strategies have been explored for the

functionalization of hydrogels with bioactive peptides that impart cell adhesion

domains (288). The fibronectin-derived adhesion peptide RGD (arginine-glycine-

aspartic acid) is often bond to the hydrogel network to promote integrin-mediated cell

binding and allow cell attachment and spreading (276, 289). Other sequences derived

from collagen and decorin have also been immobilized into hydrogels with beneficial

outcomes regarding chondrogenic differentiation of MSCs and matrix retention (290,

291). Since chondrocytes in native ECM display a round morphology, the benefit of

RGD functionalization in chondrocytes is still controversial. For example,

chondrogenesis of bone marrow stromal cells embedded within RGD-functionalized

alginate hydrogels was inhibited in response to TGF-β1 and dexamethasone regarding

gene expression and matrix synthesis (292). Another work showed that

photocrosslinked RGD-modified PEG hydrogels enhanced cartilage-specific gene

expression and matrix synthesis by bovine chondrocytes, but only in the presence of

dynamic mechanical stimulation (293). In fact, most of the hydrogel systems used for

chondrocyte encapsulation did not require cell-adhesion domains to support tissue

formation (277, 286). Overall, this data demonstrates that the immobilization of cell-

adhesion domains in hydrogels perform a key role in terms of cell phenotype and

chondrogenesis, further studies are necessary to elucidate the cell-specific influence

of peptide ligands.

Natural ECM is continuously remodelled by cell-secreted enzymes, which is required

to create space in the matrix for cell migration, proliferation and tissue deposition.

Thus, the development of hydrogels susceptible to hydrolytic and/or cell-driven

degradation is becoming increasingly recognized as a key feature to improve the ECM

biomimicry (294). Hydrolytic degradation of hydrogels is commonly obtained by

hydrolysis of ester linkages, leading to changes in the overall network properties.

While enzymatic degradation depends on the cell type and levels of circulating

74

enzymes, hydrolytic degradation is spontaneous and occurs via bulk and/or surface

erosion mechanisms. Cell-driven hydrogel degradation via cleavage of protease-

sensitive peptides occurs in a more localized manner, at the pericellular space, better

recapitulating the remodelling of natural ECM. Hydrogels based on animal polymers

such as fibrin, collagen and hyaluronic acid are naturally degraded, whereas some

plant-based polymers (e.g., alginate, pectin) and synthetic polymers are often modified

to make them susceptible to enzymatic degradation (288). This has been accomplished

by the incorporation of peptide sequences recognized by specific proteases. These

peptides can be grafted into the polymer backbone or used as crosslinker agents with

a dual function – allow gel formation and promote matrix degradation. As an example,

MMP7-degradable hydrogels based on recombinant Streptococcal collagen-like 2

(Scl2) proteins and functionalized with peptides that bind to hyaluronic acid and

chondroitin sulphate were synthetized for cartilage tissue engineering (295). Cell-

degradable hydrogels with embedded hMSCs and functionalized with GAG-binding

peptides significantly enhanced chondrogenic differentiation and gene expression of

COL2A1, ACAN, and SOX9, compared to non-functionalized hydrogels.

2.2.5.4 Tethering and controlled delivery of biochemical factors

Cartilage formation is regulated by a coordinated spatiotemporal delivery of

biochemical factors that interact with cells to induce chondrogenesis. To this end, a

major role of ECM is to serve as a storage depot for growth factors, regulating their

spatiotemporal and tissue-specific presentation to the cells. Its ability to locally bind,

store, and release growth factors is essential to regulate their bioavailability,

bioactivity and stability in order to elicit a biological response (296). The binding of

growth factors to the ECM, for example, via electrostatic interactions to heparan

sulfate, is fundamental to enhance their activity in the vicinity of cells, prolong their

action and protect them from degradation (297).

To recreate the dynamic presentation of bioactive cues available to the cells in native

cartilage, hydrogels have been designed with the ability to sequester and/or deliver

biochemical cues such as growth factors. This has been achieved through different

strategies including the covalent immobilization (298), affinity binding (299), and

encapsulation within carrier vehicles (300). Members of the transforming growth

factor beta (TGF-β) superfamily such as TGF-β1 play a role in regulating

chondrogenesis during development. To allow the immobilization into the hydrogel

75

network towards improved bioactivity and cell presentation, growth factors are usually

modified with photoreactive groups. In one example, TGF-β1 was reacted with 2-

iminothiolane, yielding thiolated TGF-β1 that can homogeneously bond throughout

the gel network of non-degradable PEG hydrogels during photopolymerization

without impairing its bioactivity (298). Thiolated TGF-β1 was covalently linked to the

norbornene end groups of PEG hydrogels crosslinked by photoinitiated step-growth

polymerization using an MMP-degradable peptide (KCGPQGIWGQCK) as a

crosslinker (301). The ability of chondrocytes to promote in situ hydrogel degradation

was firstly confirmed using a fluorogenic peptide sensor to determine the extent of

MMP‐sensitive sequence cleavage (Figure 2.17). In this hydrogel, GAG and collagen

deposition was found to be restricted to the pericellular space, which was attributed to

the limited chondrocyte degradation of the gel network to allow diffuse matrix

production. When hydrogels were used to co-encapsulate chondrocytes and hMSCs, a

higher degradation rate was observed, along with increased GAG and collagen

deposition at 14 days of culture compared to non-degradable hydrogels. This data

highlights the importance of cell-degradable motifs and localized presentation of

biochemical cues in photocrosslinked hydrogels to support the formation of cartilage

specific ECM.

76

Figure 2.17: Effect of chondrocytes embedded within cell-degradable hydrogels. A)

Schematic illustration of cell-laden hydrogel formation with tethered MMP

fluorescent sensor (Dab‐ GGPQG↓IWGQK‐Fl‐AhxC), and TGF‐β1; hydrogels are

crosslinked using either an MMP‐degradable peptide sequence

(KCGPQG↓IWGQCK) or non‐degradable (3.5 kDa PEG dithiol) linker. B)

Determination of in situ cleavage of fluorescent sensor by chondrocytes. C) GAG

staining of sections from cell-laden hydrogels after 28 days of culture: nuclei stained

black and GAGs stained red. D) Collagen staining of sections obtained from cell-laden

hydrogels at day 28: nuclei stained black and collagen stained blue (scale bars: 100

μm) (185).

Conclusions and future perspectives

Naturally derived hydrogels are suitable materials for cartilage applications. They are

biocompatible and biodegradable with a chemical and physical structure that

resembles the ECM structure of natural tissues. Different light-mediated

77

photopolymerization strategies have been explored for the fabrication of scaffolds

with micro- and nano-scale resolutions. These strategies proceed under biocompatible

conditions in the presence of cells and biochemical signals allowing the fabrication of

cell-laden bioconstructs. In order to improve the biological performance of both

scaffolds and cell-laden constructs different biofunctionalisation strategies have been

explored. Many naturally derived hydrogels degrade through enzymatic and/or

hydrolytic mechanisms, adding another layer of complexity to control their

degradation rate. This is an important limitation in terms of the design of suitable

constructs for cartilage tissue engineering.

Another critical challenge in cartilage tissue engineering is the integration and

engineering of hydrogels that are specific for each region of the osteochondral tissue.

The structure and properties of constructs for osteochondral applications must address

the characteristics of the two comprising tissues, articular cartilage and subchondral

bone, which still represents a major research challenge. The section of such constructs

related to the cartilage zone must present adequate mechanical properties to resist

mechanical loading and friction whilst promoting ECM formation and chondrogenic

expression of mesenchymal stem cells or chondrocytes, inhibiting hypertrophic

differentiation and mineralization of chondrocytes in the upper regions of the tissue.

Contrary, the section of the constructs corresponding to subchondral bone must

promote the formation of a blood vessel network, stimulate osteoblast proliferation,

and osteogenic differentiation of mesenchymal stem cells whilst remaining

biomechanically suitable until successful tissue regeneration is complete. How to

engineer hydrogels to promote or inhibit specific cellular behavioural pathways is a

major challenge, for example, a osteochondral hydrogel must promote hypertrophic

chondrocytes in the calcified zone and flatter and aligned chondrocytes in the

superficial surface. This will demand the development of multi-material and multi-

scale hydrogels that are specifically biofunctionalized to promote the complex

behaviour and organisation of chondrocyte cells. Furthermore, top-down fabrication

techniques will need to improve their resolution and multi-material processing

capabilities thus driving demand for hybrid biomanufacturing systems that incorporate

multiple fabrication technologies into a single system. Whilst bottom-up fabrication

processes will require further understanding of the fundamental developmental

biology of osteochondral tissue.

78

Chapter 3

Development and

characterisation of a

photocurable alginate

bioink for 3D bioprinting†

† This Chapter is based on the paper: Hussein Mishbak, Glen Cooper, Paulo Bartolo, “Development

and characterisation of a photocurable alginate bioink for 3D bioprinting”, International Journal of

Bioprinting, 5(2): 189, 2019 http://dx.doi.org/10.18063/ijb.v5i2.189. A preliminary version of this

manuscript was also presented at the International Conference on Engineering Science and Applications

(Tokyo, Japan, 2017) and was awarded as the best paper.

http://dx.doi.org/10.18063/ijb.v5i2.189

79

Alginate is a biocompatible material suitable for biomedical applications, which can

be processed under mild conditions upon irradiation. This Chapter investigates the

preparation and the rheological behaviour of different pre-polymerised and

polymerised alginate-methacrylate systems for 3D photopolymerisation bioprinting.

The effect of the functionalization time on the mechanical, morphological, swelling

and degradation characteristics of crosslinked alginate hydrogel is also discussed.

Alginate was chemically modified with methacrylate groups and different reaction

times considered. Photocurable alginate systems were prepared by dissolving

functionalized alginate with 0.5-1.5% photoinitiator solution and crosslinked by

ultraviolet (UV) light (8 mW/cm2).

Introduction

Three different approaches are being explored for tissue engineering applications

(302, 303). The first approach, cell therapy, is based on harvesting cells, sorting,

expanding and implanted them. This is a simple process but presents limited outcomes

as it is difficult to keep the cells in the desired region for clinically relevant periods of

time (289). The second approach, scaffold-based approach, is based on the use of

three-dimensional support structures that provide the necessary environment for cell

attachment, differentiation and proliferation (304). In this approach, scaffolds can be

directly implanted after fabrication or seeded with cells and pre-cultured in a

bioreactor before implantation (305). Finally, the third approach (bioprinting) uses

bioinks (hydrogels and cells) to create cell-laden constructs. This is a highly relevant

approach allowing in situ printing (172, 306). Techniques such as inkjet bioprinting,

extrusion-based and photopolymerization-based process are being explored. Among

them, photopolymerization is a very versatile method allowing a rapid crosslinking

under biocompatible reaction conditions without the use of solvents (172, 302).

Additionally, photocurable hydrogels are particularly relevant for biomedical

applications due to the advantage of being able to encapsulate cells and the mild

processing conditions that allow their in-situ crosslinking within a patient during a

surgical procedure

Suitable hydrogels for bioprinting must be biocompatible, biodegradable, present

appropriate mechanical properties, which depend on the type of tissue, good

printability (307) and shear thinning properties to facilitate the printing process (308).

In the case of photopolymerization bioprinting systems the amount and type of

80

photoinitiator is also critical as determines the crosslinking density, cytotoxicity,

mechanical properties and biocompatibility (309). The increase of crosslinking density

is usually associated to an increase of printability and mechanical properties and a

decrease of biocompatibility due to the reduction of free space to accommodate cell

proliferation (310).

Alginate is a suitable material for bioprinting (311). It is a natural water-soluble linear

polysaccharide derived from alginic acid, extracted from several species of brown

algae, such as Laminaria Hyperborea, Ascophyllum Nodosum and Macrocystis

Pyrifera (312-314). Its structure contains 1,4-linked -D-mannuronic (M) and -L-

guluronic (G) acid residues (Figure 3.1), arranged in a non-regular and block-wise

fashion along the chain (315-317). Alginate shows good biocompatibility, low

cytotoxicity and high-water content (high swelling ratio), mimicking the structure of

the natural extra-cellular matrix (318-322). These properties make alginate a suitable

material for wound dressings, drug delivery systems and soft tissue engineering

applications (323).

This Chapter investigates the preparation of alginate-based systems for UV bioprinting

applications. The material is characterised both before and after the curing process and

the effect of functionalization time on the rheological, mechanical, morphological,

swelling and degradation properties investigated. The effect of photoinitiator

concentration in terms of rheological and mechanical properties is also assessed and

discussed.

Figure 3.1: Alginate showing a linkage between the mannuronic and guluronic acid

(324).

81

Materials and methods

Synthesis of methacrylate alginate

Photocurable alginates were prepared through a functionalization mechanism with

methacrylate anhydride (MA) (325). Briefly, sodium alginate powder (W201502 ) 1%,

2% and 3% w/v (Sigma-Aldrich, UK) was dissolved in Dulbecco's Phosphate Buffered

Saline (DPBS) (Sigma-Aldrich, UK) and then mixed with methacrylate anhydride

(MA) (Sigma-Aldrich, UK) at 15 mL MA/g of alginate under vigorous stirring. The

pH of the solution was kept around 7.4-8.0 during the reaction time by adding 5M of

NaOH. Maintaining a higher pH during the reaction is crucial since higher pH

enhances the reaction among the amine and hydroxyl groups, which lead to higher

degree of modification (326). Two different reaction times (8 and 24 hours) were

considered to assess the effect of the reaction time on the degree of functionalisation.

After the chemical modification reaction, the polymeric solution was precipitated and

totally mixed in 100 mL of ethanol (100% ethanol) and dried in an oven overnight at

50 ºC. The precipitate polymer was dissolved in distilled water (diH2O), loaded in the

dialysis tubes membranes (SnakeSkin Dialysis Tubing 3.5K MWCO from Thermo

Fisher Scientific, UK), sealing both sides and dialyzed the solution against NaCl for 7

days with periodic water changes every day. The solution was freeze at -80ºC and the

polymer recovered by lyophilisation.

Characterisation of pre-polymerised methacrylate alginate

3.2.2.1 Nuclear magnetic resonance

The chemical structure of functionalized alginate was assessed through Nuclear

Magnetic Resonance Spectroscopy (1HNMR), using the B400 Bruker Avance III 400

MHz (Billerica, Massachusetts, USA). Polymeric materials were dissolved in

Deuterium oxide (D2O) (Sigma-Aldrich, UK), transferred to NMR tubes and the

spectra acquired with 128 scans.

3.2.2.2 Rheological characterisation

The rheological tests were performed to characterise the viscoelastic behaviour of both

pre-polymerised and polymerised alginate. The rheological assessment of pre-

polymerised alginate systems was carried out using the DHR2 TA Instrument (USA).

Samples were placed between two parallel plates and two different tests (rotational

and oscillation) were considered. Rotational tests were performed to evaluate the

viscosity and the material strength. In these tests it is assumed that the material flows

82

by applying a stress, being the response measured alongside time (temperature was

not considered in this work). A controlled stress was applied, and the resulting

movement measured. Oscillation tests were considered to evaluate the viscoelastic

behaviour of the material and dynamic moduli. The viscoelastic behaviour was

characterized by measuring the energy stored (storage modulus, Gʹ) in the material

during shearing and the energy subsequently lost (loss modulus, Gʺ). Shear strain was

controlled by varying the oscillation amplitude.

To assess the rheological changes during the photopolymerization process, the

rheological tests were carried out using the Bohlin Gemini system (Malvern

Instruments) equipped with the OmniCure® S1000 light source irradiating in the range

of 254-450 nm wavelength, and oscillation tests were considered. The light intensity

was 10 mW/ cm2.

The rheological behaviour of the materials is described by the following equation:

τ = ɳ ��n (3.1)

where τ is the shear stress (Pa), �� is the shear rate (s-1), ɳ the consistency index or

equivalent viscosity (Pa. s) and n is the power law index (dimensionless) that varies

as follows:

n<1: shear-thinning system;

n=1: Newtonian system;

n>1: shear-thickening system.

Hydrogel formation

Photocrosslinked alginate methacrylate hydrogels were prepared by dissolving 2%

w/v alginate-methacrylate with different photoinitiator concentration solutions 0.5%

w/v, 1% w/v and 1.5% w/v of VA-086 photoinitiator solution (2,2'-Azobis[2-methyl-

N-(2-hydroxyethyl) propionamide azo initiator (Wako Pure Chemical Industries,

USA). Crosslinked disks were produced by pipetting the alginate solution in an acrylic

mould with 8 mm diameter and 4 mm height. Photopolymerization was conducted

using a 365nm UV light (Model Dymax 2000-EC, Dymax Europe GmbH, Wiesbaden,

Germany) irradiating at 8 mW/cm2 during 8 min.

The photopolymerization process of alginate-methacrylate is a radicular

polymerization process started by the absorption of UV light by the photoinitiators,

followed by the generation of free radicals and a crosslinking chain reaction (327).

The mechanism is briefly presented in Figure 3.2.

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Figure 3.2: Schematic representation of the photopolymerization process of alginate-

methacrylate. a) After exposing the polymer solution to UV radiation, the

photoinitiators generated free radicals that react with the vinyl methylene starting the

crosslinking reaction; b) The reaction propagates with macro-radicals reacting with

unreacted carbon-carbon double bonds; c) At the end through a bimolecular

termination mechanism a 3D network of crosslinked hydrogel is formed. Adapted

from (328).

Characterisation of photocrosslinked hydrogels

3.2.4.1 Morphological characterisation

The morphology of the internal structure of the hydrogels was investigated through

Scanning Electron Microscopy (SEM), using the Hitachi S3000N VPSEM system.

Alginate samples were produced using cylindrical moulds (8 mm diameter and 4 mm

high). After hydrogel formation, samples were extensively washed in distilled water,

frozen at -80°C and lyophilized. Samples were fixed on stubs using a double-sided

adhesive tape and sputter coated with platinum sputter-coating.

3.2.4.2 Mechanical characterisation

Compression tests were performed at constant strain rate using the Instron 3344

machine equipped with a 10-N load cell (Instron, Buckinghamshire, UK). Crosslinked

alginate hydrogel disks were prepared as described in Section (3.2.3) and maintained

in distilled water at 37ºC following the protocol described by Jeon (325). After 24

hours of incubation, swollen alginate-methacrylate hydrogel disks were measured

using callipers to determine both the diameter and thickness, and unconfined

84

compression tests were performed on the hydrogel disks at room temperature, 0.5

mm/min of speed at a rate of 20% strain. Compressive modulus was determined from

the slope of stress versus strain plots and limited to the first 10% of strain as

recommended for cartilage applications (329).

3.2.4.3 Swelling and degradation characterisation

Alginate methacrylate hydrogel disks were frozen at -80 ºC, then, lyophilized and the

dry weights (Wi) were measured. Afterwards, the dried hydrogel samples were

immersed in distilled water (diH2O) and the same number of samples were also

immersed in Dulbecco’s Modified Eagle’s Medium - high glucose (DMEM) (Sigma-

UK) diluted with 10% FBS (Fetal Bovine Serum) (Thermofisher, UK) at pH 7 and

incubated at 37 ºC to reach an equilibrium swelling state. The distilled water and

DMEM were replaced every 1-2 days. Over the course of 3 weeks, samples were

removed from the DMEM/diH2O and the swollen hydrogel sample weights (Ws)

measured. The swelling ratio (Q) was calculated according to the following equation:

𝑄 =Ws

Wi (3.2)

where Wi is the dry weight and Ws is the weight of the swollen hydrogel sample. After

this, the swollen hydrogels were lyophilized and weighed again. The percentage of

mass loss was calculated as follows:

(Wi-Wd)/Wi x 100 (3.3)

where Wd is the weight after lyophilisation (N=3 for each time point).

Results and discussion

Alginate functionalization

The alginate modification with methacrylate anhydride was performed under standard

conditions, allowing the introduction of photo-reactive methacrylate groups into the

polymer backbone, as confirmed by 1HNMR analysis (Figures 3.3 to 3.5). Results

indicate the present of new characteristic peaks of methacrylate (MA) at 5.63 ppm and

6.09 ppm attributed to the methylene group in the vinyl bond, and a peak at 1.82 ppm

assigned to the methyl group, which are not present in the non-modified polymer,

showing that the polymers were successfully functionalized. Two different

functionalisation reaction times (8 and 24 hours) were also considered and the degree

of modification determined by dividing the relative integrations of methylene to

carbohydrate protons (330, 331). Results, presented in Table 3.1, show that the degree

85

of modification increases by increasing the reaction time, reaching a maximum value

of 33 % at 24 hours.

Table 3.1: Effect of reaction time on the modification degree of alginate methacrylate.

Composition (% w/v) Reaction time (h) Degree of modification (%)

2% 8 21

2% 24 33





corresponds to the methyl group.

86

Figure 3.4: a) Non-functionalized alginate 2% w/v, b) Functionalized alginate (2%

w/v) after 8 hours of reaction, c) Functionalized alginate 2% w/v after 24 hours of

reaction. The functionalization is confirmed by the presence of new peaks in the

spectra at 5.63 ppm and 6.09 ppm attributed to the methylene group and a peak at 1.82

ppm that corresponds to the methyl group.

87





corresponds to the methyl group.

Rheological behaviour of pre-polymerised alginate methacrylate systems

Figures 3.6 to 3.9 show the stress versus shear rate behaviour of solutions containing

1, 2 and 3% w/v of alginate-methacrylate reacted for 8 and 24 hours. Key rheological

parameters are presented in Tables 3.2 and 3.3. A non-linear behaviour is observed for

all samples. Samples containing 2% w/v obtained after 24 hours of reaction show a

clear Bingham behaviour (332). Results also show a decrease of the equivalent

viscosity by increasing the alginate concentration for samples obtained after 24 hours

of reaction, while no trend was observed for samples obtained after 8 hours of reaction.

The flow behaviour is also closer to a Newtonian fluid for samples containing high

concentrations of alginate. Based on these results the system containing 2% w/v of

88

alginate was selected as presents less variation of viscosity with the reaction time and

a clear shear-thinning behaviour being also less dependent with the reaction time

compared to the other systems.

Table 3.2: Rheological constants for solutions containing 15 mL of methacrylate

and different alginate concentrations, 24 h.

Composition (%w/v) Equivalent viscosity (Pa. s) Power law constant

1 4.36 ± 0.012 0.23

2 0.54 ± 0.014 0.68

3 0.16 ± 0.020 0.72

Table 3.3: Rheological constants for solutions containing 15 mL of methacrylate

and different alginate concentrations, 8h.

Composition (%w/v) Equivalent viscosity (Pa. s) Power law constant

1 0.09 ± 0.011 0.81

2 0.48 ± 0.012 0.48

3 0.08± 0.010 0.74

Figure 3.6: Stress versus shear rate profiles for alginate-methacrylate solutions

containing different alginate-methacrylate concentrations, reacted for 8 hours. a) 1%

w/v of alginate; b) 2% w/v of alginate; c) 3% w/v of alginate, and d) comparison of

all compositions.

89

Figure 3.7: Stress versus shear rate for solutions containing different alginate-

methacrylate concentrations reacted for 24 hours. a) 1% w/v of alginate, b) 2% w/v of

alginate, c) 3% w/v of alginate, and d) comparison of all compositions.


methacrylate concentrations reacted for 8 hours. a) 1% w/v; b) 2% w/v; c) 3% w/v d)

comparison of all compositions.

90


methacrylate concentrations reacted for 24 hours. a) 1% w/v; b) 2% w/v; c) 3% w/v

d) comparison of all compositions.

Figure 3.10 and 3.11 show the variation of storage modulus (G'), loss modulus (G''),

complex modulus and tanδ as a function of frequency and strain for systems containing

2% w/v alginate-methacrylate concentrations and different reaction times. The results

show that both the storage and loss modulus are higher in samples obtained after high

functionalisation times. In both cases the complex modulus increases with frequency,

while the variation of G' and G'' with strain shows a shift point after which G'' becomes

higher than G'. This shift occurs at high frequencies for alginate samples obtained with

high functionalisation times.

91

Figure 3.10: 2% wt. Alginate solution reacted for 8 hours. a) Storage and loss modulus

vs strain b) Storage and Loss modulus vs frequency; c) Complex modulus vs

frequency; d) tan δ vs frequency.

Figure 3.11: 2% wt. Alginate solution reacted for 24 hours. a) Storage and loss

modulus vs strain; b) Storage and Loss modulus vs frequency; c) Complex modulus

vs frequency; d) tan δ vs frequency.

92

Rheological changes during the photopolymerization process

Based on the characterisation of the pre-polymerised samples only systems containing

2% w/v of alginate were considered for photopolymerization studies. The curing

kinetics was assessed by monitoring the variation of Gʹ and G" at room temperature

through a controlled frequency of 1Hz. Photo-rheology was used to characterise the

curing process (photopolymerization) of functionalised alginate polymers obtained

after 8 and 24 hours of reaction time, mixed with 0.5% w/v, 1% w/v and 1.5% w/v of

VA-086 photoinitiator.

As observed from Figures 3.12 and 3.13 by increasing the curing time Gʹ increases

becoming significantly higher than Gʺ showing the liquid to solid phase change of the

material. It is also possible to observe that by increasing the photoinitiator

concentration the gelation time, which corresponds to the time point where the Gʹ

curve crosses the Gʺ curve, decreases. It is also possible to observe that by increasing

the amount of photoinitiator Gʹ seems to tend to a plateau which corresponds to a

vitrification stage of the curing process (333, 334). The possible explanation for this

observation is that, for the layer thickness considered in this study (~100 µm), the

photoinitiator concentration is approaching a critical value. By increasing the

photoinitiator concentration above the critical values, the polymerization occurs very

fast at the polymer surface, reducing the light penetration and, consequently, the

overall polymerization reduces.


photoinitiators functionalized for 8 hours a) 0.5% w/v of PI; b) 1% w/v of PI; c) 1.5%

w/v of PI; d) 0.05% w/v of PI.

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photoinitiators functionalized for 8 hours a) 0.5% w/v of PI ; b) 1% w/v of PI; c) 1.5%

w/v of PI.

Viscoelastic properties of formed hydrogels

Hydrogel disks (4 mm of height and 8 mm of diameter) were produced using an acrylic

machined mould. Samples of 2 % w/v of alginate-methacrylate obtained after both 8

and 24 hours of reaction time containing different concentrations of photoinitiator

(0.5, 1 and 1.5% w/v) were polymerised during 8 minutes under a light intensity of 8

mW/cm2. Produced disks were then assessed and both Gʹ and G" measured at room

temperature through a controlled frequency of 1Hz as presented in Figures 3.14 and

3.15. In the case of alginate-methacrylate samples obtained after 8 hours of reaction

time it’s possible to observe that there is no significant change of both Gʹ and G" (the

storage modulus is always higher than the elastic modulus) with the increase in the

photoinitiator concentration. In the case of alginate-methacrylate samples obtained

after 24 hours of reaction time, results show that by increasing the photoinitiator

concentration Gʹ and G" increases. In this case, it is also possible to observe that the

difference between Gʹ and G" increases with the increase of photoinitiator

concentration, which is associated to the high crosslinked density.

The viscoelastic nature of the crosslinked disks is also observed from the Gʹ and G"

versus strain graphs. In this case, it is possible to observe for all samples that, till a

critical strain value, the storage modulus is higher than the loss modulus. After the

strain critical value, the loss modulus become more significant due to the break of the

crosslinked network. In the case of alginate-methacrylate samples obtained after 24

hours of reaction time, the storage modulus is always higher than the storage modulus

of alginate-methacrylate samples obtained after 8 hours of reaction time. This is due

to the high-crosslinking density of the alginate-methacrylate samples obtained after 24

hours of reaction time.

94

Figure 3.14: 2% wt methacrylate alginate hydrogel with different concentration of

VA-086 functionalized for 8 hours a) 0.5% w/v of PI; b) 1% w/v of PI; C) 1.5% w/v

of PI.

95

Figure 3.15: 2% wt methacrylate alginate with different concentration of VA-086

functionalized for 24 hours. a) 0.5% w/v of PI; b) 1% w/v of PI; C) 1.5% w/v of PI.

Internal morphology of alginate hydrogels

The internal structure of crosslinked alginate is presented in Figure 3.16. SEM images

were obtained for samples containing 2% w/v of alginate prepared during 8 and 24

hours of reaction time and 1% wt of photoinitiator. Results show that the reaction time

influences the hydrogel morphology, with high reaction times being associated to

structures presenting both small pore size and number of pores. In addition, results

seem to indicate that long reaction times generate structures with more closed pores.

Results also show associated porosity of 94.5% ± 0.75 (292.5 µm ± 10 pore size) and

90.8% ± 0.21(252.5 µm ± 20 pore size) for 8 and 24 h respectively.

96

Figure 3.16: SEM images of crosslinked methacrylate-alginate hydrogel structures

obtained from alginate-methacrylate at different reaction times: 8 and 24 hours.

Mechanical characterisation

The mechanical performance of crosslinked alginate structures is presented in Figures

3.17 and 3.18. As observed, high compression moduli were obtained for crosslinked

disks produced with high functionalization times (13.16 kPa for 24 hours of reaction

time and 2.63 kPa for 8 hours of reaction time) and 0.5% w/v photoinitiator

concentration. These results can be explained by the high crosslinking density that

characterizes the structures obtained from alginate samples functionalized during long

reaction times and the corresponding internal morphology characterized by small size

and low number of pores. It also possible to observe that the mechanical properties

increase by increasing the photoinitiator concentration. For samples containing 1.5%

w/v of photoinitiator, compression moduli were obtained (75.4 kPa for 24 hours of

reaction time and 7.23 kPa for 8 hours of reaction time).


photoinitiator concentrations a) 0.5% w/v; b) 1% w/v; and c) 1.5 % w/v.

97


photoinitiator concentrations a) 0.5% w/v; b) 1% w/v; and c) 1.5 % w/v.

Swelling and degradation kinetics

The swelling and degradation behaviour of crosslinked alginate hydrogel disks are

presented in Figure 3.19. Results show that samples absorb both DMEM and distilled

water until reaching a state of equilibrium. This state is accomplished when the

osmotic pressure from the swelling and the elasticity of the hydrogel network is equal.

It is also possible to observe that crosslinked samples prepared with alginate-

methacrylate obtained after 24 hours of functionalization present low swelling ratio,

which shows that the degree of crosslinking controls the swelling properties of the

hydrogel. In all samples, the equilibrium status was reached at day 3. Crosslinked

alginate structures based on functionalized alginate prepared during 24 hour of

reaction time swelled up to 155%, while alginate structures based on functionalized

alginate prepared during 8 hour of reaction time swelled up to 190 %. Moreover, it is

also possible to observe that the level of functionalization of the pre-polymerised

alginate not only determines the internal morphology of the crosslinked structures but

also the degradation process as shown in Figures 3.19b and d.

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Figure 3.19: Swelling and degradation rate for 2% wt. functionalized alginate (24 and

8) hours reaction time in a) distilled water b); degradation rate in distilled water; c)

DMEM; d) degradation rate in DMEM.

Conclusion

This Chapter describes the synthesis and characterisation of alginate systems for UV-

based bioprinting applications. The alginate was successfully functionalized in the

presence of methacrylate in order to introduce the necessary number of unsaturation

allowing its crosslinking upon photopolymerization. Two different functionalization

reaction times were considered and photocurable systems containing different

photoinitiator concentrations prepared. From the results it is possible to conclude that

high functionalization reaction times originates crosslinked structures with less

porosity, smaller pores and a larger number of closed pores, less swelling, higher

degradation properties and higher mechanical stiffness. By increasing photoinitiator

concentration it was possible to observe an increase of mechanical properties and

gelation time. Moreover, for high values of photoinitiator concentration the reaction

tends to reach verification.

99

Chapter 4

Photocurable alginate

bioink development for

cartilage replacement

bioprinting†

† This Chapter based on the book Chapter: Hussein Mishbak, Enes Aslan, Glen Cooper, Paulo Bartolo,

“Photocurable alginate bioink development for cartilage replacement bioprinting”, Progress in Digital

and Physical Manufacturing. Lecture Notes in Mechanical Engineering, Springer. 2020. ISBN. 978-3-

030-29040-5

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Bioink design and assessment for tissue engineering replacement is a key topic of

research. This Chapter investigates suitable photocurable alginate bioink macromers

for bioprinting of 3D constructs for cartilage replacements. Alginate chemically

modified with methacrylate anhydride groups was assessed through different

techniques. 2% Alginate methacrylate (AlgMA) solutions containing different contents

of methacrylate at different reaction times were assessed. The methacrylate

functionalization and the degree of modification through various reaction times were

assessed through nuclear magnetic resonance (NMR), showing the ability to tune the

unsaturation degree by changing the reaction conditions. AlgMA synthesized using

different reaction times were also characterized by rheological analysis to investigate

the effect of reaction time on the rheological behavior. Results show that all Alginate

methacrylate samples exhibit a shear thinning behavior and the initial viscosity is

affected by the reaction time and methacrylate concentrations. These results indicate

that both the reaction conditions and methacrylate concentrations are of paramount

importance to control the extent methacrylation and rheological properties, which is

critical to design bioinks with appropriate properties of extrusion printing.

Biocompatibility and ecotoxicity were assessed through chondrocyte encapsulation.

Introduction

Articular cartilage is a non-vascularized tissue, covering the distal ends of the bones,

formed by a low density of cells called chondrocytes that maintain the structural and

functional integrity of the extracellular matrix (335-337). Cartilage is characterized by

a distinct zonal structure comprising, in descending order, the articular or superficial

surface, the middle (or transitional) zone, the deep zone, and the tidemark that

separates the non-calcified and calcified regions (181). However, it lacks inherent self-

repair capacity and hence the zonal structure and function are often irreversibly lost

following trauma and disease (338). As a result, cartilage defects are prone to develop

into osteoarthritis (OA), which is the predominant cartilage disease, characterized by

loss of cartilage, pain and debilitation (339). Despite the current therapeutic strategies

such microfracture, autologous chondrocyte implantation (ACI), matrix-assisted

autologous chondrocyte transplantation/implantation (MACT/MACI), autologous

matrix-induced chondrogenesis (AMIC), and osteochondral auto- and allografts (340,

341), articular cartilage repair is still a major clinical challenge. Tissue engineering

strategy emerged as a potential therapy through the fabrication of constructs

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mimicking the ECM structure of articular cartilage (342). Through this approach

different biocompatible and biodegradable materials with or without cells were

explored.

This Chapter focus on the use of photocurable alginate as a carrier of chondrocyte cells

for cartilage repair. Alginate material considered in this Chapter was previously

characterized in Chapter III and the rheological properties were further assessed in

order to guarantee good printability. Two different methacrylate concentration (15 and

25 mL for each gram of alginate) and two different reaction times (8 and 24 hrs) were

considered for each methacrylate concentration. Finally, cell-laden constructs were

produced, and cytotoxicity and cell proliferation tests performed.


Synthesis of methacrylated alginate

Methacrylate alginate was prepared following a published protocol (343). In this

study, 2% wt of sodium alginate powder (W201502) was completely dissolved in

Dulbecco's Phosphate-Buffered Saline (DPBS) (Sigma-Aldrich, UK). Afterwards,

methacrylate anhydride (Sigma-Aldrich, UK) was added to alginate solution. The pH

of the solution was kept around 7-8.0 during the reaction time by adding 5M of NaOH.

After the chemical modification, the polymer solution was precipitated with 100%

ethanol, dried in an oven overnight at 40-50 Cº, then diluted with distilled water and

purified through dialysis for 6 days (SnakeSkin™ Dialysis Tubing, 3.5K MWCO-

Thermo fisher, UK). The solution was frozen at -80ºC and the polymer recovered by

lyophilization. Two different methacrylate concentration (15 and 25 mL for each gram

of alginate) were used and two different reaction times (8 and 24 hrs) were considered

for each methacrylate concentration. Crosslinked discs were produced by dissolving

the functionalized alginate (AlgMA) polymer in a 1% w/v photoinitiator solution (PI)

VA-086, 2,2'-Azobis[2-methyl-N-(2-hydroxyethyl) propionamide azo initiator (Wako

Pure Chemical Industries, USA). The photocurable material was then pipetting into a

custom-made cylindrical Teflon mould (diameter: 8 mm; height 4 mm).

Photopolymerization was conducted using a 365 nm UV light (Dymax 2000-EC,

Dymax, Germany) irradiating at 8 mW/cm2 for 8 min.

HNMR

1HNMR was used to characterize the chemical modification of alginate. Methacrylate

alginate solution was dissolved in deuterium oxide D2O and placed in an NMR tube.

102

The 1HNMR spectra were recorded using the B400 Bruker Avance III 400 MHz

(Billerica, Massachusetts, USA) and the spectra acquired with 128 scans.

Rheological characterization

The rheological tests were carried out using the DHR2 TA Instrument (USA).

Rotational tests were considered, and all measurements were conducted at room

temperature.

Cell viability and proliferation

The cytotoxicity of the hydrogels was investigated through the encapsulation of

human chondrocyte cells (Cell Applications, USA) and a live/dead assessment

(Thermofisher Scientific, UK). Chondrocyte passage 4 were encapsulated at a density

of 0.75x106 cell/mL on sterilized hydrogel precursor solution. 160 µL of gel-cell

solution was pipetted into a 24-well custom mould and a UV radiation (365 nm of

wavelength emitted by a Dymax 2000-EC Lamp (Dymax 2000-EC, Dymax,

Germany) during 8 min with a light intensity of 8 mW/cm2. After 3 and 7 days the

cytotoxicity was analyzed by removing cell culture media, washing gently in DPBS

adding a live dead stain solution, and incubating for 30 min. The hydrogels were then

observed using a confocal fluorescence microscope (TCS-SP5, Leica, Germany). The

Alamar blue assay was also used to evaluate cell metabolism and proliferation after 1,

3, 5 and 7 days of culture.

Result and discussion

Chemical modification through 1HNMR

1HMNR results of non-functionalized and functionalized alginate are show in Figure

4.1. In the case of functionalized alginate results show the appearance of characteristic

peaks of methacrylate (MA) at 5.53 ppm and 6.10 ppm attributed to the methylene

group in vinyl bond, and a peak at 1.82 ppm associated to the methyl group. These

peaks are not presented in the nonmodified alginate (Figure 4.1a). The effect of

reaction time and methacrylate anhydride concentration on the degree of modification

(i.e., on the extent of methacrylation) determined by dividing the relative integrations

of methylene to carbohydrate protons (344), is shown in Table 4.1. In the case of 15

ml MA, Results show that the degree of modification increases by increasing the

reaction time, reaching a maximum value of 72.1% at 24 hours. However, by

increasing the methacrylate anhydride concentration up to 25 ml MA, the degree of

modification decreases, reaching 40.3% for 24h reaction time. We hypothesized that

103

the reduction on the degree of modification can be a result of the unreacted

methacrylate, hydrolysis in aqueous conditions and elevated temperature (40ºC)

during the reaction as well as possible reactions between methacrylate groups during

dialysis, as previously reported (234). These results suggest that the modification

efficiency with methacrylic anhydride is significantly affected by the methacrylate

concentration.

Figure 4.1: 1H NMR spectra of a) unmodified alginate and; b) 2% wt methacrylate

alginate (15 mL MA) after 8 hrs of reaction; c) 2% wt methacrylated alginate (25 mL

MA) after 8 hrs of reaction; d) 2% wt methacrylate alginate (15 mL MA) after 24 hrs

of reaction.

104

Table 4.1: Effect of reaction time on the degree of modification.

Composition (2 % w/v) Reaction time (hours) DoM (%)

AlgMA Low MA (15 ml) 8 55.9

AlgMA Low MA (15 ml) 24 72.1

AlgMA High MA (25 ml) 8 49.6

AlgMA High MA (25 ml) 24 40.35

Rheological characterization of alginate methacrylate

Figure 4.2 shows the variation of both the stress versus shear rate and the variation of

viscosity versus shear rate considering different reaction rates and methacrylate

concentration. Results show that all samples present a shear thinning behavior making

them suitable for printing. It is also possible to observe that by increasing the reaction

time the maximum stresses supported by the non-polymerised AlgMA increases. For

longer reaction times and high concentration of methacrylate mechanical properties

increases. Therefore, low methacrylate concentration (15 mL/g of alginate) will be

considered.

105

Figure 4.2: Rheological characterization of 2 % wt methacrylate alginate samples. a)

Stress vs shear rate profile for samples functionalized for 8 hrs reaction time with

different methacrylate concentrations; b) Stress vs shear rate profile for samples

functionalized for 24 hrs reaction time with different methacrylate concentrations; c)

Viscoelastic behaviour low methacrylate concentration (15 mL); d) Viscoelastic

behaviour high methacrylate concentration (25 mL).

Cell viability assessment

Cell viability assessment was examined using a live/dead assay (Figures 4.3 a, b, c and

d). As observed for both the 3D cell laden constructs and 2D cell culture after 3 and 7

days of cell encapsulation, cell viability is higher than 85% with only few dead cells

being observed. The results demonstrate that the functionalized hydrogel presents low

cytotoxicity being a suitable bioink materials. The metabolic and proliferation of

human chondrocyte cells was also assessed using the Alamar blue assay (Figure 4.3e).

Results show that the encapsulated cells with AlgMA hydrogels present good

proliferation rate at day 1 and 3, while after day 5 cell proliferation starts to decrease.

This can be explained by the confined space available for cells due to the absence of

any degradable sites. Moreover, these results show that photocrosslinkable AlgMA

hydrogels are biocompatible and adequately support the encapsulated cells.

106

Figure 4.3: Live/dead results at a) day 3, in AlgMA hydrogels; b) day 3, in the control;

c) day 7, in AlgMA hydrogels; d) day 7, in the control; and Alamar blue results e)

Fluorescence intensity as a function of time for different time point (1, 3, 5 and 7 days).

Note: in the live/dead results green corresponds to live cells and red to dead cells.

Conclusions

A bioink for cartilage applications is presented. Alginate was functionalized with

methacrylate groups and the effect of reaction time and methacrylation concentrations

on the degree of substitution and rheological properties were assessed through NMR

and rheological analysis, respectively. NMR data clearly show the ability to tailor the

methacrylation degree by changing the reaction time and methacrylate anhydride

concentration, enabling to tailor the properties of the bioink. Successful cell laden

constructs were produced, showing that the bioink is able to support cell proliferation

after 7 days of photopolymerization. No cytotoxicity emerging from both the alginate

and the photoinitiator was observed.

107

Chapter 5

Photocrosslinkable gelatin

methacrylate (GelMA)

hydrogels: Towards 3D

cell culture and

bioprinting for cartilage

applications†

† This Chapter based on the paper: Hussein Mishbak, Glen Cooper, Paulo Bartolo, “Photocrosslinkable

Gelatin Methacrylate (GelMA) Hydrogels: towards 3D cell culture and bioprinting for cartilage

applications”, Journal of Applied Polymer Science, submitted.

108

Tissue engineering This Chapter investigates the use of gelatin methacrylate (GelMA)

as a potential photocurable bioink for cartilage applications. The methacrylate

functionalisation process is discussed, and the level of functionalisation investigated.

Photocurable constructs are produced using ultraviolet radiation (8 mW/cm2) and 8

min of irradiation time. The influence of the functionalisation time, polymer

concentration, degree of modification and viscoelastic behaviour on the mechanical

strength and cell viability is also discussed. Results show that the increase of polymer

concentration, photoinitiator concentration and reaction time increase the

mechanical properties. Moreover, compression results show acceptable values for

cartilage applications (up to 76 kPa). Rheology shows that the GelMA presents shear

thinning behaviour and proper crosslinking structure to support encapsulated cells

during the printing process. High cell viability (85-90%) is also observed.

Introduction

Osteoarthritis (OA) is a critical clinical irreversibly cartilage defect developed after

trauma or disease (175-178, 345, 346). Currently, OA is treated through a wide range

of therapies including chondrocyte implantation (347), bone marrow stimulation

(347), mosaicplasty (348), osteochondral implantation (349) and autologous and

allogeneic tissue transplantation (350). However, these therapies present several

limitations such as the formation of inferior fibrocartilage, donor site morbidity,

limited prevention of continued tissue degeneration and in some cases two surgeries

are required (351-353). Cell laden approaches have the potential to solve these

limitations. In this case additive manufacturing technologies are used to create the

constructs using bioinks made of suitable hydrogels and encapsulated cells. However,

the design of these bioinks is complex requiring complex mechanical-rheological-

degradation-and biological relationships to be addressed. Moreover, selected

hydrogels must present a structure resembling the structure of the native tissue and

good bio affinity to support cell encapsulation and sustain cell differentiation,

proliferation, cell-cell networks and extra-cellular matrix (ECM) formation (354).

Cell laden cartilage constructs can also be produced using both chondrocytes or

mesenchymal stem cells (MSCs) (355). Several crosslinking methods have been

explored to create cell laden constructs using different types of bioinks and successful

results were obtained for different types of tissues such as skin, (356, 357), bone (358),

cardiac (359), and liver (360) . Among these methods, photoinitiated crosslinking is

109

particularly relevant as it is a relatively simple process, fast and can be performed

under physiological conditions (361). For cartilage, different photocrosslinkable

hydrogels were investigated (362, 363).

Among the different possible natural hydrogel materials, gelatin obtained from partial

hydrolysis of collagen and presenting arginyl-glycyl-aspartic acid (RGD) cell

adhesion motifs and matrix-metalloproteinase (MMP) degradation sites (356, 365), is

an ideal candidate for cartilage bioinks.

This Chapter investigates the use of gelatin as a potential photocurable bioink for

cartilage applications. The methacrylate functionalisation process is discussed, and the

level of functionalisation investigated. Photocurable constructs are produced using

ultraviolet radiation (8 mW/cm2) and 8 min of irradiation time. The influence of the

functionalisation time, polymer concentration, degree of modification and viscoelastic

behaviour on the mechanical strength and cell viability is also discussed.

Experimental

Synthesis of methacrylated gelatin

Photocrosslinkable gelatin was prepared through methacrylate functionalization

(Figure 5.1) following a protocol previously reported by Van den Bulcke et. al (366).

Briefly, powder of gelatin bovine skin type B (G9382) (Sigma-Aldrich, UK) was

dissolved at a concentration of 12.5% wt in Dulbecco's Phosphate Buffered Saline

DPBS (Sigma-Aldrich, UK) under a constant temperature of 50 °C. Then, 10% v/v

(10x molar excess) of methacrylate anhydride (MA) (Sigma-Aldrich, UK) was added

under vigorous stirring. The pH of the gelatin solution was kept around 7.4-8.0 during

the reaction period. Maintaining a higher pH during the reaction time is crucial to

enhance the reaction between the amine and hydroxyl groups, which lead to high

degree of modification. Different reaction times (2, 4 and 6 hrs) were considered.

Afterwards, the solution was diluted with of DPBS at 40 ºC, loaded at high

permeability dialysis tubes (12-14 kDa) (Spectra/Por 2, US) and purified against

distilled water for 7 days at 45 °C, changing water every day to remove the methacrylic

acid and other impurities. Finally, the pH of the final product was adjusted to 7~7.4

and frozen at -80 °C and the polymer recovered by lyophilization for 72 hours. The

final product (modified polymer) was frozen at -20 °C until further use.

110

Figure 5.1: Schematic crosslinking mechanism of gelatin and methacrylate.

Photocurable hydrogel solutions

After gelatin functionalisation, 10 and 15% wt of gelatin methacrylate (GelMA)

concentrations were diluted with a photoinitiator solution (VA-086) (2,2'-Azobis[2-

methyl-N-(2-hydroxyethyl) propionamide azo initiator (Wako Pure Chemical

Industries, USA). Two different photoinitiator concentrations were considered (1 and

1.5% wt). Crosslinked discs were produced by pipetting the photocrosslinkable

GelMA solution in an acrylic mould with a diameter of 8 mm, and 4 mm of height,

then irradiated with a UV light source (365 nm wavelength) using a UV lamp model

Dymax 2000-EC (Dymax Europe GmbH, Wiesbaden, Germany). Light intensity at

the material surface was fixed at 8 mW/cm2 and the material was irradiated for 8

minutes.

HNMR characterization

The chemical structure of functionalized gelatin was assessed using the B400 Bruker

Avance III 400 MHz. 150 mg of GelMA was dissolved in 600 μL of deuterium oxide

(D2O) (Sigma-Aldrich, UK), and then transferred to NMR tubes (VWR, UK) and the

spectra acquired with 128 scans.


The rheological assessment of non-polymerised gelatin solutions were carryout using

the discovery hybrid rheometer system DHR2 (TA Instrument, USA). Samples were

placed between the test pad and cone geometry (40 mm dimeter) fixed at the rheometer

head. Two different tests (rotational and oscillation) were considered and performed

at room temperature Rotational tests were considered to evaluate viscosity and

material strength. It is assumed that the material flows by applying a stress being the

response measured alongside time. A controlled stress was applied, and the resulting

movement measured at constant frequency (1 Hz). Oscillation tests (amplitude and

frequency sweep tests) were considered to evaluate the viscoelastic behaviour of the

material and the dynamic modulus. During the frequency sweep samples were exposed

to small-deformation oscillations covering a range of frequencies (0.01-1000 Hz) to

assess short and long-time scale responses to deformations. While, during an

111

amplitude sweep, the shear stress is varied, and the frequency is kept constant (1Hz).

The elastic or storage modulus (G') and the viscous or loss modulus (G") were

measured at room temperature. The viscoelastic behaviour was characterized by

measuring the energy stored (storage modulus, G') in the material during shearing and

the energy subsequently lost (loss modulus, G''). Shear strain was controlled by

varying the oscillation amplitude. For photocrosslinked hydrogels, the viscoelastic

properties were examined. Crosslinked hydrogels discs were placed between two

parallel plates and oscillatory tests were conducted using the DHR2 rheometer with a

20 mm diameter plate.

Scanning Electron Microscopy

The internal morphology of the hydrogel structure was observed using the Hitachi

S3000N VPSEM scanning electron microscopy system (Hitachi, Japan). Two

different hydrogel compositions were considered (10 and 15% wt of GelMA

concentrations), and both were irradiated at 8 mW/cm2 for 8 min. Gelatin methacrylate

hydrogels specimens were frozen at -80 °C for 3 days following lyophilization.

Finally, all samples were sputter-coated with gold by sputter-coating.

Mechanical characterization

Compression tests were conducted using the Instron 3344 machine (Instron, UK)

equipped with a 10-N load cell. Crosslinked gelatin hydrogel discs were prepared as

previously described and maintained in distilled water at 37 ºC. After 24 hrs of

incubation, swollen gelatin hydrogel discs were measured using callipers to determine

both the diameter and thickness. Unconfined compression tests were performed on

the hydrogel discs at room temperature and wet conditions, using a constant crosshead

speed of 0.5 mm/min, at a rate of 10% strain. Compressive young modulus was

determined from the slope of stress versus strain plots and limited to the first 10% of

strain as recommended for cartilage applications (329).

Swelling and degradation characterization

Crosslinked gelatin methacrylate hydrogel discs were frozen at -80 ºC, lyophilized and

then the dry weights (Wi) were measured. Then, dried hydrogel samples were

immersed in Dulbecco’s Modified Eagle’s Medium - high glucose (DMEM) (Sigma-

UK) diluted with 10% FBS (Fetal Bovine Serum) (Thermofisher, UK) at pH 7.4 and

incubated at 37 °C to reach an equilibrium swelling state. The DMEM was replaced

every 2 days. Over the course of 3 weeks, samples were removed from the DMEM

112

and the swollen hydrogel sample weights (Ws) measured. The swelling ratio (Q) was

calculated as follows:

Q=Ws/Wi (5.1)

After weighting the swollen hydrogels, the samples were freeze-dried and weighted

again to measure the weight after lyophilization (Wd). The percentage of mass loss

ratio was calculated according to the following equation:

((Wi-Wd)/Wi) *100 (5.2)

Measurements were performed in triplicate.

hADSCs Cell encapsulation

The gelatin methacrylate hydrogels (GelMA) were cytotoxically assessed, considering

both live dead and alamar blue assays. Human adipose-derived stem cells (hADSCs,

passage 4) (Stempro, Invitrogen, Waltham, MA, USA) were encapsulated with

hydrogel solution at 150000 cell/mL. 200 µL of the gel/cell (cell suspension) solution

was pipetted into custom-made cylindrical teflon moulds (diameter: 10 mm; height 4

mm), then exposed to 365 nm UV light, irradiating at 8 mW/cm2 for 8 min to cast the

encapsulated cell-hydrogel discs. Cell laden constructs were transferred into 24-well

plate and cell culture media (600 µl) was added until the hydrogel discs are fully

covered. The number of encapsulated cells was 30000 cell/well as recommended by

the well-plate manufacturer. The cytotoxicity was assessed at day 3, 7 and 19 by

removing cell culture media, washing gently in DPBS, and then live/dead stains

(Thermofisher Scientific, UK) were added. The Live/dead stain was prepared by

adding 5µl of the supplied calcein AM and 15µl of 2mM Ethidium homodimer (EthD-

1) stock solution to 10 mL of sterile, tissue culture-grade Dulbecco's phosphate-

buffered saline (DPBS), vortexing to ensure thorough mixing. Afterwards, 600 µl of

stain were added to each well to ensure that the hydrogel discs were fully covered.

Samples were kept in the incubator for 30 min prior to imaging. An inverted

fluorescence microscope (Leica DMI6000 B, Leica Microsystems, Wetzlar, Germany)

was used to obtain images of the hydrogel constructs. Images were processed and

analysed using ImageJ software. The Alamar blue assay was used to evaluate cell

bioactivity after 1, 3, 5 and 7 days of culture. For each time point, samples were

washed twice in DPBS and alamar blue solution 0.01% v/v (Sigma-Aldrich, UK) was

added to each well (600 µl of cell culture medium) and incubated for 4 hrs under

standard conditions (37 ºC, 5% of CO2 and 95% of humidity). After incubation, 150

113

µL of each sample solution was transferred to a 96-well plate and the fluorescence

intensity was measured at 540 nm excitation wavelength and 590 nm emission

wavelength with a spectrophotometer (Sunrise, Tecan, Männedorf, Zurich,

Switzerland). 2D cell culture was assessed as a control point. The experiments were

performed in triplicate (n=3).

Data analysis

All data are represented as mean ± standard deviation. Cell proliferation results were

subjected to one-way analysis of variance (one-way ANOVA) and post hoc Tukey’s

test using GraphPad Prism software version 8.0. Significance levels were set at p <

0.05.


Characterization of gelatin and GelMA by HNMR

Figure 5.2a shows the resonance characteristic peaks of non-functionalized gelatin.

The following peaks were identified: methyl group of amino acids, valine, leucine and

isoleucine (Peak 1 at 0.5-0.7 ppm), methylene residues of threonine (Peak 2 at 0.9-1.1

ppm), alanine (Peak 3 at 1.1-1.3 ppm), lysine (Peak 4 at 1.2-1.4 ppm, and Peak 7 at

2.6-2.7 ppm), arginine (Peak 5 at 1.6-1.8 ppm, and Peak 8 at 2.8-3.0 ppm), methylene

resonances of aspartate (Peak 6 at 2.2-2.5 ppm), and proline (Peak 9 at 3.2-3.4 ppm).

These peaks are also present in the functionalized gelatin (Figures 5.2b and c). New

peaks at 1.6-2.19 ppm, 5.69 ppm and 5.93 ppm correspond to methylene protons

(H2C=C(CH3)) and confirm the successful functionalization with methacrylic

anhydride. After confirming the chemical functionalization of gelatin, the degree of

functionalization (DoF) was determined following the procedure previously described

by Billet et al (366), and the results are presented in Table 5.1. As observed the degree

of functionalization increases with the increase of the functionalization reaction time.

After 2 hours GelMA presents low levels of C=C bonds on the material backbone,

which will originate less dense crosslinked polymer networks upon irradiation, while

after 6 hours GelMA presents high levels of C=C bonds on the backbone, which will

originate more dense crosslinked networks. No significant increase on the degree of

functionalization was observed between 4 and 6 hours of reaction. Therefore, only

samples obtained after 2 and 6 hrs of functionalization were considered for further

analysis.

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Table 5.1: Effect of reaction time on the degree of functionalization.

Figure 5.2: 1H NMR spectra for non-functionalized and functionalized gelatin. (a)

Non-functionalized gelatin (b) GelMA functionalized for 2 hrs; c) GelMA

functionalized for 4 hrs; d) GelMA functionalized for 6 hrs.

Reaction time (hrs) Degree of functionalization (%)

2 40

4 52

6 54

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Prepolymer rheology

The stress versus shear rate and the variation of viscosity with shear rate for pre-

polymerised GelMA samples are shown in Figure 5.3. As observed, the materials

exhibit a shear-thinning behaviour. Results also show that the viscosity of the polymer

solution is dependent of the shear rate, decreasing by increasing the shear rate. This is

an important characteristic for 3D bioprinting applications as less pressure is required

during the printing process. It’s also possible to observe that the polymer concentration

has an impact on the maximum stress for a specific shear rate with high values being

obtained for long functionalization reaction times. Results show that for both GelMA

compositions viscosity decreases by increasing the functionalization time. For the

same functionalization time and different material compositions results show no

significant differences of viscosity for low shear rates, but for high shear rates, the

viscosity is significantly low in the case of high concentration of gelatin methacrylate

Figure 5.3: Rheological behaviour of 10 and 15% wt of GelMA samples considering

2 and 6 hrs of functionalization time. (a)10% wt of GelMA functionalized for 2 hrs;

(b) 10% wt of GelMA functionalized for 6 hrs; (c) 15% wt of GelMA functionalized

for 2 hrs; (d) 15% wt of GelMA functionalized for 6 hrs.

116

Since the polymers exhibit a shear-thinning behaviour, a power law model can be used

to fit the experimental data:

τ = 𝜂γn (5.3)

where τ is the shear stress (Pa), γ is the shear rate (s-1), 𝜂 the consistency index or

equivalent viscosity (Pa. s) and n is the power law index (dimensionless).

Table 5.2 shows the values of the power law coefficients for the different materials. A

good approximation (R2 around 0.97) was obtained between the power law and the

experimental data.

Table 5.2: Power law coefficients for different reaction times and GelMA

concentration.

Polymer concentration - Reaction time (h) n value R2

10% wt - 2h 0.4978 0.9652

10% wt - 6h 0.5613 0.9834

15% wt - 2h 0.825 0.999

15% wt - 6h 0.5519 0.9868

Viscoelastic changes for pre-polymerised solutions were assessed using oscillation

tests (Figure 5.4).

For both concentrations, results show that both the storage and loss modulus are high

when the reaction time is long (6 hrs).

Results for GelMA samples functionalized during 2 hrs show that at low frequencies

the storage modulus is higher than the loss modulus. While for high frequencies, the

loss modulus increases and become higher than the storage modulus. For GelMA

samples functionalised during 6 hrs it is possible to observe that the loss modulus is

higher than the storage modulus for all frequencies.

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Figure 5.4: Oscillatory tests: storage and loss modulus of 10% wt GelMA for 2 and 6

hrs; (a) Amplitude sweep for 10 and 15% wt GelMA functionalized for 2hrs; (b)

Frequency sweep for 10 and 15% wt GelMA functionalized for 2hrs; (c) Amplitude

sweep for 10 and 15% wt GelMA functionalized for 6hrs (d) Frequency sweep for 10

and 15% wt GelMA functionalized for 6hrs.

Figures 5.5 and 5.6 show the variation of G' and G'' as a function of strain rate and

frequencies of crosslinked hydrogels. Amplitude sweep results show that for the two

different functionalisation times, G' is higher in the case of samples obtained from

solutions containing higher concentration of GelMA. For the same GelMA

concentration, G' also increases by increasing the functionalization time. The increase

of G' is associated with the increase in the crosslinking density. In all cases it is

possible to observe a critical strain rate value. Bellow this threshold strain, G' is higher

than G'', while above the threshold G'' becomes higher than G'. In the case of frequency

sweep, results show that the G' is always higher than G'' showing a dominant solid like

behaviour.

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of frequency for 10 and 15% wt GelMA and 1.5% wt of PI.

119






of frequency for 10 and 15% wt GelMA and 1.5% wt of PI.

Morphological characterization

The internal pore size is crucial as it determines the internal free space to encapsulate

cells and to support cell growth and proliferation. In the case of cartilage bioinks for

3D bioprinting the references are chondrocytes cells and human adipose-derived stem

cells presenting around 20 m of diameter (367, 368). Figure 5.7 shows the internal

morphology of crosslinked GelMA hydrogels. Results show that crosslinked

structures obtained from solutions containing 10% wt of GelMA (2 and 6 hrs of

reaction time) and 1% wt of photoinitiator present pore sizes with 100 to 150 m of

dimeter (Figures 5.7 a and b). The increase of photoinitiator concentration reduces

both pore size and number of pores Figures 5.7c and d as it contributes to high

crosslinking density. Polymer solutions containing 15% wt of GelMA result in high

density crosslinked structures with small pores (23 to 25 m) (Figures 5.7e and f) or

almost no pores (Figures 5.7g and h). As observed the most suitable system for cell

120

encapsulation seems to be the one based on solutions containing 10% of GelMA and

1% wt of photoinitiator as it allows more free space to accommodate cell encapsulation

and proliferation. Therefore, this is the composition further considered.

Figure 5.7: Cross-section images of GelMA crosslinked structures considering

different GelMA concentrations, reaction times and PI concentration; (a) 10% of

GelMA, 2 hrs of functionalization time and 1% wt of PI; (b) 10% of GelMA, 6 hrs of

functionalization time and 1% wt of PI; (c) 10% of GelMA, 2 hrs of reaction time and

1.5% wt of PI; (d) 10% GelMA, 6 hrs of reaction time and 1.5% wt of PI; (e) 15% of

GelMA, 2 hrs of reaction time and 1% wt of PI; (f) 15% of GelMA, 6 hrs of reaction

time and 1% wt of PI; g) 15% of GelMA, 2 hrs of reaction time and 1.5% wt of PI; (h)

15% of GelMA, 6 hrs of reaction time and 1.5% wt of PI.

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Mechanical characterization

Compressive modulus of crosslinked GelMA discs were determined one day after

equilibrium swelling and the results are presented in Figure 5.8. As observed, high

compression modulus is obtained with high functionalisation times and high

photoinitiators concentrations. For the same functionalisation time and photoinitiator

concentration, compressive modulus increases by increasing the gelatin concentration.

These results can be explained by the high crosslinking density that characterises the

structures obtained from solutions containing high GelMA concentration,

functionalised during long reaction times and containing high concentration of

photoinitiators and the corresponding internal structure characterised by small size and

low numbers of pores.

Figure 5.8: Unconfined compression tests for gelatin methacrylate hydrogels samples

with different photoinitiator concentrations a) 10% wt GelMA functionalized for 2h

containing 1% wt of PI; (b) 10% wt GelMA functionalized for 2h containing 1.5% wt

of PI; (c) 10% wt GelMA functionalized for 6h containing 1% wt of PI; (d) 10% wt

GelMA functionalized for 6h containing 1.5% wt of PI.

Swelling and degradation characterization

The swelling and degradation behaviour of crosslinked GelMA disks are presented in

Figure 5.9. Results show that all samples absorb DMEM until reaching a state of

equilibrium. This state is accomplished when the osmotic pressure from the swelling

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and the elasticity of the hydrogel network is equal. In all samples, the equilibrium

status was reached at day three. It is also possible to observe that crosslinked samples

prepared with GelMA obtained after 6 hrs of functionalization present low swelling

ratio, which shows that the degree of crosslinking controls the swelling properties of

the hydrogel. Crosslinked GelMA structures based on GelMA prepared during 6 hours

of reaction time swelled up to 14%, while GelMA structures based on functionalized

gelatin prepared during 2 hour of reaction time swelled up to 20% (Figure 5.9a).

Moreover, it is also possible to observe that the level of functionalization of the pre-

polymerised gelatin not only determines the internal morphology of the crosslinked

structures but also the degradation process (Figure 5.9b). Results show that samples

produced using GelMA obtained considering high reaction times and solutions

containing high polymer concentration (15% wt) show lower degradation rates (20%

of mass loss for crosslinked GelMA obtained from gelatin functionalised during 6 hrs

and solutions containing 15% wt of GelMA, and 35% of mass loss for crosslinked

GelMA obtained from gelatine functionalized during 2 hrs and solutions containing

10% wt of gelatin).

Figure 5.9: Swelling and degradation profiles as a function of functionalization time

for different functionalized gelatin methacrylate systems; a) Swelling in DMEM ; b)

degradation rate in DMEM.

Cell viability and proliferation

The live/dead assay was used to investigate the potential toxicity of both GelMA

hydrogels, photoinitiator concentration and the effect of irradiation conditions on the

encapsulated cells. Results after 3, 7 and 19 days of cell-encapsulation are presented

in Figure 5.10. Results show that most of the hADSCs are alive (green) with few dead

cells (red) at day 3, indicating good biocompatibility for the produced materials. At

day 7, cells show an elongated cell morphology and a cell-cell networking can be

observed at day 19, which is more intensive in samples prepared with GelMA obtained

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at high functionalization times. This might be attributed to the higher functionalization

degree of GelMA obtained after 6 hrs of functionalization.

Cell proliferation was assessed using the alamar blue assay and the results are

presented in Figure 5.11, confirming the bioactivity of encapsulated hADSCs. At day

1, the 6 hrs GelMA hydrogel shows higher fluorescence intensity in comparison to the

2 hrs hydrogel group. However, this difference reduces after day 1. At day 5, both

GelMA groups shares similar fluorescence intensity and 2hrs GelMA hydrogel group

shows higher intensity at day 7. This might be attributed to the faster degradation of

2hrs hydrogel group which releases more cells attaching to the cell culture plate. These

results, which are aligned with the results presented in Figure 5.10c, shows that 6 hrs

hydrogel presents better cell affinity for a long cell culture time.

Figure 5.10: Live/dead images of adipose-derived mesenchymal stem cells

encapsulated in samples based on 10% wt GelMA dissolved at 1% wt of PI solution

and exposed to UV light (Light intensity of 8 mW/cm2) for 8 min a) After 3 days of

encapsulation; b) After 7 days of encapsulation; c) After 19 days of encapsulation.

Note: Green corresponds to live cells and red to dead cells.

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Figure 5.11: Alamar blue fluorescence intensity at different time points (1, 3, 5 and 7

days) considering samples functionalized at two different reaction times (2 and 6 hrs)

Conclusion

Photocrosslinkable hydrogels have shown great potential for cartilage repair, due to

their unique properties. GelMA systems were fully characterised in terms mechanical,

photopolymerization process, degradation, swelling and biological properties. Results

show the ability to tune these characteristics by controlling the functionalization

process. Moreover, the biological affinity between gelatin and cells were confirmed

being possible to observe that both material systems and processing conditions do not

induce any negative effect on cell survival. High cell attachment and proliferation was

observed. Results demonstrates the viability of using GelMA as a bioink for cartilage

applications.

Time (days)

0

20000

40000

60000

6h

Control

2h

D3D1 D5 D7

*

*

Flu

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Chapter 6 Development and

characterisation of a

photocurable alginate-

gelatin for cartilage

applications†

† This Chapter based on the paper: Hussein Mishbak†, Cian Vyas†, Marie O'Brien, Glen Cooper, Paulo

Bartolo. “Development and characterisation of a photocurable alginate-gelatin for cartilage

applications”, Biomaterials Science and Engineering Journal, in preparation.

126

This Chapter focus on the design of three different photocrosslinkable hydrogels based

on alginate and gelatin for cartilage applications. The synthesis, morphological,

rheological, mechanical, and biological characterisation of AlgMA, GelMA, and

hybrid alginate-gelatin Alg-Gel(MA) methacrylate hydrogels are detailed. Results

show that the hybrid hydrogels display superior mechanical properties than the single

component hydrogels and present a more stable degradation behaviour. The hybrid

hydrogels also demonstrate high cell viability, improved cell adhesion, and increased

production of cartilage tissue markers such as collagen and glycosaminoglycan

GAGs. Both AlgMA and GelMA materials considered in this Chapter were

functionalized as described in Chapters III and V.

Introduction

Articular cartilage is a tissue which covers the ends of bones at joints and provides a

smooth and lubricated surface for articulation during movement and mechanical load

transmission (17, 369). The tissue has low metabolic activity and lacks a vasculature

thus disease or trauma can severely impact tissue functionality (369). Articular

cartilage heals poorly, typically forming fibrocartilage, and current therapies have

limited clinical efficacy in halting the degradation of the tissue which ultimately can

result in the requirement of a total joint replacement implant (370, 371). This has

motivated the development of tissue engineering strategies utilising biomaterials and

scaffolds to promote cartilage regeneration (85, 110, 372, 373). A promising class of

biomaterials are hydrogels, three-dimensional (3D) hydrated polymer networks, which

have properties that resemble the extracellular matrix (ECM) of biological tissues

(120, 374-377). Hydrogels have been investigated for tissue engineering applications

due to their biocompatibility, oxygen and nutrient permeability, and the ability to tailor

their properties (372, 378-381). Biodegradable and biocompatible synthetic and

natural based hydrogels have been explored to promote the repair and regeneration of

different tissues such as skin, nerve and cartilage (382-384).

Synthetic hydrogels have been widely investigated as a potential suitable biomaterial

for articular cartilage tissue engineering applications due to their ability to be

engineered to mimic the ECM and tailored to have similar properties in terms of

mechanical, morphological and biological properties (385-388). However, synthetic

hydrogels have limitations regarding their biocompatibility and biodegradability due

to the lack of inherent adhesion and degradation sites which requires the incorporation

127

of motifs such as the tripeptide Arg-Gly-Asp (RGD) sequences and matrix

metalloproteinases (389).

Alternatively, natural based hydrogels are attractive biomaterials due to their

biocompatibility and biodegradability and have been explored for cartilage tissue

engineering applications (390-393). Particular relevant are alginate and gelatin.

Alginate has been investigated due to its low cost, biocompatibility, hydrophilic

structure, stiff polymer backbone, ease of modification and processing (394, 395).

Alginate hydrogel based-scaffolds can be produced via ionic-crosslinking, phase

transition, free radical polymerisation, and click conjugation (396). Free radical

photopolymerization has gained tremendous attention due to the ease of use and rapid

gelation kinetics. Furthermore, photopolymerisation enables polymer solutions (pre-

polymerised monomers) containing cells or biomolecules to be printed using 3D

bioprinting technologies for the development of scaffolds or using in vivo minimally

invasive delivery technique (e.g. injection) with subsequently crosslinking through

light exposure (113, 383, 397). However, alginate has significant limitations which

limits its applicability as a single material for tissue engineering applications such as

the lack of cell binding motifs and slow gelation speed which affects gel uniformity

and strength (396). Alternatively, gelatin, obtained from the hydrolysis of collagen,

has been investigated due to its biocompatibility, biodegradability, promotion of cell

adhesion and proliferation (398, 399). The main advantages of gelatin are its

biocompatibility since it has high cell affinity and cell adhesion sites are abundant

(RGD peptides) and its ability to be degraded by native enzymatic processes (400,

401). However, gelatin has limitations such as poor mechanical properties (brittle) and

short biodegradation profile (401). Due to their limitations both alginate and gelatine

are not totally suitable for tissue engineering applications and for cartilage applications

in particular. However, the combination of both materials has the potential to

overcome their individual limitations.

This Chapter investigates the potential of using a hybrid AlgMA and GelMA

photocrosslinking hydrogel system for cartilage applications. The synthesis,

morphological, rheological, mechanical, and biological characterisation of AlgMA,

GelMA, and hybrid alginate-gelatin Alg-Gel(MA) methacrylate hydrogels are

detailed.

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Photocrosslinkable polymer preparation

Photocrosslinkable polymers were synthesised by reacting alginate and gelatin with

methacrylic anhydride to introduce methacrylate moieties into the polymer backbone

(326, 402, 403).

Alginate methacrylate preparation (AlgMA)

Alginate methacrylate AlgMA was prepared by dissolving sodium alginate (Sigma-

Aldrich, UK) at a concentration of 2% wt. in an aqueous solution containing

Dulbecco's phosphate buffered saline (DPBS) and dimethyl sulfoxide (DMSO),

(Sigma Aldrich, UK) at a ratio of 85:15. The dissolved solution was then mixed with

20% v/v methacrylic anhydride (MA) (Sigma Aldrich, UK) for 24 h, at room

temperature (RT), as previously described (404). The pH of the solution was

maintained between 7.4-8.0 by adding 5M NaOH (Sigma Aldrich, UK). After

chemical modification, the polymer solution was precipitated with 100% ethanol

(absolute ethanol ≥99.8%, AnalaR NORMAPUR®, WVR, UK), dried in an oven

overnight at 40-50°C, and purified through dialysis tubes (3.5 kDa) (Spectra/Por 2,

Fisher Scientific, USA) for 5 days. The solution was frozen at -80°C for 2 days and

the polymer recovered by lyophilisation for 3 days.

Similarly, GelMA was prepared by dissolving gelatin bovine skin type B (Sigma

Aldrich, UK) at a concentration of 12.5% wt. in DPBS at a temperature of 45°C. After

complete dissolution of gelatin a 10% v/v MA solution was added under vigorous

stirring. The pH was kept at 8 during the reaction time of 6 h. The solution was purified

through dialysis tubes (12-14 kDa) (Spectra/Por 2, Fisher Scientific, USA) for 5 days.

Then, the solution was frozen at -80ºC for 2 days and the polymer was recovered by

lyophilisation for 3 days.

Hydrogel formation

6.2.3.1 Photoinitiator solution preparation

Photoinitiator solution was prepared by dissolving the photoinitiator powder, 2,2'-

Azobis[2-methyl-N-(2-hydroxyethyl) propionamide azo initiator (VA-086) (Wako

Pure Chemical Industries, USA) in DPBS (Sigma-Aldrich, UK) at 40°C for 30 min,

to a final concentration of 1% wt. The PI concentration was previously investigated

for both alginate and gelatin single systems and the results presented in chapters III

and V.

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6.2.3.2 Single hydrogel systems

The AlgMA and GelMA hydrogels are produced by dissolving 2% wt. and 10% wt.

respectively of the lyophilised polymer in the 1% wt. photoinitiator solution.

Crosslinked hydrogel discs were produced by pipetting the AlgMA and GelMA

solutions into custom-made cylindrical acrylic moulds (diameter: 8 mm, height: 4 mm)

and irradiated with ultraviolet (UV) light (365 nm wavelength) using a UV lamp

(Dymax 2000-EC, Dymax Europe GmbH, Germany). The light intensity at the

material surface was fixed at 8 mW/cm2 and samples were irradiated during 8 min.

The exposure conditions were selected using 365 nm and between 5–20 mW/cm2 for

2–20 min since these low UV light intensities have been reported to have a lower

impact on cell viability, proliferation and gene expression (405) (Figure 6.1).

Figure 6.1: Photocrosslinking mechanism to form a) AlgMA and b) GelMA hydrogels

using UV irradiation.

6.2.3.3 Hybrid hydrogel system

The hybrid hydrogel systems were synthesised by mixing 2% wt. AlgMA and 10%

wt. GelMA solutions at ratios of 75:25 AlgMA/GelMA, 50:50 AlgMA/GelMA, and

25:75 AlgMA/GelMA. These alginate-gelatin methacrylate hybrid hydrogels (Alg-

Gel(MA)) were crosslinked as previously described (Figure 6.2).

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Figure 6.2: Photocrosslinking of hybrid alginate-gelatin systems when exposed to UV

irradiation.

Structural conformation of the hybrid hydrogels

Proton nuclear magnetic resonance (NMR) spectroscopy (Bruker 400, Bruker

Corporation, USA) was used to assess the degree of methacrylate modification for all

three systems (AlgMA, GelMA, and Alg-Gel(MA)). Samples were prepared by

dissolving 150 mg of each polymer in 750 µl deuterium oxide (D2O) (Sigma Aldrich,

UK) and then transferred to NMR tubes (VWR, UK). The spectra were acquired at

400 MHz with 128 scans. The degree of methacrylation was calculated following the

procedure previously described by Billet et al (234).

Rheological characterisation

Rheological characterisation was used to determine the general flow behaviour and

viscoelastic characteristics of the functionalized pre- and polymerized hydrogels.

Rotational and oscillatory rheology (Discovery Hybrid Rheometer, TA Instruments,

USA) was considered. All samples (n=3) were measured at room temperature (RT).

Rotational tests were performed by loading the polymer samples using a cone and

plate geometry (60 mm diameter and 200 µm gap size), applying a torque to the top

plate exerts a rotational shear stress on the material to flow and the resulting shear rate

and movement are measured. The rotational speed depends on the viscosity computed

by means of stress and shear rate. By means of the flow curve it is possible to measure

the viscosity as a function of shear rate (0.01 – 1000 s-1).

The oscillatory tests were performed to measure the viscoelastic behaviour and the

stiffness of the specimens and dynamic moduli. Amplitude and frequency sweeps were

considered. A range of frequencies (0.01-100 Hz) was used to assess short and long-

time scale responses to deformations at a constant 20% strain, while the strain rate

varies during the amplitude sweep test, and the frequency was constant at 1 Hz. The

storage (G') and the loss (G") modulus were measured.

For photocrosslinked hydrogels, oscillatory tests were performed to evaluate the

hydrogel stiffness by subjecting the hydrogel to a shear stress. The hydrogel samples

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were carefully placed onto the surface of the plate (20 mm in diameter) lowered to a

500 µm gap distance, shear stress was applied, and then, G' and G'' were measured as

a function of strain (0.1-500%).

Structure morphology and porosity

The hydrogel’s internal morphology was investigated using scanning electron

microscopy (SEM, S3000N, Hitachi, Japan). Cylindrical hydrogels samples (8 mm

diameter, 4 mm height) were frozen at -80°C for 2 days then lyophilised to recover the

dried hydrogel constructs. The samples were cut to obtain a cross-section and sputter

coated with gold prior to imaging.

Porosity is an important property of the hydrogel’s structure due to its influence on the

mechanical performance of the scaffold; fluid permeability; the supply of oxygen,

nutrients and removal of waste products; and the free volume provided for cells to

migrate and proliferate. The percentage of porosity of the hydrogels (n=3) was

calculated using a gravimetric method. Dry samples (n=3) were obtained by

lyophilisation and then the samples were placed in a desiccator under vacuum (VWR,

UK) for three days and the volume was determined by using a calliper. The porosity

was determined using the following equation:

𝑃𝑜𝑟𝑜𝑠𝑖𝑡𝑦 (%) = (1 − 𝑃ℎ𝑦𝑑𝑟𝑜𝑔𝑒𝑙

𝑃𝑚𝑎𝑡𝑒𝑟𝑖𝑎𝑙) × 100 (6.1)

where Phydrogel is the experimental density of the 3D hydrogel construct and Pmaterial is

the theoretical density of the polymer. The theoretical density for the hybrid systems

was calculated as follows:

𝑃𝑚𝑎𝑡𝑒𝑟𝑖𝑎𝑙 = 𝑋𝐴𝑙𝑔𝑀𝐴𝑃𝐴𝑙𝑔𝑀𝐴 + 𝑋𝐺𝑒𝑙𝑀𝐴 𝑃𝐺𝑒𝑙𝑀𝐴 (6.2)

where XAlgMA and XGelMA indicate the volume fractions and the densities of alginate, and

gelatin, PAlgMA and PGelMA, are 1.64 g/cm3 and 0.98 g/cm3, respectively.

The experimental apparent hydrogel density, Phydrogel, can be determined according to

the following equation:

𝑃ℎ𝑦𝑑𝑟𝑜𝑔𝑒𝑙 =𝑚𝑎𝑠𝑠

𝑣𝑜𝑙𝑢𝑚𝑒 (6.3)

where mass is the dry weight of the specimen and the volume is based on a

measurement of the dried sample.

The average pore size of the hydrogels was semi-quantified using ImageJ software

analysis of the SEM images through calculation of the Feret's diameter, which is an

approximation method to calculate the pore size dimeter (406).

132

Swelling and degradation kinetics of the hybrid hydrogels

The hydrogel discs were produced as previously mentioned in section (6.2.3) (n=3, at

each timepoint), frozen at -80°C for 2 days, and then lyophilised for 3 days to recover

dry hydrogel structures. After lyophilisation, the initial dry disc weights were

measured (Wi). The dry hydrogel samples were immersed in Dulbecco’s Modified

Eagle’s Medium (DMEM) (Sigma Aldrich, UK) supplemented with 10% fetal bovine

serum (FBS) (Thermofisher, UK) at pH 7, and incubated at 37°C to reach an

equilibrium swelling state. The swelling weights (Ws) were measured over the course

of 20 days and the DMEM was changed every two days.

The swelling ratio (Q) was calculated as follows:

𝑄 = 𝑊𝑠

𝑊𝑖 (6.4)

Where Wi is the initial dry weight and Ws is the swollen hydrogel sample weight.

For the degradation kinetics, the swollen hydrogels samples were frozen and

lyophilised at each timepoint and the dry weight (Wd) measured. The mass loss

percentage was calculated using the following equation:

𝑀𝑎𝑠𝑠 𝑙𝑜𝑠𝑠 (%) = (𝑊𝑖−𝑊𝑑

𝑊𝑖) × 100 (6.5)

Mechanical assessment

The mechanical behaviour of the crosslinked hydrogels was characterised through

unconfined compression testing using Instron 3344 single column system with a 10 N

load cell (Instron 3344, Instron, UK). Hydrogel discs (n=3) were prepared as

previously described using a custom-made cylindrical acrylic mould with a 1:1

dimension ratio (diameter 10 mm and height 10 mm) and incubated in DPBS at 37°C

for 24 h. The hydrogel discs were measured after 24 h to determine the diameter and

thickness, and unconfined compression tests were performed on the hydrogel discs at

RT. The compressive strain rate was fixed at 0.5 mm/min, and samples were loaded

up to a maximum strain of 20%. The compressive Young’s Modulus (E) was

calculated from the slope of the initial linear region of the stress-strain curve (elastic

deformation region) by calculating the tangent modulus from the curve fit equation at

20 % strain.

Biological assessment

The photocrosslinked hydrogels were biologically evaluated through the

encapsulation of human chondrocyte cells (Cell Applications, USA) and assessed for

133

cell viability, proliferation, and ECM production. The AlgMA, GelMA, and Alg-

Gel(MA) 50:50 photocrosslinked hydrogels were selected for evaluation.

6.2.9.1 Cell culture and encapsulation

Human chondrocyte cells were cultured until passage four in a complete human

chondrocyte growth medium (Cell Applications, USA). Cells were harvested and

encapsulated at density of 0.6 x 106 cells/mL in the hydrogel solution. 160 µL of the

hydrogel/cell solution was pipetted into a custom-made cylindrical acrylic mould

(diameter 8 mm and height 4 mm) and exposed to UV light (wavelength 356 nm),

irradiated at 8 mW/cm2 for 8 min to cause gelation of the encapsulated hydrogel/cell

solution. Cell-laden hydrogels (n=3) were transferred into 24-well plates and cell

culture media added

6.2.9.2 Cell viability, proliferation, and metabolic activity

The cytotoxicity of the hydrogels was observed using a live/dead assay kit

(Thermofisher Scientific, UK). Cell viability was observed on day 3, 7 and 16 by

removing cell culture media, washing gently in DPBS, adding a 2 µM calcein-AM and

4 µM ethidium homodimer-1 solution, and incubating for 25 min. The hydrogels were

then observed using a confocal fluorescence microscope occupied with fluorescence

filter sets (A, I3 and N2.1) (TCS-SP5, Leica, Germany). Images were processed and

analysed using ImageJ software (n=3) (407).

Cell proliferation and metabolic activity was evaluated by using the resazurin (Alamar

blue) assay at day 1, 3, 5, 7 and 14 days of culture. At each time point, samples (n=3)

were washed twice in DPBS and 75 µL of Alamar blue solution 0.01% (wt. ) (Sigma-

Aldrich, UK) was added to each well in 750 µl of cell culture medium and incubated

for 4 h under standard conditions (37°C, 5% CO2, and 95% humidity). After

incubation, 150 µL of each sample solution was transferred to a 96-well plate and the

fluorescence intensity was measured at 540 nm excitation wavelength and 590 nm

emission wavelength with a plate reader (Infinite 200, Tecan, Männedorf, Zurich,

Switzerland). 2D tissue culture plastic (TCP) was assessed as a positive control for

both cell viability and proliferation.

Histological analysis

Histological analysis of the cell encapsulated hydrogels was investigated at day 14,

21, 28, and 35 to assess the amount of deposited ECM. Three histological stains were

used: Gomori trichrome, aggrecan immunohistochemistry, and safranin O.

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Gomori trichrome is used to quantify the total collagen secretion in the hydrogel as

collagen is a main structural component of cartilage representing more than 60% of its

dry weight (17). Gomori trichrome bind to the nuclei (dark purple) and collagen (green

or blue). Aggrecan immunohistochemistry is used to quantify the proteoglycan, which

represents more than 20% of the cartilage dry weight (408). Aggrecan, is specific

proteoglycan in cartilage tissue, and its crucial in a cartilage functioning (17, 409).

Safranin O used to identify the chondrocytes and the detection cartilage formation,

and it stains cell nuclei (dark red).

At each time-point, the crosslinked hydrogel systems were fixed with 10% (v/v)

formalin (Sigma Aldrich, UK) for 20 min at RT. After fixation the samples were

processed in a Tissue-Tek Vacuum Infiltration Processor (Bayer Diagnostics,

Newbury, UK). The processed tissue was embedded in Paraplast Plus paraffin medium

(Sigma Aldrich, UK) and sectioned at 5 µm before mounting on glass slides.

Gomori trichrome staining sections were de-waxed in xylene for 5 mins then hydrated

through descending grades of ethanol down to water. They were stained with Mayers

haematoxylin (Sigma Aldrich, UK) for 5 min, washed with water for 5 min and then

stained with Gomori (Sigma, UK) for 5 min. After rinsing with 0.2% acetic acid,

sections were blotted dry and dehydrated in ethanol before immersing in xylene and

finally cover-slipped.

Aggrecan immunohistochemistry staining sections were de-waxed in xylene for 5 min

then hydrated through descending grades of ethanol down to water. They were then

stained using Aggrecan (LSBio - LS-C359243) antibody on the Thermo Autostainer

480S Immunohistochemistry stainer. Briefly, antigen retrieval was performed using

citrate buffer pH 6. The primary antibody (Aggrecan, 1: 500) was added for 30 min

then the UltraVision™ Quanto Detection System HRP DAB (Thermo Scientific, UK)

was used. Sections were counterstained with Mayers Haematoxylin (Sigma Aldrich,

UK) then dehydrated in ethanol, immersed in xylene (Fisher Scientific, UK) and cover

slipped.

Safranin O staining sections were de-waxed in xylene for 5 min then hydrated through

descending grades of ethanol down to water. They were stained with Weigerts

haematoxylin (Fisher Scientific, UK) for 10 min, washed in water for 5 min and then

stained with 0.05% Fast Green (Sigma Aldrich, UK) for 5 min. After briefly rinsing

with 1% acetic sections were then stained with Safranin O (Sigma Aldrich, UK) for 5

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min, dehydrated in ethanol, immersed in xylene (Fisher Scientific, UK) and cover

slipped.

All sections were imaged on a Leica DM2700M microscope (Leica Biosystems, UK)

at x10 magnification and images were processed using ImageJ analysis software (407).

Glycosaminoglycan quantification

Glycosaminoglycan (GAGs) is a key marker of articular cartilage matrix production.

A dimethylmethylene blue (DMB) assay kit (Biocolor Ltd, UK) was used according

to the manufacturer’s instructions using a papain digestion method to quantify the

production of GAGs in the hydrogel samples (n=3, at each timepoint) and TCP control

at day 7, 14, 21, 28, and 35.

The papain extraction reagent was prepared by mixing 50 mL of prepared 0.2M

sodium phosphate buffer (pH 6.4) (Sigma Aldrich, UK) with 400 mg sodium acetate,

200 mg EDTA disodium salt, and 40 mg cysteine HCl (Sigma Aldrich, UK). Then, 5

mg of papain (Sigma, UK) was dissolved in 250 µL of deionised water before addition

to the extraction solution.

Cell culture media was removed from the samples (n=3), washed twice with DPBS,

and the samples digested for 3.5 h at 65°C by adding 1 mL of the papain digestion

solution to each well. After digestion, 50 µL of supernatant was added to 50 µL

deionised water, 1 mL DMB dye was added, and then mixed for 30 min on a

mechanical shaker. Samples were centrifuged at 12,000 rpm for 10 min and the

supernatants removed. The precipitated GAGs were dissociated by adding 0.5 mL

dissociation reagent, vortexed, and centrifuged at 12,000 rpm for 5 min.

The GAGs were measured by transferring the sample solution to a 96-well plate. A

microplate reader (Infinite 200, Tecan, Männedorf, Zurich, Switzerland) was used and

the absorbance measured at 656 nm.

Statistical analysis

The statistical analysis of cell proliferation, histology, and biochemical

characterisation (GAGs quantification and collagen content) were performed using

GraphPad Prism software version 8.0.1 (GraphPad Software, California, USA).

Significance levels were set at p < 0.05. Metabolic activity and proliferation were


test.

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Functionalisation confirmation and rheological characterisation

Methacrylate amino groups were successfully bonded to the alginate and gelatin

polymer backbones as confirmed by HNMR (Table 6.1 and Figure 6.3). Results show

that the reaction with methacrylate anhydride was successful and the polymers were

functionalized with methacrylate groups grafted onto the alginate and gelatin structure

(Figures 6.3 c and d). AlgMA shows methacrylation moieties at 6.2 ppm and 5.7 ppm,

corresponding to the protons on the alkene of the methacrylate, and at 1.7 ppm

corresponding to the methyl group on the methacrylate (CH3). GelMA shows new

peaks at 1.7-2.2 ppm, 5.69 ppm and 5.93 ppm corresponding to methylene protons

(H2C=C(CH3)), confirming the successful functionalisation with methacrylic

anhydride (Fig. 4d). The reaction between both polymers have resulted in the

appearance of new characteristic peaks of MA at 5.63 ppm and 6.09 ppm attributed to

the methylene group in the vinyl bond, and a peak at 1.82 ppm assigned to the methyl

group, which are not present in the non-modified polymers. The hybrid systems were

also successfully functionalized as shown in Figures 6.3 e, f, and g.

The degree of methacrylation was calculated and results are presented in Table 6.1.

The results show that all groups were reacted with methacrylic anhydride, and the

degree of carbon-carbon double bond substitution ranged between 55-79%. AlgMA

shows a high degree of methacrylation compared to other systems, which could be due

to using a high concentration of methacrylate anhydride (20%) reacted for long time

(24 h). GelMA shows the lowest substitution of methacrylate groups (55%). While in

the case of the hybrid systems, these values range between 65 and 71.5%.

Table 6.1: Degree of functionalization for all functionalized group polymers.

Sample (group) Degree of methacrylation

Alginate methacrylate(AlgMA) 79.0%

Gelatin methacrylate (GelMA) 55.0%

Hybrid IPN (Alg-Gel)MA 75:25 65.0%



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Figure 6.3: Representative 1H NMR spectra of a) non-modified alginate, b) non-

modified gelatin, c) AlgMA, d) GelMA, and the hybrid Alg-Gel(MA) systems e)

75:25, f) 50:50, and g) 25:75.

Rheological Characterization

The rheological behaviour of the polymers was investigated to determine their

viscoelastic properties which influences cell behaviour and their printability for the

development of a bioink. The viscoelastic behaviour of the polymers is directly related

to the molecular structure, percentage dilution, and functionalization processes. The

results show that all three systems (AlgMA, GelMA, and Alg-Gel(MA)) present shear-

thinning behaviour which is characterised by a decrease in the viscosity with the

increase of shear rate (Figure 6.4). This viscoelastic behaviour is important for

bioprinting cell laden constructs as the material will flow once a pressure is applied

and the lower pressure required to process the bioink through the printing nozzle will

improve cell viability (410). The cells also respond to the shear rate/time-dependent

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properties of their substrates since viscoelastic polymers encourage cell aggregation

and promote ECM formation (286, 411, 412).

Figure 6.4: Viscosity vs share rate profiles of AlgMA, GelMA, and Alg-Gel(MA)

(25:75, 50:50, and 75:25) samples.

The measurement of the amplitude sweep dependence of the G' and G'' allows

quantification of the linear viscoelastic region (LVER) that determines the stability of

the polymer solution. The viscoelastic behaviour is observed to be independent of

stress or strain values by monitoring the moduli (G' and G'') versus strain curve and

the critical strain point when G' crosses G''. The results show that the polymer

specimens behave as a rigid body, when low deformation is applied to the polymer

samples, G' and G'' are constant (G' is higher than G'') and the sample structure is

undisturbed (Figure 6.5 a, c and e). At higher strain rates both G' and G'' decrease and

the structure of the polymer solution is disturbed. Once the critical strain point is

identified, G' crosses G'', the LVER ends and the polymer behaves as a fluid after this

point. The material behaves as a rigid body at low strain rates but flows as a viscous

fluid at high strain rates.

After identifying the critical strain point, where G' crosses G'', frequency sweep

measurements were performed to determine the material frequency dependence over

a range of oscillation frequencies (0.1-100 Hz). The results present the effect of

frequency on the G' and G'' for all hydrogel systems (Figure 6.5 b, d, and f). For all

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hydrogels, G' is higher than G'', showing that in all cases there is a prevalence of the

elastic behaviour in comparison to the viscous one, which can also allow for high print

fidelity after the printing process. It can also be seen that the G' and G'' is not effected

by increasing the frequency, showing a plateau behaviour. The results also show that

the AlgMA system presents lower elasticity (G') and viscous properties (G'') than

GelMA and consequently lower printability characteristics. The gelatin and hybrid

systems present higher mechanical properties with higher G’ in the LVER which is

also independent of strain to higher values with a shift in the G' and G'' crossing point

(Figure s 6.5 c and e).

Figure 6.5: Oscillatory results for pre-polymerized systems a) Storage and loss

modulus as a function of strain for AlgMA systems; b) Storage and loss modulus as a

function of frequency for AlgMA systems; c) Storage and loss modulus as a function

of strain for GelMA systems; d) Storage and loss modulus as a function of strain for

GelMA systems; e) Storage and loss modulus as a function of strain for hybrid

systems; f) Storage and loss modulus as a function of frequency for hybrid systems.

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Hydrogel rheological characterisation was performed on the photocrosslinked

hydrogel discs by means of oscillation rheology test. Amplitude sweep results (G' and

G'' plotted as a function of shear strain) show that all hydrogel systems have the same

trend of G' being higher than G'' at low strain rates. however, the hydrogel crosslinking

bonds begin to break at higher strain rates, when the G' crosses G'' and both decrease

(Figure 6.6). The hybrid systems have higher mechanical elastic properties (G')

compared to the single material only systems (AlgMA and GelMA). The hybrid

hydrogel containing equal percentage of each polymer (50:50 AlgMA-GelMA)

presents the highest mechanical properties among all hybrid systems and withstands

deformation at higher strains compared to other hydrogel systems, indicating

improved stability (413).

Figure 6.6: Oscillatory results for polymerized systems a) Storage and loss modulus

as a function of strain for AlgMA systems; b) Storage and loss modulus as a function

of frequency for AlgMA systems; c) Storage and loss modulus as a function of strain

for GelMA systems; d) Storage and loss modulus as a function of strain for GelMA

systems; e) Storage and loss modulus as a function of strain for hybrid systems; f)

Storage and loss modulus as a function of frequency for hybrid systems.

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Structure morphology and porosity The hydrogels morphology was assessed using SEM (Figure 6.7). The

photocrosslinked hydrogel groups are characterised by a highly porous cellular

structure with an average pore size of ~250 µm for all systems. AlgMA has larger pore

size than GelMA, and different pore sizes were obtained for the hybrid systems as

shown in Table 6.2. This porous structure provides a route for nutrient supply, gas

exchange and metabolic activity as well as supporting and maintaining the phenotype

of the chondrocyte cells for articular cartilage applications. Accordingly, Mei Lien et

al. (414) reported that a pore size between 250-500 µm provides better ECM

production (62-64). All hydrogels provide highly porous structures (> 61%). The

AlgMA hydrogel have the highest porosity (90.8%) with an average pore size of 252.5

µm ± 20. The GelMA has the lowest porosity (61.0%) and an average pore size of

239.5 µm ± 30. The results show that the hybrid hydrogel samples present higher

porosity with higher amounts of AlgMA and that the porosity decreases by increasing

the incorporation of GelMA, which can be related to the inherent densities of the

biomaterials. These results are highly relevant for cartilage tissue engineering

applications since the highly porous structures produced, promote the secretion of

articular cartilage specific ECM and allows chondrocyte interactions and the creation

of a cell-cell network through the surrounding ECM.

Figure 6.7: Cross-section images of different hydrogel systems a) AlgMA, b)

GelMA, c) Alg-Gel(MA) 75:25, d) Alg-Gel(MA) 50:50, and e) Alg-Gel (MA) 25:75.

Scale bar: 200 µm.

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Table 6.2: Experimental density, porosity, and average pore size for the different

material systems

Sample Experimental

density (g/cm3) Porosity (%)

Average pore size

(µm)

AlgMA 0.149± 0.003 90.8%±0.21 252.5 ± 20

GelMA 0.382± 0.011 61.0%±1.16 239.5 ± 30

Alg-Gel(MA) 75:25 0.226± 0.001 84.5%±0.13 251 ± 24

Alg-Gel(MA) 50:50 0.263±0.001 79.9%±0.13 247 ± 30

Alg-Gel(MA) 25:75 0.340±0.005 70.2%±0.47 240 ± 15

Swelling and degradation kinetics and hydrogel

The swelling profile of all hydrogel groups is depicted in Figure 6.8. Results show that

the AlgMA hydrogel presents the highest swelling ratio, 50%, and GelMA hydrogels

present the lowest swelling, 18%. However, all hybrid systems show swelling ratios

ranging from 32%, 40%, and 17% for Alg-Gel(MA) 75:25, 50:50, and 25:75,

respectively. The swelling equilibrium was reached by day 3 for all hydrogels as

evidenced by a plateauing of the swelling curve. Moreover, results show that adding

AlgMA to the hybrid system improves the swelling ratio, indicating that the hydrogel

composition (ratio of biomaterials) controls the swelling properties.

The degradation profile of a hydrogel-based scaffold should match the development

and growth of new tissue with the ECM formation being able to maintain the

mechanical integrity of the structure (415). The degradation rates (percentage mass

loss) at day 22, for AlgMA was 23% , 27% for GelMA, and 17%, 20% and 47% for

the hybrid systems Alg-Gel(MA) 75:25, 50:50, and 25:75, respectively (Figure 6.8b).

However, these degradation rates can be tailored by increasing the crosslinking density

either by increasing the polymer solution/photoinitiator concentration and by

increasing the exposure time to the UV light source. The hydrogels containing gelatin

underwent higher degradation than alginate as gelatin undergoes faster rate of

hydrolytic degradation (416).

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Figure 6.8: Swelling and degradation profiles of all hydrogel groups. a) Swelling ratio

profile and b) degradation (mass loss profile).

Mechanical characterisation

The mechanical properties of the hydrogels were determined using unconfined

uniaxial compression testing (Figure 6.9). The hydrogel discs (cell-free constructs)

were measured on day 1 after incubation for 24 h at 37°C. The results show that

AlgMA and GelMA present an average compressive Young’s modulus of 28.1± 3.5

and 36.4± 6.0 kPa, respectively. The hybrid hydrogels exhibit an average of

compressive Young’s moduli 43.8 ± 3.1, 50.2 ± 4.4, and 66.2 ± 9.8 kPa for Alg-

Gel(MA) 25:75, 50:50, and 75:25 respectively. The Young’s modulus is highly

influenced by hydrogel composition, polymer concentration, photoinitiator

concentration, and exposure time as previously reported (417).

Figure 6.9: Mechanical properties of different photocrosslinked hydrogel, a) stress-

strain curve; b) Young modulus. Legend: A-alginate; G-gelatin; H-Hybrid system.

The mechanical compression test results are important to indicate the mechanical

stability of the hydrogels under potential physiological loading conditions. The

reported results within 20% strain are less than cartilage stiffness (400 to 600 kPa)

(418), however, it has been reported for other polymeric systems that similar

mechanical properties are appropriated to support the ECM formation and to maintain

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chondrogenic differentiation (419-421). Moreover, the proposed hybrid system must

also be anchored to the subchondral bone using a stiffer polymeric material.

Hydrogel encapsulated chondrocyte viability and proliferation

The viability of the encapsulated cells in the hydrogel was assessed using a live/dead

assay (Figure 6.10). At day 3, AlgMA shows a few live cells and the encapsulated

chondrocytes are dying due to lack of cell binding motifs such as RGDs which are

present in fibronectin, collagen, and laminin to which mammalian cells can bind (422,

423). However, GelMA and the hybrid systems present high cell viability. At later

timepoints (day 7 and 16) the alginate viability improves potentially due to the binding

of proteins from the cell culture media and ECM formation of the surviving cells.

These results show that cell encapsulation conditions were suitable. The photoinitiator

concentration, UV wavelength (365 nm), light intensity (8 mW/cm2), and exposure

time (8 min) had limited cytotoxic effect on the cells. Preliminary evaluation was

performed to assess the cell morphology. In the case of alginate encapsulated

chondrocyte cells, the surviving cells retained a round morphology due to high

crosslinking density of alginate and lack of binding motifs. However, the encapsulated

cells in gelatin system have shown to exhibit phenotype drift, development of

hypertrophic chondrocytes which could be due to low crosslinking density. Cells

encapsulated in the hybrid system had a cell morphology with a more spherical shape

for the entire duration of the in vitro studies compared to the AlgMA, GelMA and 2D

TCP systems. The chondrocytes cultured on 2D TCP have an elongated shape and

fibroblastic morphology.

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Figure 6.10: Cell viability of human chondrocytes encapsulated in AlgMA, GelMA,

Alg-Gel(MA) 50:50, and seeded on a TCP control at day 3, 7, and 16. Scale bar: 200

µm.

Cell proliferation and metabolic activity in the hydrogels was assessed using the

Alamar blue assay at day 1, 3, 5 ,7 and 14 (Figure 6.11). The results show that the

cellular metabolic activity increases for all hydrogels up to day 3 before decreasing

with the AlgMA hydrogel exhibiting the largest decrease in metabolic activity. The

large decrease in AlgMA may be attributed to the lack of cell binding sites in alginate

thus poor viability, as previously shown, and the degradation of the hydrogel

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preventing cell maintenance and proliferation. The GelMA and hybrid hydrogel show

higher cell metabolic activity up to day 14 compared to AlgMA, potentially due to the

presence of cell binding sites in gelatin allowing cell attachment. Furthermore, the

reduction and plateauing of chondrocyte metabolic activity after day 5 may be due to

the low metabolic and proliferative activity of chondrocytes when present in an

environment that mimics the native ECM environment of articular cartilage tissue.

Figure 6.11: Fluorescence intensity indicating the metabolic activity of human

Chondrocytes in AlgMA, GelMA, Alg-Gel (MA) 50:50 and TCP control at days 1, 3,

5, 7 and 14.

Histological assessment of hydrogel ECM formation

Histological evaluations were performed to visualise and semi-quantify the

cartilaginous specific ECM formation in the photocrosslinked hydrogels at day 14, 21,

28, and 35 (Figure 6.12). The Gomori trichrome staining results indicate that in all

experimental groups, most cells were positive, and collagen was produced, however,

the hybrid system presented higher collagen content (Figure 6.12 and 6.13). Semi-

quantitative analysis shows clear difference in the amount of collagen detected. The

amount of collagen in all hydrogels increased until day 21. At day 28 and 35, the

amounts of collagen started to decrease. A possible interpretation is due to the

degradation of the hydrogel, since at least a minimum of 20% of the hydrogel has

D1 D3 D5 D7 D14

0

2000

4000

6000

Time (days)

GelMA

Hybrid

AlgMA

Flu

ore

sc

en

ce

2D control

**

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degraded by this timepoint. Thus, cells are migrating to attach to the TCP and amount

of collagen production is reduced and being removed during standard cell culture

procedures. This shows the importance of developing a hydrogel which maintains

structural stability long enough for sufficient ECM tissue formation to support the

cells.

The expression of aggrecan is investigated through immunohistological staining. The

results show that aggrecan was deposited within all the hydrogel systems. Aggrecan

was distributed throughout the hydrogels and was highest in the Alg-Gel(MA) hybrid

system at all timepoints whilst lowest in the AlgMA (Figure 6.12 and 6.13). The peak

aggrecan content is at day 28 before decreasing potentially due to the degradation of

the hydrogels.

The safranin-O staining is present in all hydrogels with the most proteoglycans

detected in the hybrid system with considerably stronger staining than AlgMA and

GelMA.

The histological staining and semi-quantification showed that the composition,

content, and distribution of the ECM varied among the hydrogels. The AlgMA had

the poorest capacity to support cartilage specific ECM formation, as shown by the

weak intensity of staining in the samples, due to the poor cell binding capacity The

GelMA and the hybrid hydrogel enabled the deposition of ECM by the encapsulated

chondrocytes and had the highest qualitative and semi-quantitative staining. This

indicates that the hybrid hydrogel is supportive of maintaining a chondrocyte

phenotype. The hybrid Alg-Gel(MA) is stiffer than the single component hydrogels

support the importance of stiffness as factor in hydrogel design as a stiffer hydrogel

system provides better chondrocyte stimulation to maintain chondrogenic phenotype

after encapsulation in 3D, these results align with Li et al. (424) demonstration of

different chondrocyte phenotypes based on the stiffness of a gelatin hydrogel.

148

(a) Gomori

Figure 6.12: Histological analysis. a) Gomori; b) Aggrecan; c) Safranin-O production

at days 14, 21, 28 and 35 in Alginate methacrylate hydrogels, Gelatin methacrylate

hydrogels and hybrid system (50:50 Alg-Gel)MA.

149

(b) Aggrecan

Figure 6.13 (Cont.): Histological analysis. a) Gomori; b) Aggrecan; c) Safranin-O

production at days 14, 21, 28 and 35 in Alginate methacrylate hydrogels, Gelatin

methacrylate hydrogels and hybrid system (50:50 Alg-Gel)MA.

150

(c) Safranin-O

Figure 6.14 (Cont.): Histological analysis. a) Gomori; b) Aggrecan; c) Safranin-O

production at days 14, 21, 28 and 35 in Alginate methacrylate hydrogels, Gelatin

methacrylate hydrogels and hybrid system (50:50 Alg-Gel)MA.

Figure 6.15: Semi-quantification of the relative area of histological staining of the

AlgMA, GelMA, and Alg-Gel(MA) hydrogels at day 14, 21, 28, and 35 for a) Gomori

trichrome and b) aggrecan.

151

GAGs quantification

GAGs are a major component of articular cartilage and an indicator of cartilage

formation. The secretion of GAGs was detected in the AlgMA, GelMA, and hybrid

Alg-Gel(MA) hydrogels at day 7, 14, 21, 28, and 35 (Figure 6.14). The results show a

clear difference in GAGs production in the three different hydrogel systems. The

AlgMA hydrogels show the lowest GAG production due to poor cell viability and cell

attachment. However, the GelMA and hybrid hydrogels produced larger quantities of

GAGs with the hybrid hydrogel having the largest production of GAGs by day 35.

The 2D TCP control produced similar amounts to the AlgMA hydrogel even though a

larger number of cells are present in the TCP control. This discrepancy can be

attributed to the elongated chondrocyte cell morphology present in 2D cell culture and

the dedifferentiation and phenotypic changes that occur as the chondrocytes proliferate

in 2D cell culture (425). GAGs are crucial ECM factors in cartilage tissue due to their

ability to strongly bind water, create osmotic pressure, and enable attachment of

proteins. The high quantity of GAGs produced in the hybrid system indicates that the

hydrogel can support the maintenance of the chondrocyte phenotype and secretion of

a cartilage specific marker.

Figure 6.16: Assessment of GAG secretion in the photocrosslinked AlgMA, GelMA,

and Alg-GelMA hydrogels at day 7, 14, 21, 28, and 35.

D7 D14 D21 D28 D35

0.0

0.5

1.0

1.5

2.0

Time (days)

Ab

so

rban

ce (

656 n

m)

Control

Hybrid

GelMA

AlgAM

✱✱

✱✱✱

✱✱✱✱

✱✱✱✱

✱✱✱✱

✱✱✱ ✱✱✱✱

✱✱✱✱

✱

✱✱

152

Conclusions

The AlgMA, GelMA, and hybrid Alg-Gel(MA) solutions and hydrogels were

extensively characterised aiming at understanding their suitability to be used as bioink

materials for cartilage applications. All polymers were successfully functionalised

with MA as observed by HNMR. Photopolymerisation was successfully used to

polymerised functionalized hydrogels. Porous structures with appropriate pore sizes

for cartilage applications were obtained. The hybrid system demonstrates that the

combination of both AlgMA and GelMA improves the viscoelastic behaviour of pre-

polymerised materials (improved printability) and mechanical properties of

polymerised hydrogels.

The hybrid system also presents high cell viability, proliferation, and production of

ECM proteins, and GAGs. The results show that the developed hybrid hydrogel is a

suitable bioink material for cartilage applications. It is able to support chondrocytes

proliferation, cell-cell interaction and ECM formation. Moreover, the properties of this

system can be easily tailored by changing the ration between AlgMA and GelMA.

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Chapter 7 Hybrid PCL/Hydrogel scaffold

fabrication and in-process plasma

treatment using PABS†

† This Chapter based on the paper: Fengiuan Liu, Hussein Mishbak, Paulo Bartolo, “Hybrid

polycaprolactone/hydrogel scaffold fabrication and in-process plasma treatment using PABS”,

International Journal of Bioprinting, 5(1): 174, 2019. http://dx/doi.org/10.18063/ijb.v5i1.174. A

preliminary version of this manuscript was also presented at the International Conference on Progress

in Additive Manufacturing (Singapore, 2018) and was awarded as the best paper.

http://dx/doi.org/10.18063/ijb.v5i1.174

154

The development of a new tissue engineering approach to enable the neocartilage

formation fully anchored to the underlying subchondral bone is critical for tissue

engineering. In Chapters III to VI, two different hydrogel materials (alginate and

gelatin) and a hybrid system based on both alginate and gelatin were characterised

and their potential for cartilage bioinks explored. These different polymeric systems

will enable to create a cartilage construct replicating the zonal structure of cartilage.

However, it is also important to guarantee a good interface between the subchondral

bone and the cell laden construct to guarantee the stable formation of neocartilage

tissue and to enable vertical integration. This interface must be created using both

hard and soft materials.

For bone tissue engineering scaffolds, a wide range of polymeric, ceramic and

polymer-ceramic materials have been investigated (426). Among these materials

polycaprolactone (PCL) has been one of the most popular material. However, due to

its hydrophobicity produced scaffolds are limited in terms of cell-seeding and

proliferation efficiency (427). Moreover, non-uniform cell distribution along the

scaffolds with limited cell attachment in the core region is a common problem. The

absence of commercially available additive manufacturing systems able to produce

multi-type-material and gradient scaffolds makes this problem difficult to address.

This Chapter investigates the use of a novel Plasma-Assisted Bio-extrusion System

(PABS), to produce multi-material scaffolds for the interface between the subchondral

bone and cartilage consisting in a synthetic biopolymer (PCL) and hybrid natural

hydrogel system (previously described in Chapter VII). The PABS system is also used

for the scaffold surface modification (N2 plasma modification).

Introduction

Tissue engineering is promising for organ replacement which minimizes the side

effects of organ transplantation (428, 429). Biomanufacturing is the major strategy of

tissue engineering aiming at the development of biological substitutes that restore,

maintain, or improve tissue function, and it requires the combined use of additive

manufacturing (AM), biocompatible and biodegradable materials, cells and

biomolecular signals (304).

The scaffolds-based strategies (Figure 7.1) for tissue engineering has been most

commonly used, depending strongly on materials and manufacturing processes. For

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materials, five types of biomaterials have been used: acellular tissue matrices,

synthetic polymers, natural polymers, ceramics and polymer/ceramic composites

(430-434).The most commonly used biomaterial for producing scaffolds are synthetic

polymers, such as polycaprolactone (PCL). Polymeric scaffolds play a pivotal role in

tissue engineering through cell seeding, proliferation, and new tissue formation in

three dimensions, showing great promise in the research of engineering a variety of

tissues. Moreover, Scaffolds made from collagen are being rapidly replaced with

ultraporous scaffolds from biodegradable polymers.

Figure 7.1: scaffold-based approach for tissue engineering.

Biodegradable polymers are attractive candidates for scaffolding materials because

they degrade as the new tissues are formed, eventually leaving nothing foreign to the

body. The major challenges in scaffold manufacturing lies in the design and

fabrication of customizable biodegradable constructs with properties that promote cell

adhesion and cell porosity, along with sufficient mechanical properties that match the

host tissue, with predictable degradation rate and biocompatibility (435, 436).

For the fabrication methods, AM techniques have been commonly applied in scaffold

fabrication due to the superior ability in controlling pore size, pore shape and pore

distribution, and thus creating interconnected porous structures (304, 437).When

combined with clinical imaging data, these fabrication techniques can be used to

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produce constructs that are customized to the shape of the defect or injury (438). In

terms of tissue and organ manufacturing, the additive nature ensures minimal waste of

scarce and expensive building material, namely cells, growth factors and biomaterials

(439-441). Among the AM techniques, Material Extrusion has been mostly applied in

the bioengineering field due to the flexibility in material selection based on the use of

pneumatic (442, 443), piston (444) and screw-assisted (445-447) extrusion systems

enabling a wider range of materials to be applied. Some processes operate at room

temperature, thus allowing for cell encapsulation and biomolecule incorporation

without significantly affecting viability. However, cell-seeding and proliferation

efficiency is currently a big challenge due to the following limitations (448-451):

• Most AM techniques are limited to single-material fabrication, which is difficult

to provide appropriate environment for cells due to the inadequate chemical,

physical and biological cues provided during AM processes.

• Additionally, non-uniform cell distribution, especially rare cell adhesion in the

core region of scaffolds, is often caused by the tortuosity of the constructs.

• Moreover, the synthetic biopolymers, most commonly used, are hydrophobic, and

the cell colonization.

Different strategies have been explored to solve the above problems. Multi-material

have been developed and utilized to produce multiple-material scaffolds (442, 452).

However, most of these systems can only form one type of biomaterials, either soft

hydrogels containing cells or bio-signals in the scale of KPa, or rigid biopolymers and

composites in the scale of MPa, which fails in the mimicry of natural tissues.

Additionally, Additionally, low temperature plasma modification is capable of

improving the hydrophilicity of biopolymers by inducing certain functional groups on

the surface to change the chemistry, wettability and energy without altering the bulk

properties (453, 454). However, most plasma treatment can be conducted after

scaffolds printed and the penetration depth is limited, which results in non-uniform

cell distribution along the scaffold.

A novel plasma-assisted bioprinting system (PABS) has been developed in the

University of Manchester, allowing processing soft-hard biomaterial integration and

plasma surface modification layer by layer during the fabrication process in the same

chamber. This paper utilized the plasma-assisted bioextrusion system (PABS) to

produce PCL/Hydrogel hybrid scaffolds and plasma fully treated scaffolds. The

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hydrogel is assesssed with the preparation procees, functionalizaiton process and

rhology properteis, while the PCL fully treated scaffolds are both morphlogically and

bilogically assessed.

Materials and Methods

Polycaprolactone (PCL)

PCL (CAPATM6500, Mw = 50,000 g/mol), purchased from Perstorp (Cheshire, UK)

in the form of 3 mm pellets, was used to produce the scaffolds. PCL is an easy-to-

process semi-crystalline polymer with a density of 1.1 g/cm3, a melting temperature

between 58-60 °C, and a glass transition temperature of - 60 °C.

Hybrid hydrogel) methacrylate anhydride (Alg-Gel)MA preparation

Hybrid hydrogel system platform were considered in this paper. hydrogel which made

of mixing alginate methacrylate and gelatin methacrylate at 50:50% v/v.

Functionalization process for both polymer (alginate and gelatin) is necessary to

introduce the carbon-carbon double bond into the polymer chains that eventually

convert the polymer into photopolymerized polymer. According to published

protocols (455), alginate was functionalized with methacrylate groups with minor

modifications to be photopolymerized. The 2% wt powder alginate (Sigma-Aldrich,

UK) was dissolved in Dulbecco's Phosphate Buffered Saline (DPBS) (Sigma-Aldrich,

UK), and then mixed with methacrylate anhydride (MA) (Sigma-Aldrich, UK) at 23%

v/v of alginate solution under vigorous stirring. The pH of the solution was kept around

7.4 during the reaction time by the addition of 5M NaOH. The reaction time was 24

hours. After the chemical modification, the polymer solution was precipitated with

cold ethanol, dried in an oven overnight at 45 °C and purified through dialysis for 6

days. The solution was then frozen at -80 ºC and recovered by lyophilization.

Gelatin was functionalized by dissolving gelatin bovine skin type B (Sigma-Aldrich,

UK) at a concentration of 12.5%, in Dulbecco's Phosphate Buffered Saline (DPBS)

(Sigma-Aldrich, UK) at a temperature of 45ºC. After gelatin dissolution, MA (10x

molar excess) was added under vigorous. The pH of the solution was kept around 7.0-

8.0 during the reaction. The reaction time was 6 hours. The solution was purified

through dialysis for 7 days, frozen at -80C° and the polymer recovered by

lyophilization.

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Plasma-assisted Bioextrusion System set up

Scaffolds were produces using a hybrid additive manufacturing system called PABS,

being developed at University of Manchester. PABS (Figure 7.2 a) comprises two

main units, a multi-extrusion unit and a three-inlet plasma modification unit (Figure

7.2b). The multi-extrusion unit consists of two pressure-assisted extruders and one

screw-assisted extruder, allowing operating a range of biomaterials, such as synthetic

biopolymers, hydrogels and biopolymer/ceramic composites. The extrusion unit has

four movements: one rotational movement (C1) for selecting the required extrusion

heads, a second rotational movement (C2) for driving and controlling the screw

rotational speed of the screw-assisted extruder, and two linear movements (X and Y)

in the X-Y plane. All these four movements are controlled by stepper motors and CNC

drives. The build platform moves in the z-direction and constitutes the sixth

controllable axis. The build platform was fabricated using 7076 aluminium plate (250

mm x 200 mm x 7.5 mm) and is attached to the z-axis rail guide with an L-shape

support underneath.

The plasma modification unit is mounted on the X-Y platform which is co-planar with

the extrusion platform. Both platforms share common cylindrical guide rails in the X-

direction. The Y-direction movement (V axis) for the plasma modification unit is

independently driven by a stepper motor parallel to the Y axis of the extrusion unit. A

quartz capillary (outside diameter of 7 mm; inner diameter r of 5 mm; length of 70

mm) with three gas inlets serves as the reaction jet. A tungsten rod (inner diameter of

2 mm) and one copper film (10 mm of width) wrapped around the quartz tube serve

as the high-voltage and ground electrodes, respectively. The electrode is connected to

a high-voltage DC power supply (applied voltage of 10 kV; frequency of 50 kHz). The

plasma is generated from the top central electrode, expanding to the surrounding air

inside and outside the nozzle.

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Figure 7.2: a) Set-up of plasma-assisted bio-extrusion system; b) system diagram:

multi-extrusion system, plasma modification system and control system.

The control system consists of a motion control system, temperature control system

and gas supply control system. The motion control system utilizes a Geo Brick LV

motion controller (Delta Tau Data Systems, Inc) to manipulate all motors. The built-

in software allows for complete machine logic control, including G code execution.

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The temperature control system consists of four digital temperature controllers (P.I.D.

Omron E5CN), which are used to precisely control the temperature in the extruder

heating chamber with polyimide thermos-foil heating elements (Omega

Online/Kapton Heaters), and thermocouples (Omega). The gas supply control system

includes an air compressor which works as the main gas supply, regulators and gauges

(SMC, UK), which enables pressure on/off control and manually adjustment. Power

supply (230V ac, PHOENIX CONTACT) offers 24v output voltage and 40A current

for all the control system. Safety considerations were taken into account regarding the

electrical parts, including fuses, main circuit breakers (MCBs), main filter, push

buttons and estops (all supplied by OneCall Electronics Company, UK), and circuits.

The control software was developed in MATLAB as previously reported, and G code

files was generated containing all the instructions for the fabrication process (456).

This configuration enables multi-material dispensing and plasma modification in a

sequential mode. The system is able to achieve a maximum linear velocity of 20 mm/s

and resolution of 0.05 mm.

Scaffolds printing strategies

Scaffolds with a cross-section of 10x10 mm and a height of 3 mm were fabricated

using the single laydown pattern of 0/90°.and a filament distance of 1 mm with pore

size of 1 mm (slice thickness of 0.5 mm; heating temperature of 90°C; extrusion screw

rotational speed of 15 rpm; and nozzle tip size of 0.5 mm.). The strategies for printing

hybrid scaffolds and plasma treated scaffolds are as follows (Figure 7.3):

• Hybrid PCL/hydrogel scaffolds fabrication (Figure 7.3a): After the scaffold being

printed with screw-assisted extruder, the extruder selection unit rotated with 120°,

and then the hydrogel solution was printed in the pores using the pressure-assisted

extruder.

• Full layer treated scaffolds fabrication (Figure 7.3b): The N2 plasma modification

was performed after one-layer PCL was deposited, and the process was repeated

until the last layer of PCL was complete. The treatment was conducted at a

pressure of 0.689 bar and a flow rate of 5 L/mm. The deposition speed of the

plasma jet was 3 mm/s and each layer was subjected to the plasma treatment for

one minute. The distance from the bottom of the jet to the surface of the PCL

filaments was 10 mm.

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Figure 7.3: a) Sequence of operations to fabricate hybrid PCL/hydrogel scaffolds; b)

Sequence of operations to fabricate full-layer plasma treated scaffolds.

Nuclear Magnetic Resonance characterisation

Hybrid hydrogel were characterized by means of 1HNMR to confirm the

functionalization process. The chemical structure of functionalized hybrid hydrogel

system was assessed through Nuclear Magnetic Resonance Spectroscopy (1HNMR),

using the B400 (Bruker Avance III, USA) 400 MHz. Polymeric materials were

dissolved in Deuterium oxide (D2O), transferred to NMR tubes and the spectra

acquired with 128 scans.

Rheological characterisation

To assessing the mechanical properties of the hybrid hydrogel quantitatively, a

deformation rheology tests are performed on polymeric hydrogels discs. The

rheological analyses were carried out using the DHR2 Rheometer (TA Instrument,

USA). Small amplitude oscillatory shear measurements tests were considered, the

shear storage modulus, G′, loss modulus, G″ were quantified. It is assumed that the

material flows by applying a stress being the response measured. A controlled stress

was applied, and the resulting movement measured. Small strain oscillations have

been used to measure viscoelastic properties without destroying the sample structure

the oscillatory test provides a mechanical spectrum for the material.

Morphology Characterization

Scanning electron microscopy (SEM) was used to assess the morphology and surface

characteristic of printed scaffolds. The scaffolds were gold/palladium coated using a

Q150T turbo-pumped sputter coater (Quorum technologies, UK) and imaged at 10 kV

(Hitachi S3000N, Japan). These images were then analysed using ImageJ software.

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Wettability measurement

Water Contact Angle (WCA) measurements on the flat surfaces of untreated and

plasma-treated PCL scaffolds were carried out with a commercial KSV CAM 200

system (KSV Instruments, Finland). The system is equipped with a CCD video camera

and a micrometric liquid dispenser to drop 2 µL of distilled water on the surface of the

scaffold. The measurements of the contact angles are automatically calculated with

the instrument software.

Biological tests

In vitro biological assessments were conducted with human adipose-derived stem cells

(hADSCs) (STEMPRO, Invitrogen, Waltham, MA, USA). Before cell seeding,

scaffolds were sterilized by soaking in 70% ethanol for 2 hours. After sterilization,

samples were rinsed twice in phosphate buffered saline (PBS) (Gibco, ThermoFisher

Scientific, Waltham, MA, USA), transferred to 24-well plates and air-dried for 24

hours at room temperature 50,000 cells were seeded on each sample, including

plasma-treated and untreated scaffolds.

Cell viability/proliferation behavior and the percentage of cells attached to the

scaffolds (cell-seeding efficiency) were assessed through Alamar Blue assay (also

termed the Resazurin assay, reagents from Sigma-Aldrich, UK). Cell

viability/proliferation was measured at 1, 3, 7 and 14 days after cell seeding. For each

measurement, cell-seeded scaffolds were transferred to a new 24-well plate and 0.7 ml

of Alamar Blue solution was added to each well, the plate was incubated for 4 hours

under standard condition (37 ºC, 5% CO2 and 95% humidity). After incubation, 150

µL of each sample solution was transferred to a 96-well plate and the fluorescence

intensity measured at 540 nm excitation wavelength and 590 nm emission wavelength

with a spectrophotometer (Sunrise, Tecan, Männedorf, Zurich, Switzerland).

Cell attachment and distribution are assessed using laser confocal microscopy, with

cell nuclei stained. At day one of cell culture and after 14 days, scaffolds were removed

from 24-well cell culture plate, rinsed twice in phosphate-buffered saline (PBS)

(Gibco, ThermoFisher Scientific, Waltham, MA, USA), fixed with 10% neutral

buffered formalin (Sigma-Aldrich, Dorset, UK) for 30 minutes at room temperature

After fixation, samples were rinsed twice with PBS for the removal of formalin, then

permeabilised with 0.1 % Triton-X100 (Sigma-Aldrich, Dorset, UK) in PBS at room

temperature for 10 minutes, rinse twice for the removal of Triton-X100. Cell nuclei

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were stained blue by soaking scaffolds in a PBS solution containing 4′,6-Diamidine-

2′-phenylindole dihydrochloride (DAPI) (Sigma-Aldrich, Dorset, UK) at the

manufacturer recommended concentration. Samples were left in the staining solution

for 10 min prior to removal, rinsed twice thoroughly with PBS. Confocal images were

obtained using a 3D rendering mode on a Leica TCS SP5 (Leica, Milton Keynes, UK)

confocal microscope, and cell colonization was quantified using a standard z-stack

method. All images were taken at the centre area of the scaffolds and the experiments

were performed three times in triplicate.

Data analysis

All data are represented as mean ± standard deviation. Biological results were


test using GraphPad Prism software. Significance levels were set at p < 0.05.

Results

Hybrid Scaffold Fabrication

The hybrid scaffold, consisting of PCL and hydrogel solution, was successfully

fabricated using PABS. Figure 7.4a shows the image of the hybrid scaffold, where the

porous structures hold the hydrogel solution without any deformation. The position of

the deposited hydrogel was precisely controlled.

Full-layer treated PCL scaffold

Figure 7.4b shows that the printed full-layer N2 plasma treated scaffolds has a well-

bonded interconnected structure with uniform pore distribution and pore size in the

range of ~500 µm (Figure 7.4c). The N2 plasma treatment was conducted at a pressure

of 0.689 bar and a flow rate of 5 L/mm. The deposition speed of the plasma jet was 3

mm/s and each layer was subjected to the plasma treatment for one minute. The

distance from the bottom of the jet to the surface of the PCL filaments was 10 mm.

Figure 7.4d presents the filament surface after N2 plasma modification, where lines in

the direction of plasma movement can be observed making the surface roughness

increased.

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Figure 7.4: (a) Hybrid polycaprolactone (PCL)/hydrogel scaffold; (b) Scanning

electron microscopy (SEM) image of top view of PCL/hydrogel scaffold; (c) SEM

image of side view of PCL/ hydrogel scaffold; (d) photo of a full-layer N2 plasma-

treated PCL scaffold; (E) SEM image of top view of full-layer N2 plasma treated PCL

scaffold; (f) SEM image of filament surface of full-layer N2 plasma-treated PCL

scaffold.

Wettability assessment of full-layer plasma modified scaffold

WCA measurements were performed on untreated and fully treated plasma PCL

scaffold surfaces to determine the effect of plasma modification on the surface

wettability. Table 7.1 highlights the WCA results at different time points after the

droplet was dropped on the surface of the scaffolds. The results show that in the case

of an untreated PCL scaffold, there are no significant changes in the WCA values with

time with values varying between 83.2±2.0° and 80.9±2.7°. For treated scaffolds, the

WCA value, at 0 s, was lower (63.0±3.1°), leading to a fully wetting value of 26.7±0.9

° at 0.5 s. At 3 s, the droplet was fully absorbed.

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Figure 7.5: a) Temporal variation of fluorescence intensity of cell-seeded PCL

scaffolds with and without N2 plasma treatment; b) Confocal microscope images of

untreated and full-layer treated of cell seeded scaffolds, 1 day and 14 days after cell

culture. Scale bar 250 µm.

Table 7.1: Temporal variation of water contact angles for treated and untreated

scaffolds.

Time PCL scaffolds N2 plasma fully treated

0s 83.2±2.0 ° 63.0±3.1 °

0.5s 82.9±1.2 ° 26.7±0.9 °

3s 80.9±2.7 ° Fully absorbed

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Biological assessment of full-layer plasma modified scaffold

The adhesion and proliferation of hADSCs cells on plasma full-layer modified PCL

scaffolds was studied and compared with untreated ones. The biological

characterization was assessed using Alamar Blue Assay. The fluorescence intensity of

cell seeded scaffolds measured at four different culture time points (Day 1, 3, 7 and

14) is shown in Figure 7.5a. Higher fluorescence intensity corresponds to more

metabolically active cells. As observed, cell proliferation increases with time in all

types of scaffolds, suggesting that they are suitable structures for cell attachment and

proliferation. However, a fast proliferation rate is observed in the case of plasma

treated scaffolds. The different performance is statistically significant after day 3.

Confocal microscopy images (Figure 7.5b) present the cell attachment and distribution

after cell seeding (day 1) and proliferation (day 14). It can be observed that plasma

treated scaffolds presented higher numbers of cells than untreated scaffolds.

Additionally, it is also possible to observe that plasma surface treated scaffolds

presented best cell attachment and dispersion.

Discussion

The printed hybrid PCL scaffolds filled with hydrogel present the printability of

PBAS, enabling the soft-hard material integration. The photo and SEM images of the

whole printed structure indicate that the plasma modification process does not affect

the physical appearance of the structure and potentially no effect on the mechanical

performances (457). Moreover, SEM image of the filament surface confirms the

increased surface roughness is due to the etching process (458), which results in the

stripping off the topmost layer of the polymer filament. From this study, the increase

surface roughness is also resulted from the linear scratches attributed to the gas flow

by the plasma jetting directly after the PCL deposition when the material is still in the

molten status. As the printed material is not totally cooled down, when the plasma

modification occurs, the gas flow effect is stronger, significantly influencing the

surface topography. Moreover, the increase of surface roughness is beneficial for cell

colonization in the scaffolds (459).

The wettability results reveal that the hydrophilicity of the surface is dramatically

improved due to the ionisable groups introduced on the surface by the N2 plasma,

which enhances the hydrogel bonding with water. Compared with the results published

in (454,460), the absorption speed is dramatically increased, due to chemical

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heterogeneity on the surface of each layer because of the plasma modification layer

upon layer. However, the effect of plasma modification can last within a certain period

of time depending on the density of chemical regents deposited on the surface, and the

effect also decreases with time since the structural rearrangement of polymer chains

resulting in the decreased surface energy (461).

In addition, it is important to control hydrogels printability window (injection,

deposition and extrusion) during printing process to mimic the native organ structures.

To achieve precise printable properties, hydrogels have been designed to achieve

adjustable viscoelastic behaviour, and the SEM images confirm the liquid-like-gel

structure before printing and quick gelation kinetics to produce structure with good

fidelity. Hybrid hydrogel system presents an adjustable viscoelastic behaviour that

could be changed/ tailored with temperature, shear thinning and polymer

concentrations. The rheological assessment result confirms that the hybrid hydrogel

system has clear linear viscoelastic region (LVE) which characterize by both storage

moduli G′ and loss moduli G″ are independent of shear strain and the system is solid-

like-gel (G′ ≫ G″). These characteristics make it easy to inject hybrid hydrogel

system (Alg-Gel)MA.

Consistent with the other studies (462, 463), a significant increase in the biological

behavior is observed after plasma modification. Since the hydrophilicity has been

improved with the plasma modification process, the environment is changed to be

more suitable for cell colonization and proliferation. Also, with the full-layer plasma

modification capability, the plasma active species can reach the walls of the pores on

each layer of the scaffold, leading to the uniform cell distribution along the scaffold.

This will enable higher rate of tissue formation in the clinical research.

Conclusions

This paper presents a novel additive manufacturing system comprising a multi-

material printing unit and a plasma jet unit. To assess the system, hybrid PCL/hydrogel

scaffolds and full-layer plasma treated PCL scaffolds were produced. The effect of

plasma treatment on PCL samples was examined. WCA results confirmed that the

hydrophilic character of the PCL samples increased due to the nitrogen groups

introduced by the plasma jetting on the scaffold filaments. It was also possible to

observe that the plasma treatment positively influences cell attachment and

proliferation. Applications that may benefit from this technology include hybrid tissue,

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which has compositional variations depending on the region or organ-like structure

that require continuous vascular network to facilitate nutrient diffusion.

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Chapter 8

Conclusions and future

work

170

Conclusions

Cartilage is an avascular tissue with limited regenerative capabilities. Its structure, role

and main diseases and clinical therapies were detailed in Chapter II (research objective

1). Despite different treatments have been proposed and being clinically used they still

present several limitations such as high cost, poor clinical outcomes, and limited

temporal efficacy. In order to address these limitations cartilage tissue engineering

approaches based on the use of biocompatible and biodegradable materials and

additive manufacturing technologies to produce scaffolds for cartilage repair or cell

laden constructs based on the use of bioinks have been proposed. The scaffold-based

approach is based on the use of 3D structures that will provide the substrate for cell

adhesion, proliferation and differentiation. In this approach scaffolds can be used as

acellular structures and implanted into the defect, recruiting cells from the surrounding

tissue; or seeded with cells and pre-cultured in a bioreactor prior implantation. The

success of this approach has been limited and the scaffolds being proposed so far are

not able to recapitulate the zonal structure of cartilage. Cell laden constructs seems to

be the ideal strategy not only because it allows to print a cartilage-like construct

encapsulating relevant cartilage cells and materials with a structure similar to cartilage

but also because it allows the development of future in situ printing approaches.

Bioprinting approaches to produce cell laden constructs for cartilage repair have been

proposed. However, few limitations still exist related to materials (e.g. tailoring

viscoelasticity), printing process (e.g. shear stresses and shape fidelity) and risk of

contamination during the bioprinting process. The development of novel constructs

for cartilage applications is critical and requires the:

• Design of suitable materials, able to be printed and presenting suitable

mechanical, swelling, degradability and morphology and able to sustain cell

growth and ECM formation;

• Design of a multi-layer construct able to replicate the zonal structure of

cartilage;

• Selection of the best additive manufacturing technique (e.g. crosslink

mechanism).

As discussed in Chapter 2, natural-based hydrogels are ideal candidates for cartilage

applications and photocrosslinking reactions a suitable method to produce cell laden

constructs. However, only two photocurable hydrogels are commercially available

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(ChonDuxTM and GelrinC). This an area that requires more research not only in terms

of design the hydrogels (e.g. level of C=C in the pre-polymer chain) but also on

investigating the effect of the curing process on both printability and cell survival.

However, this is not a simple task as high levels of C=C bonds increase the

crosslinking density, improving mechanical properties and shape fidelity during the

fabrication process but decreases biocompatibility as the free space available to

support cell growth is reduced. Similarly, high levels of photoinitiator increase the

curing kinetics and crosslinking density but might also induces some toxicity. Light

intensity and irradiation time must also be optimised in order to avoid any negative

effects on cells. Some of these correlations were investigated in this research work and

the results presented in Chapters III to VI.

Based on the aims and research objectives presented in Chapter I, the main novelty

of this research is the development of a novel multi-material cell laden construct

for cartilage based on two different polymer materials (alginate and gelatin) and

a hybrid system based on the combination of both alginate and gelatin. Single type

materials are considered for both the superficial zone and the deeper region of

cartilage, while the hybrid system is considered for the transitional zone. Both alginate

and gelatin have been explored before for different tissue engineering applications but

their combination to create a zonal construct such the one proposed by this

dissertation was never investigated before. Each material was investigated in detail

as a single material and the results presented in Chapters III and IV for alginate and

Chapter V for gelatin. Each material was successfully functionalized with

methacrylate anhydride as observed through NMR and it was possible to control the

functionalization process by changing the polymer concentration, methacrylate

anhydride concentration and reaction time (research objective 2). Key conclusions

emerging from this part of the work are:

• By increasing the polymer concentration, keeping the reaction time and the

methacrylate anhydride concentration constant, the degree of modification

(level of unsaturation) decreases;

• By increasing the methacrylate anhydride concentration, keeping polymer

concentration and reaction time constant, the degree of modification

decreases;

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• By increasing the reaction time, keeping both methacrylate anhydride and

polymer concentration constant, the degree of modification increases.

The effect of the functionalization process on the rheological, printability, mechanical,

morphological, swelling and degradation properties was also investigated for both

materials (Chapters III and V) (research objective 3). Results show that is possible to

tailor the properties of each single material and the main conclusions are:

• The increase of the degree of modification increases the crosslinking density,

printability and mechanical properties;

• The increase of the degree of modification decreases the swelling properties

and slow down the degradation process;

• The increase of the degree of modification produces structures with high

number of pores with small pore sizes.

Based on the results presented in Chapters III to V the following single material

systems were considered (research objective 3):

• AlgMA (2% wt of alginate and 20% v/v of methacrylate anhydride)

functionalized during 24 h and dissolved in a solution containing 1% wt of

photoinitiator;

• GelMA (12.5% wt of gelatin and 10% v/v of methacrylate anhydride)

functionalized during 6 h and dissolved in a solution containing 1% wt of

photoinitiator;

These systems were used to encapsulate mesenchymal stem cells (Chapter V) and

human chondrocytes (Chapter IV) and cell laden constructs were produced using a

UV lamp (8 mW/cm2 of light intensity) and 8 min of irradiation time (research

objective 3). Results show that:

• All material showed no toxicity effects;

• All materials were able to support cell growth. However, fast cell

proliferation was observed for GelMA, which can be attributed to the

presence of specific cell adhesion motifs as discussed in Chapter V.

Chapter VI investigates hybrid systems obtained through the combination of different

ratios between alginate and gelatin. These materials were investigated before and after

photopolymerisation and the effect of material composition in terms of rheological,

mechanical, swelling, degradation properties and biological properties was

investigated (research objective 4). Results show that:

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• AlgMA and GelMA hybrid systems are able to support cell growth and ECM

formation;

• The addition of AlgMA increases the mechanical properties, swelling

characteristics and slow down the degradation process;

• The levels of collagen, aggrecan and GAGs formation were significantly

higher in the hybrid systems in comparison to the single material systems.

Moreover, GelMA presents higher levels of ECM formation in comparison to

AlgMA.

Finally, Chapter VII investigates the use of a scaffold produced with a novel additive

manufacturing system and comprising two main materials (PCL and hybrid

alginate/gelatin system). This construct was investigated aiming to define a proper

vertical integration between the subchondral bone and cartilage (research objective 5).

Due to hydrophobic the nature of PCL (a material commonly used for bone tissue

engineering scaffolds) plasma treatment was used. Results showed that it was not only

possible to create a construct using these two different polymeric systems but also that

the plasma treatment positively influenced cell attachment and proliferation.

Future works

This thesis discusses the use of hydrogel-based constructs scaffolds for tissue

engineering produced for cartilage replacements through additive manufacturing. An

extensive in vitro work was conducted to design and assess the materials and the 3D

produced scaffolds. Despite the contributions to the current knowledge provided by

this research work, it was possible also to open new research questions that will be

addressed in the future. Moreover, additional work should be conducted to complete

the current research. Planned activities include:

(1) A zonal construct containing layers of alginate, gelatin and alginate/gelatin will be

produce encapsulating chondrocytes but also stem cells and growth factors. The

construct will be characterized in terms of overall mechanical properties, adhesion

between layers and the ability to sustain cell adhesion, proliferation and

differentiation. The zonal construct will be produced using the materials

investigated in these research work;

(2) Investigate new functionalization methods for both alginate and gelatin by

introducing C=C on the hydroxyl (e.g. oxidation, sulfation and copolymerization)

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and carboxyl (esterification and amidation) groups in order to improve both

mechanical, rheological, biological and degradability characteristics;

(3) Investigate the effect of light intensity, irradiation time and photoinitiator

concentration on the different properties of the printed constructs;

(4) The lubrication properties of cartilage, particularly the upper zone of cartilage, is

an important property not considered in this research. The coefficient of friction,

under both dynamic and static loading conditions, of the different materials as a

function of functionalization and processing conditions will be assessed;

(5) Cartilage is an avascular tissue and as a consequence chondrogenesis improves

under hypoxia conditions. The effect of the functionalization process and printing

conditions in terms of oxygen permeability will be investigated allowing to design

constructs creating a more suitable biological environment;

(6) In order to have a deeper understanding on the interactions between the hydrogel

and cells, additional protein and gene expression tests will be conducted. This will

include the quantification of the transcription factor SOX-6 and SOX-9, Collagen

II, and cartilage oligomeric protein (COMP) of chondrocytes after long periods of

culture;

(7) Free radical polymerization was used in this research. This is a chain-growth

reaction allowing the spatiotemporal control of the polymeric network but

producing heterogeneous networks. In the future, other chemistries such as

photoclick reactions (step-growth mechanism) will be explored to allow fast and

more efficient reactions with minimal network defects;

(8) Scaffolds will be also characterized in vivo. Initially, constructs will be

subcutaneously implanted into the flanks of nude mice to investigate swelling,

degradability and neocartilage formation. Later, more detailed studies will be

performed considering osteochondral defects created in rabbits.

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