Biomaterials for organ and tissue regeneration: new technologies and future prospects 9780081029060, 9780081029077, 1771771771, 0081029063, 0081029071

Table of contents :
Front Cover......Page 1
Biomaterials for Organ and Tissue Regeneration......Page 4
Copyright Page......Page 5
Contents......Page 6
List of contributors......Page 20
Preface......Page 28
Acknowledgment......Page 30
1 Properties and forms of biomaterials......Page 32
1.1 Introduction......Page 34
1.2 Many facets of new biomaterials: new naturally sourced biomaterials, new synthetic biomaterials, materiomics, metabioma.........Page 35
1.3 Off-shoot technologies linked to biomaterials and tissue engineering: biorobotics, bioinks, and bioprinting......Page 39
1.4 Biomaterial risk assessment......Page 41
References......Page 46
2.1 Introduction......Page 50
2.2.2 Mechanical properties......Page 51
2.2.4 Control of porosity, pore size, and pore connectivity......Page 53
2.3.1 Surface energy-hydrophilicity......Page 54
2.3.4 Protein adsorption......Page 56
2.3.5 Versatile modification of the biomaterials’ surface chemistry......Page 57
2.3.7 Antibacterial properties......Page 58
2.4 Properties of biomimetic biomaterials......Page 59
2.6 Conclusion and perspectives......Page 60
References......Page 61
Abbreviations......Page 64
3.1 Introduction......Page 65
3.2 A brief history of composites......Page 66
3.3.1 Mechanical properties......Page 67
3.3.2 Biodegradation properties......Page 70
3.4.1 Bone......Page 74
3.4.1.1 Composite with bioactive ceramics/glasses......Page 75
3.4.1.1.1 Composites with calcium phosphate–based bioceramics......Page 76
3.4.1.1.2 Composites with bioactive glasses......Page 77
3.4.1.2 Nanocomposite materials mimicking mineralized collagen fibrils......Page 80
3.4.1.4 Magnetized composites......Page 81
3.4.2.1 Composites of natural polymers......Page 82
3.4.2.2.2 Oligo(poly(ethylene glycol)fumarate) gel/gelatin microparticles......Page 83
3.4.2.3 Composites of synthetic polymers......Page 84
3.4.3 Tendon, ligament, and enthesis regeneration......Page 85
3.4.4 Skeletal muscle regeneration......Page 86
3.4.4.2 Synthetic composites......Page 87
3.5.2 Importance of vasculature and innervation for bone regeneration......Page 88
3.6.1 Composite biomaterial approach for nerve regeneration......Page 89
3.6.1.1 Composites of synthetic polymers (copolymers)......Page 90
3.6.1.2 Composites containing carbon nanostructures......Page 91
3.6.1.3 Polymer–ceramics composites......Page 92
3.6.1.5.1 Multichannel nerve guide conduit......Page 93
3.6.1.5.2 Electrospun fibrous nerve guide conduit......Page 94
3.6.1.5.3 Nerve guide conduit with intraluminal guidance......Page 95
3.7.1 Fabrication of polymer-based composite scaffolds that incorporate a vascular network......Page 96
3.7.2.1 Synthetic polymer–based composites containing collagen and gelatin......Page 97
3.8 Conclusion and future prospects......Page 98
References......Page 99
4.1 Introduction......Page 114
4.2 Analytical methodologies for protein identification and monitoring......Page 115
4.3 Mass spectrometry–based proteomic analysis......Page 118
4.3.1 Sample preparation for proteomics experiments......Page 119
4.3.2 Peptide mixture analysis by liquid chromatography coupled to tandem mass spectrometry......Page 121
4.3.3 Protein identification......Page 123
4.3.4 Applications of proteomic analysis to the development of biomaterials for tissue regeneration......Page 125
4.4 Analytical methodologies adapted to protein structural characterization......Page 129
4.4.1 Sample preparation......Page 131
4.4.2 Tandem mass spectrometry–based analysis of posttranslational modifications......Page 132
4.5 Conclusion......Page 135
References......Page 136
5.1 Introduction......Page 144
5.2.2 Delivery systems for tissue and organ regeneration......Page 145
5.3 Nanoscale-delivery systems for regeneration purposes......Page 146
5.3.1.1 Nanoparticles......Page 147
Carbohydrate-based nanoparticles......Page 148
Protein-based nanoparticles......Page 151
Nanoparticles of synthetic polymers......Page 152
5.3.1.1.2 Lipid nanoparticles......Page 155
Ceramic nanoparticles......Page 156
Metallic nanoparticles......Page 160
Quantum dots......Page 162
5.3.1.2.1 Carbon nanotubes......Page 164
5.3.1.2.3 Metallic nanotubes......Page 165
5.3.1.3.2 Composite nanorods......Page 167
5.3.1.4.2 Organosilicon nanocages......Page 168
5.4.1 Microfluidic devices for production of nanoparticles......Page 170
5.4.2 Recruiting 3D printing and nanoparticles for tissue engineering applications......Page 171
5.5 Conclusion and future perspectives......Page 172
References......Page 173
6.1.1 From artificial to bioinspired materials: challenges at the cell–material interface......Page 194
6.2 Engineering the cell–material interface......Page 195
6.2.1 Surface mimics of the extracellular matrix......Page 196
6.2.2 Chemical and spatial control of cell adhesion to surface materials......Page 197
6.3 Delivery strategies for growth factors at the cell–material interface......Page 200
6.3.1 Biofunctionalization strategies for tailoring the spatiotemporal delivery of growth factors......Page 201
6.3.2 Guidance of cell responses by growth factors complexed with surface materials......Page 202
6.4 Conclusion and outlook......Page 203
References......Page 204
7.2.1 Self-renewal......Page 208
7.2.2 Potency......Page 210
7.3.1 Embryonic stem cells......Page 211
7.3.2.1 Reprograming techniques......Page 213
7.3.3 Adult stem cells......Page 214
7.3.3.2 Mesenchymal stem cells......Page 215
7.3.3.2.2 Adipose-derived stem cells......Page 216
7.3.3.2.3 Bone marrow–derived mesenchymal stem cells and adipose-derived stem cells......Page 218
7.4 Applications of stem cells in tissue engineering......Page 219
7.5 Conclusion......Page 223
References......Page 224
Abbreviations......Page 228
8.1 Introduction......Page 229
8.2.1.1 Just a matter of recognition......Page 230
8.2.1.2 Just a matter of amplification and increased efficiency......Page 232
8.2.2 Tissue regeneration/wound healing (“why immune system is so important”)......Page 233
8.2.3 Immune response to biomaterials: when all goes wrong, that is, the foreign body reaction......Page 236
8.3.1 Myeloid cells......Page 240
8.3.2 Innate-like lymphocytes......Page 244
8.3.3 Lymphocytes......Page 247
8.4 Immune cell sourcing......Page 251
8.5 In vivo testing......Page 253
References......Page 254
9.1 Introduction......Page 262
9.1.1 How to modulate cell adhesion, cell migration, and cell extrusion?......Page 263
9.1.2 Synthetic matrices to control cell programming and reprogramming......Page 266
9.1.3 Nuclear mechanics and mechanical memory......Page 270
9.2 Conclusion......Page 272
References......Page 273
2 Biomaterials use in organ specific applications......Page 278
10.1 Introduction......Page 280
10.2.1 Arterial tissue......Page 281
10.2.2 Cardiac tissue......Page 282
10.3 Cardiovascular disease......Page 283
10.4 Coronary artery bypass grafting......Page 284
10.4.1 Vascular grafts......Page 285
10.4.2 Role of biomechanical compliance......Page 286
10.5 Tissue-engineered blood vessels......Page 288
10.5.1 Biomaterials for tissue-engineered blood vessels......Page 292
10.5.2 Stem cells in tissue-engineered blood vessel applications......Page 293
10.6 Electrospinning of tissue-engineered blood vessels......Page 294
10.6.3 Collector systems for creating electrospun vessels......Page 295
10.7 Future outlook for cardiovascular tissue engineering......Page 297
References......Page 298
11.2 The small intestine—structural organization and function......Page 304
11.3.1 Modeling the small intestine in vitro by two-dimensional monolayer cell cultures......Page 309
11.3.2 Small intestinal organoids—artificial mini organs grown in vitro......Page 311
11.3.2.1 Organoid cell sources......Page 312
11.3.2.2 Organoids as in vitro tools to model or study intestinal diseases......Page 313
11.4 Small intestinal tissue engineering in the Transwell—when cells meet scaffolds......Page 314
11.5 Next-generation models—integration of microenvironmental factors......Page 315
11.6 Outlook......Page 319
References......Page 320
12.3 Pancreatic islet......Page 330
12.3.1.3 δ-Cells: somatostatin......Page 332
12.3.2 β-Cells role and insulin function......Page 333
12.4 Diabetes......Page 334
12.4.2 Type 2 diabetes......Page 335
12.4.5.1 Acute complications......Page 336
12.5.1.2 Interstitial blood measurements......Page 337
12.5.2.1 Multiple daily injections......Page 338
12.5.2.3 Artificial pancreas......Page 339
12.5.3.1 Pancreas transplantation......Page 340
12.5.3.2 Pancreatic islet transplantation......Page 341
12.6.1 Low isolation yield and high pancreas requirement......Page 342
12.6.4 Instant blood-mediated inflammatory reaction......Page 343
12.7 Improvements in islet transplantation (Fig. 12.7)......Page 344
12.8.1.1 Pig islet function......Page 347
12.8.1.2 Risk of zoonosis......Page 348
12.8.2.1 Embryonic stem cells......Page 349
12.8.2.2 Induced pluripotent stem cells......Page 350
12.9.1 Definition......Page 351
12.9.2 Microencapsulation......Page 352
12.9.3 Macroencapsulation......Page 353
References......Page 356
13.2.1 Diabetic foot ulcer complications......Page 366
13.2.2 Cellular and molecular events in diabetic foot ulcer......Page 367
13.2.3 Implication of advanced glycation end products on cell function......Page 368
13.2.3.3 Keratinocytes......Page 369
13.2.4 Epigenetic changes related to diabetic foot ulcer......Page 370
13.3.2 Oxygenation......Page 371
13.3.5 Cross-linking of biopolymers......Page 372
13.4 Biomaterials supporting the administration of bioactive agents......Page 374
13.4.1 Therapy based on growth factors......Page 375
13.4.2 Pharmacological treatment alternative to growth factors......Page 376
13.4.4 Gene therapy–based approaches......Page 378
13.4.5 Cell therapy–based approaches......Page 379
13.5.1 Natural extracellular matrix biomaterials......Page 381
13.5.2 Peptide-based biomaterials......Page 382
13.5.3 Inorganic agent–containing dressings......Page 383
References......Page 384
14.1 Background......Page 394
14.2 Bone morphogenetic protein for bone regeneration......Page 395
14.2.1 Bone morphogenetic protein delivery via carrier materials......Page 396
14.2.2 Clinical products......Page 399
14.3.2 Complications and risks......Page 401
14.3.3 Implant considerations......Page 402
14.4 Current strategies......Page 403
14.4.1 Minimizing dose......Page 404
14.4.2.1 Microparticle bone morphogenetic protein encapsulation......Page 405
14.4.2.2 Matrix bone morphogenetic protein encapsulation......Page 407
14.4.2.3 Matrix bone morphogenetic protein surface coatings......Page 409
14.4.3.1 Bioactive materials......Page 410
14.4.3.2 Advanced biomanufacturing......Page 411
References......Page 412
15.2 The adipose cells for the reconstruction of in vitro models......Page 424
15.2.2 Primary bone marrow mesenchymal stem cells......Page 425
15.2.6 Human cells or other species?......Page 426
15.3 Current existing in vitro adipose tissues models......Page 427
15.3.1 Reconstruction of an in vitro adipose tissue without scaffold......Page 429
15.3.3 Natural components in adipose tissue engineering......Page 430
15.4 Medical applications of adipose tissues grafts......Page 436
15.4.2 For reconstructive surgery......Page 437
15.5.1 Vascularized adipose tissues......Page 438
15.5.2.4 Muscle cells......Page 439
15.5.3.1 “White” (fat storage adipose tissue)......Page 440
15.5.3.4 “Pink” (breast adipose tissue)......Page 441
15.6 Conclusion......Page 442
References......Page 443
16.1.2 Endothelial cells, pericytes, and astrocytes......Page 456
16.1.3 Extracellular matrix......Page 457
16.2 Spheroids......Page 458
16.3.2 Extracellular matrix channels......Page 459
16.4 Sprouts and guided capillaries growth......Page 460
16.5.1 Capillaries self-organization on top of Matrigel®......Page 461
16.5.3 Capillaries self-organization in device-free hydrogels......Page 462
16.6 Current challenges in translational research......Page 463
16.7 Implantation prospects......Page 464
References......Page 465
17.1 Introduction......Page 472
17.2.1 Kidney tissue engineering......Page 475
17.2.2 Bladder tissue engineering......Page 478
17.3 Conclusion and future perspectives......Page 481
References......Page 482
Further reading......Page 486
18.2.1 Embryology......Page 488
18.2.2.1 Descriptive anatomy......Page 489
18.2.2.2 Endoscopic anatomy......Page 490
18.2.3.1 Anatomy......Page 491
18.2.3.2 Histology......Page 492
18.2.3.3 Physiology......Page 493
18.3.1 Laryngeal diseases......Page 494
18.3.2.2 Oeso-tracheal fistulas......Page 495
18.4.2 Biomaterials......Page 496
18.4.3.1.1 Manufactured scaffolds......Page 498
18.4.3.2.1 Cartilage......Page 499
18.4.3.2.2 Respiratory epithelium......Page 500
18.5.2 Autografts......Page 501
18.5.3 Allografts......Page 502
References......Page 503
19.1.1.2 Red blood cells or erythrocytes......Page 508
19.1.1.3 White blood cells or leukocytes......Page 509
19.1.2.1 History......Page 510
19.1.2.2 Platelet action......Page 512
19.1.3.2 Centrifugation......Page 514
19.1.3.4 Activation method to induce platelet degranulation and release of growth factors......Page 515
19.2.1 An autologous cell culture supplement......Page 516
19.2.2 Platelet-rich plasma in tissue-engineered constructs......Page 517
19.3 Platelet-rich plasma in regenerative medicine......Page 519
References......Page 523
3 Emerging and enabling technologies for biomaterials in tissue regeneration......Page 528
20.1 Introduction......Page 530
20.2 Conventional hydrogels and their limitations......Page 531
20.2.1.1 Polysaccharides-based hydrogels......Page 532
20.2.1.2 Protein-based hydrogels......Page 534
20.2.2 Synthetic polymers......Page 536
20.3.1.1 Incorporation of prefabricated nanomaterials......Page 539
20.3.2 Nanoparticles in tissue engineering......Page 540
20.4.1 Tailored mechanical and structural properties......Page 544
20.4.2 Enhanced electrical conductivity......Page 546
20.4.3 Enhanced availability of biological factors and drugs......Page 547
20.4.4 Cellular reprogramming......Page 551
20.5 Conclusion and future directions......Page 552
Acknowledgments......Page 553
References......Page 554
21.1 Introduction......Page 560
21.2.1 Graphene......Page 561
21.3 Function mimetic carbon-based engineered tissues......Page 563
21.3.1 Skeletal muscle regeneration......Page 564
21.3.2 Cardiac tissue regeneration......Page 566
21.3.3 Neural tissue regeneration......Page 568
21.4 Bone regeneration......Page 570
21.5.2 Biodegradability......Page 574
21.6 Conclusion and future perspectives......Page 575
References......Page 576
22.1 Introduction......Page 582
22.2.1 Hyaluronic acid......Page 583
22.3 Cross-linking chemistry of hyaluronic acid......Page 586
22.3.1 Schiff-base cross-linking hydrogels......Page 587
22.3.2 Diels–Alder click cross-linked hydrogel......Page 591
22.4.1 Hyaluronic acid–based scaffolds......Page 592
References......Page 594
23.2 Design considerations of microfluidics chips......Page 598
23.2.1 Photolithography......Page 599
23.2.2 Microcontact printing......Page 601
23.2.3 Micropatterning of cells on microchannels......Page 603
23.2.4 Cryopreservation techniques of cells for tissue engineering......Page 605
23.3.1 Composite microparticles......Page 607
23.3.2 Particulate biomaterials at the nanoscale......Page 608
23.3.3 Fibrous biomaterials at micro- and nanoscale......Page 609
23.4 Methods for cell patterning and cultivation......Page 610
23.4.2 Bioreactors......Page 612
23.4.3 Microfluidic devices for cell manipulation......Page 614
23.4.4 Microenvironment on cell integrity......Page 615
23.5.2 Vascular tissue......Page 616
23.5.3 Liver......Page 619
23.5.4 Bone......Page 620
23.6 Conclusion......Page 621
References......Page 622
Further reading......Page 626
24.1 Introduction......Page 630
24.2 Biomechanics and mechanobiology......Page 632
24.3 Mechanobiology and biomaterials functionality......Page 634
24.4 Methods and challenges......Page 639
24.5 Biomaterials evaluation: a practical example......Page 642
24.6 Conclusion and outlook......Page 649
References......Page 650
25.1.1.1 Photo-patterned microgels......Page 660
25.1.1.2 Emulsion-based microgels......Page 662
25.1.1.3 Bioactive microfibers......Page 663
25.1.2 Microfluidic tissue models and microphysiological systems......Page 664
25.1.3 Emerging biofabrication technologies......Page 667
25.1.4 Stem cell technology—biomaterial interface......Page 669
25.1.5 Organoids......Page 670
25.1.6 Rationale design of biomaterials for disease modeling......Page 672
25.1.7 Biocompatibility......Page 673
25.1.9 Vascularity......Page 675
25.1.11 Electrical conductivity......Page 678
25.2.1 Inflammatory response and cancer modeling......Page 679
25.2.2 Cardiovascular diseases......Page 681
25.2.3 Skin diseases......Page 682
25.2.4 Gastrointestinal diseases......Page 683
25.2.5 Neurological disorders......Page 684
25.3 Conclusion......Page 685
References......Page 686
26.1 Introduction......Page 700
26.3 Selection parameters for biomedical applications......Page 702
26.3.1 Lung......Page 704
26.3.2 Brain......Page 706
26.3.3 Heart......Page 707
26.3.4 Kidney......Page 708
26.3.5 Liver......Page 709
26.3.7 Muscle......Page 710
26.3.8 Bone......Page 712
26.3.9 Multiorgan......Page 713
26.4 Organ-on-chip platforms to mimic human pathophysiology......Page 715
26.5.1 Space......Page 717
26.5.2 Military......Page 718
26.6.1 Elastomers......Page 719
26.7 Thermoplastics......Page 720
26.8 Hydrogels......Page 721
26.9 Biomaterials for tissue fabrication for organ-on-chip platforms......Page 722
26.10 Challenges and outlook......Page 726
References......Page 727
Further reading......Page 738
27.1 The role of bioreactors in tissue engineering......Page 740
27.2.1 Stirred bioreactors......Page 742
27.2.2 Wave bioreactors......Page 743
27.2.3 Parallel-plate bioreactors and parallel-plate flow chamber bioreactors......Page 745
27.2.4 Rotating wall vessel (reduced gravity) bioreactors......Page 746
27.2.5 Strain bioreactors......Page 748
27.2.6 Perfusion bioreactors......Page 753
27.2.7 Hollow-fiber bioreactors......Page 756
27.2.8 Microfluidic bioreactors......Page 758
27.2.9 Combined systems......Page 760
27.3 Cell-seeding techniques for bioreactors......Page 762
27.4 Design considerations and future outlook......Page 764
References......Page 770
Further reading......Page 783
28.1 Introduction......Page 784
28.1.1 General overview......Page 785
28.1.2 Review of the lung and liver cell line models......Page 788
28.2.1 Finite-difference method......Page 791
28.2.2 Finite element modeling......Page 793
28.3 Modeling of bioreactor for lung cells......Page 794
28.4 Mathematical modeling of liver cells......Page 804
28.5 Conclusion......Page 815
References......Page 817
Further reading......Page 821
Index......Page 822
Back Cover......Page 847

Citation preview

Biomaterials for Organ and Tissue Regeneration

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Woodhead Publishing Series in Biomaterials

Biomaterials for Organ and Tissue Regeneration New Technologies and Future Prospects

Edited by

Nihal Engin Vrana Helena Knopf-Marques Julien Barthes

Woodhead Publishing is an imprint of Elsevier The Officers’ Mess Business Centre, Royston Road, Duxford, CB22 4QH, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom Copyright © 2020 Elsevier Ltd. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-08-102906-0 (print) ISBN: 978-0-08-102907-7 (online) For information on all Woodhead Publishing publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisitions Editor: Sabrina Webber Editorial Project Manager: Peter Adamson Production Project Manager: Joy Christel Neumarin Honest Thangiah Cover Designer: Christian J. Bilbow Cover Image Credit: Ilker Gazioglu Typeset by MPS Limited, Chennai, India

Contents

List of contributors Preface Acknowledgment

Section 1 1

2

xix xxvii xxix

Properties and forms of biomaterials

Introduction to biomaterials for tissue/organ regeneration Nihal Engin Vrana 1.1 Introduction 1.2 Many facets of new biomaterials: new naturally sourced biomaterials, new synthetic biomaterials, materiomics, metabiomaterials 1.3 Off-shoot technologies linked to biomaterials and tissue engineering: biorobotics, bioinks, and bioprinting 1.4 Biomaterial risk assessment 1.5 Conclusion Acknowledgment References Physicochemical properties of biomaterials Vincent Ball 2.1 Introduction 2.2 Bulk properties of biomaterials 2.2.1 Shape and size control 2.2.2 Mechanical properties 2.2.3 Corrosion and degradation in a given chemical environment 2.2.4 Control of porosity, pore size, and pore connectivity 2.3 Surface properties of biomaterials 2.3.1 Surface energy-hydrophilicity 2.3.2 Lack of toxicity, of unfavorable immunological response, hemocompatibility 2.3.3 Surface topography 2.3.4 Protein adsorption 2.3.5 Versatile modification of the biomaterials’ surface chemistry 2.3.6 Degradability of surface coatings 2.3.7 Antibacterial properties 2.3.8 Active biomaterials

1 3 3 4 8 10 15 15 15 19 19 20 20 20 22 22 23 23 25 25 25 26 27 27 28

vi

Contents

2.4 2.5

3

4

Properties of biomimetic biomaterials Real-time monitoring of an implanted biomaterial and personalized implants 2.6 Conclusion and perspectives References

28

Polymer-based composites for musculoskeletal regenerative medicine Patrina S.P. Poh, Maria A. Woodruff and Elena Garcı´a-Gareta Abbreviations 3.1 Introduction 3.2 A brief history of composites 3.3 Polymer-based composites scaffold characteristics 3.3.1 Mechanical properties 3.3.2 Biodegradation properties 3.4 Polymer-based composite scaffolds for specific musculoskeletal tissue regeneration 3.4.1 Bone 3.4.2 Cartilage and osteochondral regeneration 3.4.3 Tendon, ligament, and enthesis regeneration 3.4.4 Skeletal muscle regeneration 3.5 The necessity for nerve and vascular regeneration 3.5.1 Importance of vasculature and innervation for skeletal muscle regeneration 3.5.2 Importance of vasculature and innervation for bone regeneration 3.6 Nerve regeneration 3.6.1 Composite biomaterial approach for nerve regeneration 3.7 Vascular regeneration 3.7.1 Fabrication of polymer-based composite scaffolds that incorporate a vascular network 3.7.2 Engineering of small diameter blood vessels (,6 mm) with polymer-based composites 3.8 Conclusion and future prospects Acknowledgments References

33

Emerging biotechnological approaches with respect to tissue regeneration: from improving biomaterial incorporation to comprehensive omics monitoring Rabah Gahoual, Yannis-Nicolas Franc¸ois, Nathalie Mignet and Pascal Houze´ 4.1 Introduction 4.2 Analytical methodologies for protein identification and monitoring 4.3 Mass spectrometry
based proteomic analysis 4.3.1 Sample preparation for proteomics experiments

29 29 30

33 34 35 36 36 39 43 43 51 54 55 57 57 57 58 58 65 65 66 67 68 68

83

83 84 87 88

Contents

vii

4.3.2

Peptide mixture analysis by liquid chromatography coupled to tandem mass spectrometry 4.3.3 Protein identification 4.3.4 Applications of proteomic analysis to the development of biomaterials for tissue regeneration 4.4 Analytical methodologies adapted to protein structural characterization 4.4.1 Sample preparation 4.4.2 Tandem mass spectrometry
based analysis of posttranslational modifications 4.5 Conclusion References 5

6

Use of nanoscale-delivery systems in tissue/organ regeneration Milad Fathi-Achachelouei, Dilek Keskin and Aysen Tezcaner 5.1 Introduction 5.2 Properties and application areas of nanoscale-delivery systems in biomedical field 5.2.1 Delivery systems for therapeutic purpose 5.2.2 Delivery systems for tissue and organ regeneration 5.3 Nanoscale-delivery systems for regeneration purposes 5.3.1 Morphological classification of nanoscale-delivery systems 5.4 Emerging delivery technologies in nanoparticle area 5.4.1 Microfluidic devices for production of nanoparticles 5.4.2 Recruiting 3D printing and nanoparticles for tissue engineering applications 5.5 Conclusion and future perspectives References

90 92 94 98 100 101 104 105 113 113 114 114 114 115 116 139 139 140 141 142

Surface functionalization of biomaterials for cell biology applications 163 E. Ada Cavalcanti-Adam and Wenqian Feng 6.1 Introduction 163 6.1.1 From artificial to bioinspired materials: challenges at the cell
material interface 163 6.2 Engineering the cell
material interface 164 6.2.1 Surface mimics of the extracellular matrix 165 6.2.2 Chemical and spatial control of cell adhesion to surface materials 166 6.3 Delivery strategies for growth factors at the cell
material interface 169 6.3.1 Biofunctionalization strategies for tailoring the spatiotemporal delivery of growth factors 170 6.3.2 Guidance of cell responses by growth factors complexed with surface materials 171 6.4 Conclusion and outlook 172 References 173

viii

7

8

9

Contents

Stem cells: sources, properties, and cell types Melis Asal and Sinan Gu¨ven 7.1 Introduction 7.2 Stem cell properties 7.2.1 Self-renewal 7.2.2 Potency 7.3 Cell types 7.3.1 Embryonic stem cells 7.3.2 Induced pluripotent stem cells 7.3.3 Adult stem cells 7.4 Applications of stem cells in tissue engineering 7.5 Conclusion References

177 177 177 177 179 180 180 182 183 188 192 193

Immune cells: sources, properties, and cell types S. Jung and Florent Meyer Abbreviations 8.1 Introduction 8.2 Immune system consideration in the use of biomaterials and tissue regeneration 8.2.1 Overall description of the immune system: innate versus adaptive system (“know your basics”) 8.2.2 Tissue regeneration/wound healing (“why immune system is so important”) 8.2.3 Immune response to biomaterials: when all goes wrong, that is, the foreign body reaction 8.3 Immune cell description 8.3.1 Myeloid cells 8.3.2 Innate-like lymphocytes 8.3.3 Lymphocytes 8.4 Immune cell sourcing 8.5 In vivo testing 8.6 Conclusion References

197

Cell signaling and strategies to modulate cell behavior Claire Ehlinger, Dominique Vautier and Leyla Kocgozlu 9.1 Introduction 9.1.1 How to modulate cell adhesion, cell migration, and cell extrusion? 9.1.2 Synthetic matrices to control cell programming and reprogramming 9.1.3 Nuclear mechanics and mechanical memory 9.2 Conclusion Acknowledgments References

231

197 198 199 199 202 205 209 209 213 216 220 222 223 223

231 232 235 239 241 242 242

Contents

Section 2 10

11

12

ix

Biomaterials use in organ specific applications

Cardiovascular tissue engineering Richard A. O’Connor, Paul A. Cahill and Garrett B. McGuinness 10.1 Introduction 10.2 The cardiovascular system 10.2.1 Arterial tissue 10.2.2 Cardiac tissue 10.3 Cardiovascular disease 10.4 Coronary artery bypass grafting 10.4.1 Vascular grafts 10.4.2 Role of biomechanical compliance 10.5 Tissue-engineered blood vessels 10.5.1 Biomaterials for tissue-engineered blood vessels 10.5.2 Stem cells in tissue-engineered blood vessel applications 10.6 Electrospinning of tissue-engineered blood vessels 10.6.1 Fundamentals of electrospinning 10.6.2 Electrospinning parameters 10.6.3 Collector systems for creating electrospun vessels 10.6.4 Limitations of electrospun scaffolds 10.7 Future outlook for cardiovascular tissue engineering Acknowledgments References Bioartificial gut—current state of small intestinal tissue engineering ¨ Thomas Daullary, Christina Fey, Constantin Berger, Marco Metzger and Daniela Zdzieblo 11.1 Introduction 11.2 The small intestine—structural organization and function 11.3 Modeling the small intestine—biology meets engineering 11.3.1 Modeling the small intestine in vitro by two-dimensional monolayer cell cultures 11.3.2 Small intestinal organoids—artificial mini organs grown in vitro 11.4 Small intestinal tissue engineering in the Transwell—when cells meet scaffolds 11.5 Next-generation models—integration of microenvironmental factors 11.6 Outlook References From insulin replacement to bioengineered, encapsulated organoids Elisa Maillard and Se´verine Sigrist 12.1 Introduction 12.2 Pancreas

247 249 249 250 250 251 252 253 254 255 257 261 262 263 264 264 264 266 266 267 267 273

273 273 278 278 280 283 284 288 289 299 299 299

x

Contents

12.3

13

Pancreatic islet 12.3.1 Composition of pancreatic islets 12.3.2 β-Cells role and insulin function 12.4 Diabetes 12.4.1 Type 1 diabetes 12.4.2 Type 2 diabetes 12.4.3 Gestational diabetes 12.4.4 Other diabetes 12.4.5 Poor glycemia regulation complications 12.5 Insulin replacement for type 1 diabetes 12.5.1 Glucose measurements 12.5.2 Exogenous insulin 12.5.3 Endogenous insulin production 12.6 Islet transplantation limits (Fig. 12.5) 12.6.1 Low isolation yield and high pancreas requirement 12.6.2 Extracellular matrix destruction 12.6.3 Hypoxia 12.6.4 Instant blood-mediated inflammatory reaction 12.6.5 Autoimmunity and alloimmunity 12.6.6 Immune suppressive regimen 12.7 Improvements in islet transplantation (Fig. 12.7) 12.8 Other sources of insulin-secreting cells 12.8.1 Cells of animal origin 12.8.2 Surrogate cells 12.9 The bioartificial pancreas 12.9.1 Definition 12.9.2 Microencapsulation 12.9.3 Macroencapsulation 12.10 Conclusion References

299 301 302 303 304 304 305 305 305 306 306 307 309 311 311 312 312 312 313 313 313 316 316 318 320 320 321 322 325 325

Diabetic wound healing with engineered biomaterials Laura E. Castellano, Jorge Delgado, Arturo Vega-Gonza´lez and Birzabith Mendoza-Novelo 13.1 Introduction 13.2 Impaired wound healing under condition of diabetes 13.2.1 Diabetic foot ulcer complications 13.2.2 Cellular and molecular events in diabetic foot ulcer 13.2.3 Implication of advanced glycation end products on cell function 13.2.4 Epigenetic changes related to diabetic foot ulcer 13.3 Physicochemical aspects and fabrication of biomaterials in diabetic wound healing 13.3.1 Hydration 13.3.2 Oxygenation 13.3.3 Infection control

335

335 335 335 336 337 339 340 340 340 341

Contents

13.3.4 Nanoparticle synthesis and its bioactivity 13.3.5 Cross-linking of biopolymers 13.4 Biomaterials supporting the administration of bioactive agents 13.4.1 Therapy based on growth factors 13.4.2 Pharmacological treatment alternative to growth factors 13.4.3 Treatment with natural extracts 13.4.4 Gene therapy
based approaches 13.4.5 Cell therapy
based approaches 13.5 Biomaterials with prohealing activity 13.5.1 Natural extracellular matrix biomaterials 13.5.2 Peptide-based biomaterials 13.5.3 Inorganic agent
containing dressings 13.6 Final remarks References 14

15

Bone morphogenetic protein
assisted bone regeneration and applications in biofabrication Naomi C. Paxton, Cynthia S. Wong, Mathilde R. Desselle, Mark C. Allenby and Maria A. Woodruff 14.1 Background 14.2 Bone morphogenetic protein for bone regeneration 14.2.1 Bone morphogenetic protein delivery via carrier materials 14.2.2 Clinical products 14.3 Bone morphogenetic protein limitations 14.3.1 On- and off-label use 14.3.2 Complications and risks 14.3.3 Implant considerations 14.4 Current strategies 14.4.1 Minimizing dose 14.4.2 Controlled release systems 14.4.3 Complex biomaterials and biofabrication systems 14.5 Conclusion Acknowledgement References Adipose tissue engineering Fiona Louis and Michiya Matsusaki 15.1 Introduction 15.2 The adipose cells for the reconstruction of in vitro models 15.2.1 Preadipocytes cell lines 15.2.2 Primary bone marrow mesenchymal stem cells 15.2.3 Primary adipose
derived stem cells 15.2.4 Primary mature adipocytes 15.2.5 The importance of the vascularization in adipose models 15.2.6 Human cells or other species?

xi

341 341 343 344 345 347 347 348 350 350 351 352 353 353

363

363 364 365 368 370 370 370 371 372 373 374 379 381 381 381 393 393 393 394 394 395 395 395 395

xii

Contents

15.3

Current existing in vitro adipose tissues models 15.3.1 Reconstruction of an in vitro adipose tissue without scaffold 15.3.2 Use of synthetic scaffolds 15.3.3 Natural components in adipose tissue engineering 15.4 Medical applications of adipose tissues grafts 15.4.1 For cosmetic surgery 15.4.2 For reconstructive surgery 15.4.3 For wound healing 15.4.4 For bone healing 15.5 Further developments needed 15.5.1 Vascularized adipose tissues 15.5.2 Addition of other surrounding cells types 15.5.3 Reconstruction of the different types of adipose tissues 15.6 Conclusion References 16

17

396 398 399 399 405 406 406 407 407 407 407 408 409 411 412

Blood
brain barrier tissue engineering Agathe Figarol and Michiya Matsusaki 16.1 Introduction, specificities of the blood
brain barrier 16.1.1 A highly selective barrier 16.1.2 Endothelial cells, pericytes, and astrocytes 16.1.3 Extracellular matrix 16.1.4 First attempts in blood
brain barrier models 16.2 Spheroids 16.3 Templated vessels’ growth 16.3.1 Rigid channels 16.3.2 Extracellular matrix channels 16.4 Sprouts and guided capillaries growth 16.5 Capillaries self-organization 16.5.1 Capillaries self-organization on top of Matrigels 16.5.2 Capillary self-organization on microchips 16.5.3 Capillaries self-organization in device-free hydrogels 16.6 Current challenges in translational research 16.7 Implantation prospects 16.8 Conclusion References

425

Tissue engineering in urology Elif Vardar 17.1 Introduction 17.2 Biomaterials for urological tissues 17.2.1 Kidney tissue engineering 17.2.2 Bladder tissue engineering 17.3 Conclusion and future perspectives

441

425 425 425 426 427 427 428 428 428 429 430 430 431 431 432 433 434 434

441 444 444 447 450

Contents

Conflict of interest References Further reading 18

19

Respiratory tissue replacement and regeneration: from larynx to bronchi Lea Fath, Esteban Brenet, Dana M. Radu, Emmanuel Martinod and Christian Debry 18.1 Introduction 18.2 Normal respiratory tissue 18.2.1 Embryology 18.2.2 Larynx 18.2.3 Lower airways: trachea, carina, bronchi, bronchioles 18.3 Airways diseases 18.3.1 Laryngeal diseases 18.3.2 Tracheobronchial diseases 18.4 Replacement and regeneration strategies 18.4.1 Laryngeal transplantation 18.4.2 Biomaterials 18.4.3 Tissue engineering 18.5 Transplant 18.5.1 Nonliving tissue transplants 18.5.2 Autografts 18.5.3 Allografts 18.6 Conclusion and outlook References Platelet-rich plasma in tissue engineering Anne Lehn 19.1 Introduction 19.1.1 Blood composition 19.1.2 How does platelet-rich plasma work? 19.1.3 Preparation of platelet-rich plasma
based biomaterials 19.2 Tissue engineering 19.2.1 An autologous cell culture supplement 19.2.2 Platelet-rich plasma in tissue-engineered constructs 19.3 Platelet-rich plasma in regenerative medicine 19.4 Conclusion References

Section 3 Emerging and enabling technologies for biomaterials in tissue regeneration 20

Nanocomposite hydrogels for tissue engineering applications Azadeh Mostafavi, Jacob Quint, Carina Russell and Ali Tamayol 20.1 Introduction

xiii

451 451 455 457

457 457 457 458 460 463 463 464 465 465 465 467 470 470 470 471 472 472 477 477 477 479 483 485 485 486 488 492 492

497 499 499

xiv

Contents

20.2

Conventional hydrogels and their limitations 20.2.1 Natural polymers 20.2.2 Synthetic polymers 20.3 Nanomaterials for engineering composite hydrogel systems 20.3.1 Methods for creating nanocomposite hydrogel systems 20.3.2 Nanoparticles in tissue engineering 20.4 Properties of nanocomposite hydrogels 20.4.1 Tailored mechanical and structural properties 20.4.2 Enhanced electrical conductivity 20.4.3 Enhanced availability of biological factors and drugs 20.4.4 Cellular reprogramming 20.5 Conclusion and future directions Acknowledgments References 21

22

Functional carbon-based nanomaterials for engineered tissues toward organ regeneration Yasamin A. Jodat and Su Ryon Shin 21.1 Introduction 21.2 Characteristics of carbon-based materials used for tissue engineering 21.2.1 Graphene 21.2.2 Carbon nanotubes 21.3 Function mimetic carbon-based engineered tissues 21.3.1 Skeletal muscle regeneration 21.3.2 Cardiac tissue regeneration 21.3.3 Neural tissue regeneration 21.4 Bone regeneration 21.5 Considerations for in vivo tissue regeneration 21.5.1 Toxicity 21.5.2 Biodegradability 21.6 Conclusion and future perspectives References Hyaluronic acid
based hydrogels for tissue engineering N. Vijayakameswara Rao 22.1 Introduction 22.2 Chemical modifications of hyaluronic acid 22.2.1 Hyaluronic acid 22.3 Cross-linking chemistry of hyaluronic acid 22.3.1 Schiff-base cross-linking hydrogels 22.3.2 Diels
Alder click cross-linked hydrogel 22.3.3 Photo-cross-linking 22.3.4 The hyaluronic acid
disulfide cross-linking hydrogels

500 501 505 508 508 509 513 513 515 516 520 521 522 523

529 529 530 530 532 532 533 535 537 539 543 543 543 544 545 551 551 552 552 555 556 560 561 561

Contents

23

24

xv

22.4

Hyaluronic acid as a biomaterial in tissue engineering 22.4.1 Hyaluronic acid
based scaffolds 22.5 Conclusion Acknowledgement References

561 561 563 563 563

Microfluidics in tissue engineering Sudip Kumar Sinha and Arindam Bit 23.1 Introduction 23.2 Design considerations of microfluidics chips 23.2.1 Photolithography 23.2.2 Microcontact printing 23.2.3 Micropatterning of cells on microchannels 23.2.4 Cryopreservation techniques of cells for tissue engineering 23.3 Biomaterials at microscale 23.3.1 Composite microparticles 23.3.2 Particulate biomaterials at the nanoscale 23.3.3 Fibrous biomaterials at micro- and nanoscale 23.3.4 Sheet biomaterials 23.4 Methods for cell patterning and cultivation 23.4.1 Cell-patterning techniques 23.4.2 Bioreactors 23.4.3 Microfluidic devices for cell manipulation 23.4.4 Microenvironment on cell integrity 23.5 Microfluidic cell culture models for tissue engineering 23.5.1 Basal lamina 23.5.2 Vascular tissue 23.5.3 Liver 23.5.4 Bone 23.6 Conclusion References Further reading

567

Biomechanical characterization of engineered tissues and implants for tissue/organ replacement applications Michael Gasik 24.1 Introduction 24.2 Biomechanics and mechanobiology 24.3 Mechanobiology and biomaterials functionality 24.4 Methods and challenges 24.5 Biomaterials evaluation: a practical example 24.6 Conclusion and outlook References

567 567 568 570 572 574 576 576 577 578 579 579 581 581 583 584 585 585 585 588 589 590 591 595

599 599 601 603 608 611 618 619

xvi

25

26

Contents

In vitro disease and organ model Emal Lesha, Sheyda Darouie, Amir Seyfoori, Alireza Dolatshahi-Pirouz and Mohsen Akbari 25.1 Model development 25.1.1 Microengineered tissues 25.1.2 Microfluidic tissue models and microphysiological systems 25.1.3 Emerging biofabrication technologies 25.1.4 Stem cell technology—biomaterial interface 25.1.5 Organoids 25.1.6 Rationale design of biomaterials for disease modeling 25.1.7 Biocompatibility 25.1.8 Biodegradability 25.1.9 Vascularity 25.1.10 Mechanical properties 25.1.11 Electrical conductivity 25.2 Emerging applications and clinical considerations 25.2.1 Inflammatory response and cancer modeling 25.2.2 Cardiovascular diseases 25.2.3 Skin diseases 25.2.4 Gastrointestinal diseases 25.2.5 Neurological disorders 25.3 Conclusion References Biomaterials for on-chip organ systems Shabir Hassan, Marcel Heinrich, Berivan Cecen, Jai Prakash and Yu Shrike Zhang 26.1 Introduction 26.2 Design and biomaterial considerations for the development of specialized microphysiological systems 26.3 Selection parameters for biomedical applications 26.3.1 Lung 26.3.2 Brain 26.3.3 Heart 26.3.4 Kidney 26.3.5 Liver 26.3.6 Gut 26.3.7 Muscle 26.3.8 Bone 26.3.9 Multiorgan 26.4 Organ-on-chip platforms to mimic human pathophysiology 26.5 Applications beyond conventional research 26.5.1 Space 26.5.2 Military

629

629 629 633 636 638 639 641 642 644 644 647 647 648 648 650 651 652 653 654 655 669

669 671 671 673 675 676 677 678 679 679 681 682 684 686 686 687

Contents

26.6

27

28

xvii

Biomaterials for chip fabrication 26.6.1 Elastomers 26.7 Thermoplastics 26.8 Hydrogels 26.9 Biomaterials for tissue fabrication for organ-on-chip platforms 26.10 Challenges and outlook References Further reading

688 688 689 690 691 695 696 707

Bioreactors in tissue engineering: mimicking the microenvironment Ece Bayir, Mert Sahinler, M. Mert Celtikoglu and Aylin Sendemir 27.1 The role of bioreactors in tissue engineering 27.2 Bioreactor configurations 27.2.1 Stirred bioreactors 27.2.2 Wave bioreactors 27.2.3 Parallel-plate bioreactors and parallel-plate flow chamber bioreactors 27.2.4 Rotating wall vessel (reduced gravity) bioreactors 27.2.5 Strain bioreactors 27.2.6 Perfusion bioreactors 27.2.7 Hollow-fiber bioreactors 27.2.8 Microfluidic bioreactors 27.2.9 Combined systems 27.3 Cell-seeding techniques for bioreactors 27.4 Design considerations and future outlook 27.5 Conclusion References Further reading

709

Simulation of organ-on-a-chip systems Nenad Filipovic, Milica Nikolic and Tijana Sustersic 28.1 Introduction 28.1.1 General overview 28.1.2 Review of the lung and liver cell line models 28.2 Review of numerical solutions of developed models 28.2.1 Finite-difference method 28.2.2 Finite element modeling 28.3 Modeling of bioreactor for lung cells 28.4 Mathematical modeling of liver cells 28.5 Conclusion Acknowledgments References Further reading

753

Index

709 711 711 712 714 715 717 722 725 727 729 731 733 739 739 752

753 754 757 760 760 762 763 773 784 786 786 790 791

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List of contributors

Mohsen Akbari Laboratory for Innovations in Microengineering (LiME), Department of Mechanical Engineering, University of Victoria, Victoria, BC, Canada; Center for Advanced Materials and Related Technologies, University of Victoria, Victoria, BC, Canada Mark C. Allenby Institute of Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia Melis Asal Izmir Biomedicine and Genome Center, ˙Izmir, Turkey; Izmir International Biomedicine and Genome Institute, Dokuz Eylul University, ˙Izmir, Turkey Vincent Ball University of Strasbourg, Faculty of Dental Surgery, Strasbourg, France; French National Health Institute, Unit 1121, Strasbourg, France Ece Bayir Ege University Central Research Test and Analysis Laboratories Research and Application Center (EGE-MATAL), Izmir, Turkey Constantin Berger Tissue Engineering and Regenerative Medicine, University Hospital Wu¨rzburg, Wu¨rzburg, Germany Arindam Bit Department of Biomedical Engineering, National Institute of Technology, Raipur, India Esteban Brenet Inserm UMR 1121, Biomate´riaux et Bioinge´nierie, Strasbourg, France; Universite´ de Strasbourg, Strasbourg, France; Department of OtorhinoLaryngology, ENT Surgery, University Hospital of Reims, Reims, France Paul A. Cahill Vascular Biology & Therapeutics Group, School of Biotechnology, Dublin City University, Dublin, Ireland Laura E. Castellano Department of Chemical, Electronics and Biomedical Engineering, Science and Engineering Division, University of Guanajuato at Leo´n, Guanajuato, Mexico E. Ada Cavalcanti-Adam Max Planck Institute for Medical Research, Heidelberg, Germany

xx

List of contributors

Berivan Cecen Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States M. Mert Celtikoglu Department of Bioengineering, Ege University, Izmir, Turkey Sheyda Darouie Radboud University Medical Center, Radboud Institute for Molecular Life Sciences, Department of Dentistry-Regenerative Biomaterials, Philips van Leydenlaan 25, Nijmegen, The Netherlands Thomas D¨aullary Tissue Engineering and Regenerative Medicine, University Hospital Wu¨rzburg, Wu¨rzburg, Germany Christian Debry Department of Otorhino-Laryngology, ENT Surgery, University Hospital of Strasbourg, Strasbourg, France Jorge Delgado Department of Chemical, Electronics and Biomedical Engineering, Science and Engineering Division, University of Guanajuato at Leo´n, Guanajuato, Mexico Mathilde R. Desselle Institute of Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia Alireza Dolatshahi-Pirouz Radboud University Medical Center, Radboud Institute for Molecular Life Sciences, Department of Dentistry-Regenerative Biomaterials, Philips van Leydenlaan 25, Nijmegen, The Netherlands; Department of Health Technology, Center for Intestinal Absorption and Transport of Biopharmaceuticals, Technical University of Denmark, Lyngby, Kgs, Denmark Claire Ehlinger UMR-S 1121 Inserm, Biomaterials and Bioengineering, Strasbourg, France; Faculty of Dental Medicine, Strasbourg University, Strasbourg, France; Federation of Translational Medicine, Strasbourg, France Lea Fath Department of Otorhino-Laryngology, ENT Surgery, University Hospital of Strasbourg, Strasbourg, France; Inserm UMR 1121, Biomate´riaux et Bioinge´nierie, Strasbourg, France; Universite´ de Strasbourg, Strasbourg, France Milad Fathi-Achachelouei Department of Biomedical Engineering, Middle East Technical University, Ankara, Turkey Wenqian Feng Max Planck Institute for Medical Research, Heidelberg, Germany Christina Fey Tissue Engineering and Regenerative Medicine, University Hospital Wu¨rzburg, Wu¨rzburg, Germany; Translational Center Regenerative Therapies, Fraunhofer Institute for Silicate Research ISC, Wu¨rzburg, Germany

List of contributors

Agathe Figarol Department of Applied Engineering, Osaka University, Suita, Japan

xxi

Chemistry, Graduate School

of

Nenad Filipovic Faculty of Engineering, University of Kragujevac (FINK), Kragujevac, Serbia; Steinbeis Advanced Risk Technologies Institute doo Kragujevac (SARTIK), Kragujevac, Serbia; Bioengineering Research and Development Center (BioIRC), Kragujevac, Serbia Yannis-Nicolas Franc¸ois Laboratory of mass spectrometry of interactions and systems (LSMIS), CNRS UMR7140, University of Strasbourg, Strasbourg, France Rabah Gahoual Department of Chemical and Biological Technologies for Health (UTCBS), CNRS UMR8258 - Inserm U1022, Faculty of Pharmacy, Paris Descartes University, Paris, France Elena Garcı´a-Gareta Regenerative Biomaterials Group, RAFT Institute, Mount Vernon Hospital, Northwood, United Kingdom Michael Gasik School of Chemical Engineering, Aalto University Foundation, Espoo, Finland; School of Health Sciences, University of Eastern Piedmont, Novara, Italy Sinan Gu¨ven Izmir Biomedicine and Genome Center, ˙Izmir, Turkey; Izmir International Biomedicine and Genome Institute, Dokuz Eylul University, ˙Izmir, Turkey; Department of Medical Biology, Faculty of Medicine, Dokuz Eylul University, ˙Izmir, Turkey Shabir Hassan Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States Marcel Heinrich Section Targeted Therapeutics, Department of Biomaterials Science and Technology, Technical Medical Centre, University of Twente, Enschede, The Netherlands Pascal Houze´ Department of Chemical and Biological Technologies for Health (UTCBS), CNRS UMR8258 - Inserm U1022, Faculty of Pharmacy, Paris Descartes University, Paris, France; Biochemistry department, University Hospital Necker-Enfants Malades, Public assistance - Paris Hospitals (AP-HP), Paris, France Yasamin A. Jodat Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States

xxii

List of contributors

S. Jung Oral Biology Department, Faculty of Dentistry, University of Strasbourg, Strasbourg, France; INSERM UMR_S 1109 (Molecular ImmunoRheumatology unit), Strasbourg, France Dilek Keskin Department of Biomedical Engineering, Middle East Technical University, Ankara, Turkey; Center of Excellence in Biomaterials and Tissue Engineering (BIOMATEN), Middle East Technical University, Ankara, Turkey; Department of Engineering Sciences, Middle East Technical University, Ankara, Turkey Leyla Kocgozlu UMR-S 1121 Inserm, Biomaterials and Bioengineering, Strasbourg, France; Faculty of Dental Medicine, Strasbourg University, Strasbourg, France; Federation of Translational Medicine, Strasbourg, France; Mechanobiology Institute, National University of Singapore, Singapore, Singapore Anne Lehn Department of Pediatric Surgery, University Hospital of Strasbourg, Strasbourg, France; UMR DIATHEC, EA 7294, Translational Medicine Federation of Strasbourg (FMTS), University of Strasbourg, Strasbourg, France Emal Lesha Tufts University School of Medicine, Boston, MA, United States Fiona Louis Joint Research Laboratory (TOPPAN) for Advanced Cell Regulatory Chemistry, Graduate School of Engineering, Osaka University, Suita, Japan Elisa Maillard Strasbourg University, Strasbourg, France Emmanuel Martinod Assistance Publique - Hˆopitaux de Paris, Hˆopitaux Universitaires Paris Seine-Saint-Denis, Hˆopital Avicenne, Chirurgie Thoracique et Vasculaire, Universite´ Paris 13, France; Universite´ de Strasbourg, Strasbourg, France; Laboratory Hypoxia and the Lung INSERM UMR 1272, Universite´ Paris 13, France Michiya Matsusaki Department of Applied Chemistry, Graduate School of Engineering, Osaka University, Suita, Japan; Joint Research Laboratory (TOPPAN) for Advanced Cell Regulatory Chemistry, Graduate School of Engineering, Osaka University, Suita, Japan Garrett B. McGuinness Centre for Medical Engineering Research, School of Mechanical and Manufacturing Engineering, Dublin City University, Dublin, Ireland Birzabith Mendoza-Novelo Department of Chemical, Electronics and Biomedical Engineering, Science and Engineering Division, University of Guanajuato at Leo´n, Guanajuato, Mexico Marco Metzger Tissue Engineering and Regenerative Medicine, University Hospital Wu¨rzburg, Wu¨rzburg, Germany; Translational Center Regenerative Therapies, Fraunhofer Institute for Silicate Research ISC, Wu¨rzburg, Germany

List of contributors

xxiii

Florent Meyer Oral Biology Department, Faculty of Dentistry, University of Strasbourg, Strasbourg, France; INSERM, UMR_S1121, Biomaterials and Bioengineering, Strasbourg, France Nathalie Mignet Department of Chemical and Biological Technologies for Health (UTCBS), CNRS UMR8258 - Inserm U1022, Faculty of Pharmacy, Paris Descartes University, Paris, France Azadeh Mostafavi Department of Mechanical and Materials Engineering, University of Nebraska
Lincoln, Lincoln, NE, United States Milica Nikolic Faculty of Engineering, University of Kragujevac (FINK), Kragujevac, Serbia; Steinbeis Advanced Risk Technologies Institute doo Kragujevac (SARTIK), Kragujevac, Serbia; Bioengineering Research and Development Center (BioIRC), Kragujevac, Serbia Richard A. O’Connor Centre for Medical Engineering Research, School of Mechanical and Manufacturing Engineering, Dublin City University, Dublin, Ireland; Vascular Biology & Therapeutics Group, School of Biotechnology, Dublin City University, Dublin, Ireland Naomi C. Paxton Institute of Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia Patrina S.P. Poh Julius Wolff Institute, Charite´—Universit¨atsmedizin Berlin, Berlin, Germany Jai Prakash Section Targeted Therapeutics, Department of Biomaterials Science and Technology, Technical Medical Centre, University of Twente, Enschede, The Netherlands Jacob Quint Department of Mechanical and Materials Engineering, University of Nebraska
Lincoln, Lincoln, NE, United States Dana M. Radu Assistance Publique - Hˆopitaux de Paris, Hˆopitaux Universitaires Paris Seine-Saint-Denis, Hˆopital Avicenne, Chirurgie Thoracique et Vasculaire, Universite´ Paris 13, France N. Vijayakameswara Rao Department of Chemical Engineering, National Taiwan University of Science and Technology, Taipei Carina Russell Department of Mechanical and Materials Engineering, University of Nebraska
Lincoln, Lincoln, NE, United States Mert Sahinler Department of Bioengineering, Ege University, Izmir, Turkey

xxiv

List of contributors

Aylin Sendemir Department of Bioengineering, Ege University, Izmir, Turkey; Department of Biomedical Technologies, Ege University, Izmir, Turkey Amir Seyfoori Laboratory for Innovations in Microengineering (LiME), Department of Mechanical Engineering, University of Victoria, Victoria, BC, Canada Su Ryon Shin Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States Se´verine Sigrist Defymed S.A.S., Strasbourg, France Sudip Kumar Sinha Department of Metallurgical Engineering, National Institute of Technology, Raipur, India Tijana Sustersic Faculty of Engineering, University of Kragujevac (FINK), Kragujevac, Serbia; Steinbeis Advanced Risk Technologies Institute doo Kragujevac (SARTIK), Kragujevac, Serbia; Bioengineering Research and Development Center (BioIRC), Kragujevac, Serbia Ali Tamayol Department of Mechanical and Materials Engineering, University of Nebraska
Lincoln, Lincoln, NE, United States Aysen Tezcaner Department of Biomedical Engineering, Middle East Technical University, Ankara, Turkey; Center of Excellence in Biomaterials and Tissue Engineering (BIOMATEN), Middle East Technical University, Ankara, Turkey; Department of Engineering Sciences, Middle East Technical University, Ankara, Turkey Elif Vardar De´partement Femme-Me`re-Enfant (DFME), Centre Hospitalier Universitaire Vaudois (CHUV), Lausanne University Hospital, Lausanne, Switzerland Dominique Vautier UMR-S 1121 Inserm, Biomaterials and Bioengineering, Strasbourg, France; Faculty of Dental Medicine, Strasbourg University, Strasbourg, France; Federation of Translational Medicine, Strasbourg, France Arturo Vega-Gonza´lez Department of Chemical, Electronics and Biomedical Engineering, Science and Engineering Division, University of Guanajuato at Leo´n, Guanajuato, Mexico Nihal Engin Vrana SPARTHA Medical, Strasbourg, France; INSERM UMR 1121, Biomaterials and Bioengineering, Strasbourg, France

List of contributors

xxv

Cynthia S. Wong Institute of Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia Maria A. Woodruff Institute of Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia Daniela Zdzieblo Tissue Engineering and Regenerative Medicine, University Hospital Wu¨rzburg, Wu¨rzburg, Germany; Translational Center Regenerative Therapies, Fraunhofer Institute for Silicate Research ISC, Wu¨rzburg, Germany Yu Shrike Zhang Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States

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Preface

The chapters in the book are aimed to provide the reader with a more in-depth information about the basis of biomaterial/biological system interactions and properties in the context of their use in tissue/organ regeneration and modeling. The book is divided into three sections. The first section covers “the properties and forms of biomaterials” with a part on cell types and their properties relevant to regenerative medicine, whereas the second section focuses on “the use of biomaterials in organ-specific applications.” The last section elaborates on the “emerging and enabling technologies” related to biomaterials with a particular focus on their relevance to tissue and organ regeneration (be it for therapeutic, modeling, or other purposes). The second section is more application oriented and the chapters in this part cover the specific requirements of a given organ, the available designs in the literature used for their replacement or modeling and the current trend and future outlook starting with cardiovascular tissues, small intestine, and pancreas. This is followed by the role of tissue engineering in a specific and particularly problematic kind of wound healing, diabetic wound healing. A chapter focuses on bone tissue engineering with an emphasis on the historical use of bone morphogenetic proteins and the problems and risk that entailed their use. Whereas others, take us into the developments in miniaturized tissue/physiological zone models with the examples of adipose tissue and blood
brain barrier. After the coverage of two important organs/systems (renal systems and respiratory), the section concludes with the description and uses of a regenerative tool that has become common in clinical practice, platelet enriched plasma. The last section of the book is more technology oriented starting with the description of nanocomposite hydrogels and systems based on carbon-based biomaterials which are being actively used in biorobotics applications previously mentioned. In another single material-oriented chapter, the use of hyaluronic acid in tissue engineering and regenerative medicine with an attention to delivery systems is covered. There is a chapter that focuses on the incorporation of the microfluidic systems in biomaterials applications, while another covers an important but generally difficult to analyze aspect of biomaterial characterization, the biomechanical properties and testing of biomaterials. A chapter provides an overview of all in vitro model systems, while a subsequent one, concentrates on the biomaterials used to develop such systems. The penultimate chapter covers the main means in the biomaterial fields to apply the dynamic physiological conditions to

xxviii

Preface

tissue-engineering constructs, the types and properties of bioreactors. Simulation of organ-on-a-chip systems, concludes the book with description of the use of simulations in biomaterials field with some specific examples. We hope that the book provides an overall view of the field and will be useful for the researchers in the field and the students.

Acknowledgment

The editors would like to acknowledge the funding from the European Union’s Horizon 2020 research and innovation programme under grant agreement No 760921 (PANBioRA).

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Section 1 Properties and forms of biomaterials

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Introduction to biomaterials for tissue/organ regeneration

1

Nihal Engin Vrana1,2 1 SPARTHA Medical, Strasbourg, France, 2INSERM UMR 1121, Biomaterials and Bioengineering, Strasbourg, France

1.1

Introduction

As any broad and end point
related term, “Biomaterials” is hard to define as it encompasses a large group of materials with little similarity in their physicochemical properties. This stems from the historical fact that the conventional definition of “Biomaterials” refers not to their inherent properties but more to their intended use. In other words, “Biomaterials” as a term is generally used for any kind of material which is in contact with a biological entity in an intended way of interaction. If we narrow this definition even further (as its understanding was narrower in the beginning), one of the most common early uses was the replacement of a physiological function in human body as a therapeutic solution (i.e., the biomedical applications of biomaterials). This is generally done in the form of implants or implantable medical devices in a broader sense. These devices can be as simple as a stent to keep a vessel open or a dental implant to replace the mechanical function of teeth, which are made of one or few biomaterials. But in many cases, these devices are complex equipment such as pacemakers to keep an organ continue its function, which has many components and made of different materials [in some cases, some of them might be even toxic and need to be isolated from the rest of the body if they are crucial for the function of the overall system (such as batteries or sensor components, e.g.)]. But regardless of the complexity of the device, in the case of the replacement of a single function (such as load bearing, tissue filling, and signal transmission), the interface between the medical device and the host tissue is very well defined. The primary function is generally limited to one interaction (keeping the lumen open for the stents and sending the necessary electrical signals for the control of the beating of the heart for the pacemakers, e.g.) and any other interaction is considered detrimental. This has led to an initial preference for bioinertness where the inherent properties of the bioinert biomaterials evade any unwanted secondary interactions with the body, such as inflammation or blood coagulation. This approach kept the interaction of the medical device with the body in the predefined conditions while putting a limit to undesirable secondary effects. However, now the biomaterial field has matured as such approaches have proven to be useful and successful; they have become common tools in healthcare. Now, there are several implantable devices such as knee implants, breast implants, and cochlear implants that are implanted in Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00001-5 © 2020 Elsevier Ltd. All rights reserved.

4

Biomaterials for Organ and Tissue Regeneration

the ranges of several hundred thousand to millions each year. As a result, the ambitions had surpassed the initial goals where the aim moved from just supporting or helping a tissue/organ to completely replacing them. Maybe in the near future, designing nonexistent tissues/organs with capabilities surpassing the natural tissues will also be possible. Thus the bioinert approach to biomaterials stopped to be relevant. Today, the researchers in the biomaterial field design structures that can interact with the host tissues or external stimuli in a complex manner with a point of view of full integration. For example, recently it has been shown that by using a bilayer skin substitute containing cellular components, a prohealing response and keratinocyte activation can be obtained in the clinical case of chronic venous ulcers [1]. However, for the time point, this ambition has proven to be more difficult than the initial optimism and enthusiasm suggested [2] as the number of available, functional tissue-engineering solutions are limited to skin, cartilage, and some of the hollow organs such as bladder. Thus there is still room for improvement and innovation for biomaterials that can direct and facilitate the regeneration and/or replacement of more complex organs. The motivation for this book was to put together the different aspects of biomaterial-related technologies developed for tissue/organ regeneration while providing a strong basis for understanding biomaterial properties and uses. The requirements and functions of each tissue necessitate the use of specific kind of biomaterials in specific forms and properties. Beyond the tissue-related design constraints, the biomaterial researchers also need to design their systems for more general requirements such as anti-immunogenicity, blood compatibility, and controllable resistance to the degradation and failure due to the contact with body fluids. While the rest of the book takes a more historical approach and aims at fundamental understanding pertaining to either biomaterial properties or their applications, this chapter turns its attention more on the very recent developments to provide the reader with an overview of the current activities in the field.

1.2

Many facets of new biomaterials: new naturally sourced biomaterials, new synthetic biomaterials, materiomics, metabiomaterials

In the early days of biomaterials, whenever the need for interaction with the cells and tissues was a priority, the most commonly used materials where natural biomaterials sourced from animals, plants, and bacteria. Some of these materials were attractive as they were readily available (such as cellulose or alginate) or have biological properties that are necessary for a specific application (such as the promotion of cell adhesion by collagen/gelatin or inherent antimicrobial properties of chitin/chitosan). Today, the greatest source of new materials remains to be the nature. The oven of evolution has churned myriads of materials with very interesting properties and some of them still are waiting to be found. These materials are not only limited to the primary materials as mentioned before but also include

Introduction to biomaterials for tissue/organ regeneration

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enzymes that can cross-link biomaterials in a highly efficacious manner or degrade them in a specific way that can enable controlled delivery or 4D remodeling. Meanwhile, with the advents in genetic engineering, the ability to produce polypeptides recombinantly, advances in supramolecular chemistry and click chemistry have opened the door for designed, naturally sourced biomaterials which not only harness the intrinsic properties of natural biomaterials but also can add to these properties by linking different domains with known functions in the same polypeptide chain (or as pendant groups) or creating hybrid materials, in a similar manner to glycosaminoglycans (GAGs), which demonstrate the relevant properties of proteins, lipids, and sugars in a design-specific manner. One recent example of such systems is glyconucleolipids, which are synthetic amphiphilic biomaterials based on nucleic acid and lipid chemistry and can be used for drug delivery, biosensor coating, and also for biobattery applications [3], which can one day lead to artificial tissues with bioelectronic components for whose energy requirements there are incorporated biobatteries in the engineered tissue. One of the most commonly used synthetic polypeptide-based biomaterials is elastin-like recombinamers (ELRs). These molecules are based on elastin, a component of extracellular matrix (ECM) that provides the tissues with elasticity due to its disordered structure, which can withstand considerable levels of strain and result in a recoiling behavior upon removal of the stress. When the structure of the precursor of elastin, tropoelastin, was first investigated, a recurring pentapeptide sequence— VPGVG—has been discovered. As mentioned before, with the recombinant technologies, it has become feasible to produce such simple functional units without any secondary or tertiary structures in bacteria such as Escherichia coli [4]. This has given rise to ELRs, which are now being widely used in drug/gene delivery and tissue engineering in the form of hydrogels or porous scaffolds. As the basis of ELRs is quite simple, the potential number of derivatives is endless, and they can be turned into 3D systems either by using standard chemical cross-linking methods or using modified ELRs suitable for click chemistry or responsive to enzymatic cross-linking. ELRs provide a good example of finding out the biological building block responsible for a desired property in natural materials and expanding on this knowledge for designing new materials and advanced biosystems with new functions. A recent example is the design of an injectable system, which is composed of a specifically encoded polypeptide sequence based on ELR and polyaniline that can form porous structures of controllable dimensions in vivo once in contact with the body temperature [5]. This behavior stems from their partially ordered nature or more precisely the inherent disordered nature of its ELR component, which allows the sequence-controlled system to tap into the natural emergent properties of the system by the means of interactions between ordered and disordered domains. Such approaches (where the design of the biomaterial results in dynamic changes in the target area due to the microenvironmental properties of the implantation/injection zone) can be fine-tuned for the needs of a given tissue, for example, silk-ELR diblocks with controlled fibrillar assembly properties. These molecules can be used where the presence of anisotropic fibers is paramount for the function of the tissue such as ligament, muscle, or cornea [6].

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Biomaterials for Organ and Tissue Regeneration

A one step further than the tailored design and development polypeptide sequences is full-scale molecular bionics [7], that is, the design and development of a complex system that can mimic the unique properties of the biological systems such as dynamic responsiveness to stimuli, self-assembly, and highly selective temporally controlled interactions. In the first step, this involves using supramolecular interactions and interactions between nanoscale particles to mimic the construction of biological systems such as ECM (for mimicking the self-assembly of collagen chains into several degrees of order and their supramolecular interactions with the other components of ECM such as proteins and GAGs). Such complex, active 3D structures are what need to be created for providing suitable niches for cells in tissue-engineering applications. There are many methodologies for achieving this such as layer-by-layer (LbL) methods using electrostatic interactions and entropydriven assembly of charged natural or synthetic polymers. As LbL is an iterative process where, in each step, the material component can be changed, the level of composite properties achievable and the level of compositional control in the direction normal to the substrate is extremely high [8], which had been harnessed widely in fuel cell applications but also in biomedical applications particularly for cell/surface interaction control [9]. Our group has been developing self-antimicrobial LbL assemblies, which is only possible under narrow material coupling conditions [10]. Another potential way of molecular bionics is the self-assembling polypeptide sequences (such as low molecular weight hydrogelators whose self-assembly can be precisely controlled either in solution or on surfaces [11]) and amphiphilic molecules inspired by their natural counterparts which form fibrillar structures based on supramolecular interactions such as hydrogen bonding and π
π stacking. Such assemblies can be either used for forming bulk materials via supramolecular interactions. Another potential way is incorporating them into covalently linked structures such as a protein backbone (a gelatin-based film or hydrogel, e.g.) and introducing additional supramolecular assemblies into this backbone to achieve multifunctionality [12]. Another commonly used method is host
guest interactions, for example, based on cyclodextrin (a cup-shaped polysaccharide structure) for assembling self-healing hydrogels with controlled properties such as mechanical stiffness or transparency [13]. Self-healing is a particularly important property in applications where there are considerable stress/strain conditions involved, which can ensure the long-term functionality of the scaffold. Beyond the natural or natural precursor-based polymers, the second option is fully synthetic polymers. The initial advances in the development of biodegradable synthetic polymers for tissue/organ regeneration had toned down over time; as the rate of degradation of these materials is mostly not in-line with the timeline of tissue generation, their degradation by-products can be harmful and inflammationinducing and their overall capacity to induce specific functional responses from cells is generally limited. This has led to a different approach, which focuses on the available biomaterials in the form of decellularized tissues that keep the intricate structure of ECM while avoiding the immune reactions due to the removal of the cellular component which is the driving force of the immune response [14]. For this end, any tissue can be used to obtain decellularized structures (the decellularization

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process is carried out with many methods such as detergent treatments, enzymatic treatments, freeze-thawing, and chemical treatments), and recently the decellularization of readily made engineered tissues as off-the-shelf tailorable templates has also been developed [15]. However, there is one inherent problem with decellularized tissues from an engineering point of view, which is that the process starts with materials with predefined properties and each step deteriorates these properties no matter how well it is controlled; thus the level of control on the properties of the final system is limited. So, there is still a niche for synthetic polymers in tissue engineering and regenerative medicine (beyond their superior controlled delivery capacities); provided that their biological properties can be improved. In the meantime, the level of complexity in synthetic polymers increased also and now it is possible to synthesize polymers with inherent biological functions and also a high level of physiologically relevant stimuli responsiveness (pH, light, O2, H2O2, temperature, etc.). However, as our capacity to design and produce new biomaterials increases, the task of proper characterization of them becomes daunting. One way to circumvent this problem is the application of high throughput methodologies as used in -omics approaches to the material discovery and looking for hits with desired functions rather than trying to design biomaterials bottom-up with a hypothesis-driven approach. In fact, one of the driving forces in the increased use of decellularized templates is the sudden jump in the level of information available about their individual components, even very minute ones, obtained by -omics approaches and other precision analytical techniques that had become readily available recently. With the breakthroughs achieved by the introduction of genomics, proteomics, and metabolomics in their respective fields, there are reasonable expectations that materiomics approaches can also revolutionize the field of biomaterials [16]. One way of approaching materiomics is the use of omics techniques (particularly transcriptomics and proteomics techniques to understand the effect of a given material on cell gene expression and protein secretion patterns, respectively) to gain initial insights about the effects of a given set of materials and then applying modifications that would steer the final design toward the desired properties (be it the chondrogenic differentiation of stem cells to preferential apoptosis of cancer cells or an angiogenic effect). Then, the hit materials with the desired properties can be bulk produced and can be used for the development of 3D structures such a stimuliresponsive hydrogels [17]. Although nonhypothesis-driven, serendipitous biomaterial discovery can be criticized for the lack of understanding of the underlying mechanisms of the obtained results, its capacity to facilitate the pinpointing of biomaterials with exceptional properties cannot be understated and this branch of biomaterials research can be expected to grow exponentially. With all the tools described previously, a new step taken, in general, in materials science is to harness our capabilities to develop intricate material structures to obtain new materials with unique behavior either as a proof of concept or for specific applications. A metamaterial, by definition, is a material that has properties that are different from those observed in bulk materials. They are obtained by precise engineering of structural properties, which leads to the expected property that is beyond the capacities of the material of origin (although they are not necessarily

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Biomaterials for Organ and Tissue Regeneration

single material constructs). The resulting properties such as negative refractive index or negative Poisson’s ratio have practical applications in many fields (such as Ghost in the Shell
style invisibility cloaks) and biomaterial applications can also benefit from metamaterials, although they are not as commonplace as in the applications such as photonics. One recent example is the calibration of fatigue behavior (due to topological properties of the metamaterial) of titanium/cobalt-chromium, tantalum-based additively manufactured (via selective laser melting) metamaterials for orthopedic implant applications [18]. Naturally, certain applications will require the harnessing of all types of materials described previously which will require combinatorial biomaterial systems. Combinatorial systems are based on using mathematical models for the detection of local extrema that correspond to an optimal performance while using multiple parameters without relying on extravagantly high number of experiments. Methodologies developed for testing/experiments with rare resources for optimal understanding of the effect of the individual and combination of parameters on a given set of outcomes can be used for obtaining composite biomaterial structures optimized for multiple outputs, which is generally the requirement for tissue/organ regeneration.

1.3

Off-shoot technologies linked to biomaterials and tissue engineering: biorobotics, bioinks, and bioprinting

This high level of activity in biomaterials science had other benefits also. Tissueengineering methodologies, which basically boil down to creating 3D architectures that can be inhabited by the cells, might have applications beyond the realm of healthcare. From an engineering point of view, tissues and organs are highly specialized and efficient devices with multiple functions concentrated in relatively small volumes with highly efficient use of energy and with very strong stimuli responsiveness and also self-healing properties. Also cells as entities with cell type
specific responses to their environment are very strong tools for the development of advanced biosystems [19]. The utilization of stimuli responsive engineered tissues for robotic applications, which is a sub-branch of biorobotics, is based on these premises. For example, muscle cells that are responsive to electrical signals and fibroblasts, which are not can be micropatterned to created “living diodes” [20]. The cell encapsulation technologies developed for soft tissue engineering can also be harnessed for biorobotics applications as the contractility of muscle cells (particularly heart muscle cells) together with their responsiveness to the electrical signals (which can be amplified by using conductive hydrogels either based on conductive polymers or composites using ingredients such as graphene oxide or carbon nanotubes) can be used to created living actuators [21]. These approaches can revert

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back to help tissue-engineering applications also, as a microrobotic, steerable system which contains a suitable stem cell niche has been recently reported for cell delivery for tissue regeneration [22]. There is also a huge potential for biomaterials to adopt well-established technologies and techniques, for example, the use of electrospinning which has its roots in textile industry for nanofibrous scaffold production for tissue engineering is a good example of such technological transfers. One of the biggest problems in the tissue engineering is that of the volume; developing engineered structures that have the means to distribute the nutrients and oxygen and get rid of waste to stay alive (which is handled by the cardiovascular system in mammalians). The standard methods of scaffold formation rely on bulk materials such as additively manufactured structures, nanofibrillar thick structures produced by methods such as electrospinning, or porous structures such as foams. The amount of material available in such structures and tortuosity of the pores (when they are available) created serious diffusion limitations for nutrients, waste, and gases. One of the natural solutions to such problems is folding. By folding structures, voluminous systems can be obtained that in fact contain very little material. The well-established paper-folding techniques, generally known as origami (or kirigami if the substrate use has cuts or been cut at a later stage), are not only suitable in terms of biomaterials but can also be applied for applications in electronics, sensors, and energy applications [23]. Moreover, origami/kirigami techniques also allow the development of metamaterials, as described previously, particularly from mechanical property control and stimuli responsiveness point of view. These methods have been ingeniously applied in tissue-engineering scaffold development as folding is an efficient way of obtaining structures with lumens and in its essence, paper is a highly processed biomaterial. By using a paper scaffold (its surface modified with poly-L-lysine) folded with origami methods, laden with a hydrogel (alginate) a scaffold that was used for tracheal reconstruction in rabbits was constructed [24]. This technique is a very good example of multiscale engineering that is required for tissue engineering and regenerative medicine; the origami techniques were used for the macrostructure, while the presence of the hydrogel component was achieved by harnessing the LbL techniques at the nanoscale as described before. Although due to the additional layer of complexity required for designing origami procedures for complex architectures the examples of origami-based tissue-engineering templates are limited at the moment, the technique is gaining traction specifically in the form of advanced cell culture substrates [25] which can be further rendered functional (such as being rendered conductive by the addition of reduced graphene oxide) [26]. One manner of controlling macroscale behavior is to design microlevel architectures that would lead to highly ordered 3D structures. In tissue/organ generation, the use of porous structures as scaffolds is very common, and one way of achieving a higher order of control in such systems is using templating methodologies, which provides better structure property control that is critical for biological applications such as pore interconnectivity [27]. For example, Dehli et al. developed hydrogel foams based on gelatin methacryloyl using liquid foam templating methodologies

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Biomaterials for Organ and Tissue Regeneration

that result in highly ordered porous structures with a very precise control over pore morphology, distribution, and size [28]. Such methods highly benefit from the recent advances in microfluidics and the increased availability of the required instruments. A more active are for achieving such microscale control is 3D bioprinting [29] and more recently 4D bioprinting, which refers to the printing of cellladen stimuli responsive structures that can morph into predesigned shapes after the printing process and when required [30]. As the instrumentation for bioprinting has become readily available at a standardized manner over the last 10 years (both at simple instruments for research purposes or more therapy oriented, high-end devices that can be operated in GMP conditions), the focus turned on the development of the optimized building blocks, that is, materials with good printability generally referred to as bioinks [31]. It is important to make the distinction between bioinks, a nomenclature is recently suggested [32]. Bioinks need to have a cellular component, whereas without cells the inks are “biomaterial inks” used for creating scaffolds. However, definitions will never cover all the potential ways of achieving goals in engineering. In a recent ingenious “back to the basics” approach, Wang et al. used ice as a sacrificial scaffold for printing complex vascular structures, as a single molecule with very well-defined physicochemical behavior, ice is a more suitable sacrificial scaffold than conventional thermoresponsive polymeric sacrificial scaffolds or decross-linking/dissolution-based sacrificial layers as the printing conditions are more controlled once the system is in place for printing free-standing structures (Fig. 1.1) [33]. A water feed is kept in liquid form via insulation in a below freezing temperature platform that enabled the 3D printing of ice (or more precisely controlled freezing of layers of water that leads to 3D structures) which can be then dip coated with any given material and lyophilized to give freestanding, complex 3D architectures. As additive manufacturing is a prolific area which has firmly anchored itself in many sectors, the developments in that field will surely trickle down to bioprinting be it 3D or 4D and result in more sophisticated and functional systems.

1.4

Biomaterial risk assessment

The exponential growth in the number of available biomaterials creates a “problem of the rich”: how could we assess the potential risks pertaining to each of these materials? The utilization of the established biocompatibility testing methods would soon become intractable and also there are significant discrepancies between the biocompatibility assessment (which is highly reliant on animal testing) and the current will decrease the animal tests due to ethical concerns, cost-related considerations, and also the questions related to their relevancy for predicting the behavior in humans. Moreover, the way these tests have been designed only provides a “Yes/ No” answer to the risk-related questions and these tests do not add up to a better understanding of the differences between new materials. Thus there is a growing

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Figure 1.1 3D printed ice structures as sacrificial scaffolds for the development of vessellike structures with different biomaterials. Source: Reprinted with permission from Wang R, Ozsvar J, Aghaei-Ghareh-Bolagh B, Hiob MA, Mithieux SM, Weiss AS. Freestanding hierarchical vascular structures engineered from ice. Biomaterials 2019;192:334
45.

need in both design and the development of new testing systems that will provide more pertinent and rich information regarding the potential risks of biomaterials and also means of analyzing these data and corroborating it with the historical information to achieve an overall view of the relative risks of all available biomaterials. One of the examples of efforts for improved biomaterial risk assessment is Horizon 2020 PANBioRA project (www.panbiora.eu). PANBioRA aims to develop an integrated and automated instrument that can test biomaterials from antibody scale (via immunoprofiling) to miliscale (by using connected organs-on-a-chip) (Fig. 1.2) [34]. A recent example of how this kind of testing systems might look like is the “Foreign Body Response-on-a-chip” platform, which is a perfusable system where biomaterials are separated from the flow, containing immune cells, with a porous membrane covered with a confluent vascular endothelial layer; it creates a mimic of the extravasation of immune cells to the implantation site in vivo [35]. Other efforts involve the use of skin tissue mimic for biomaterial risk assessment [36]. Multiparametric electrochemical sensor signals have also been shown to enable the analysis of the ex vivo response to biomaterials [37]. In biomaterial risk assessment, both the type of the biomaterial [38] and form of the biomaterial (such as nanoparticles given that they are widely used in tissue engineering particularly for cytokine and growth factor delivery [39]) need to be taken into account [40]. While the risk assessment of new biomaterials themselves are a highly active research field, the use of biomaterials for the risk assessment of pharmacological

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Biomaterials for Organ and Tissue Regeneration

Figure 1.2 PANBioRA personalized or generalized biomaterials risk-assessment system under development based on immunoprofiling, real-time microscopy, electrochemical, and antibody-based biosensor read outs for immune cells and organ-on-a-chip systems supplemented with simulations. Source: Reprinted with permission from Ungemach M, Doll T, Vrana NE. How to predict adverse immune reactions to implantable biomaterials? Eur J Immunol 2019;49(4):517
20.

and cosmetics products is also evolving. Following the general public sentiment against the animal tests and the recent legislation changes particularly in Europe that ban animal testing for cosmetics and firmly establish the utilization of 3Rs (replacement, reduction, refinement) in all animal testing, brought a significant interest in biomaterial-based 3D cell culture platforms as a potential replacement. For many uses, including pharmacology (new drug discovery), toxicology (potential effects of new chemical entities), cosmetics (animal free, reliable testing of new products), disease modeling (better understanding of the mechanisms of new and rare diseases), personalization, the need is to have in vitro models which are reliable, sophisticated, and readily available. One axis of research in in vitro tissue organ models is the development of cellonly systems. These are defined as spheroids and organoids. Spheroids are spherical or quasispherical cellular aggregates formed of one or multiple cell types. It is an approximation of the 3D organization of tissues and it has been shown that in these configurations certain cell types, such as hepatocytes, demonstrate more physiological behavior. Spheroids form in the conditions where the propensity of the cells to attach a given surface (such as a low adhesion cell culture plate) is less than them to form clusters, which in turn turns into geometrical aggregates under the biophysical cues (such as size-defined microwells to obtain reproducible spheroid size). However, the control over the tissue architecture is rather poor with spheroid

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method and for finer in vitro models the attention was generally on organoids. In organoid approach, starting from stem cells and via mimicking developmental stages either via biochemical and biophysical gradients, miniaturized organ-like structure which recapitulates several components of a given tissue and organ can be obtained. However, what is gained in complexity and fidelity is lost in the sense of ease of production, cost, and availability. The utilization of biomaterial-based matrices can provide artificial organ settings where the production of the model is cheaper, faster, and has more architectural similarity to the target organ. The current challenge in in vitro systems is the incorporation of as many tissue components as possible in the final system and keep it alive for a long term (so that they can become relevant microphysiological systems). Also many tissues have multiple concomitant functions (such as epithelium which not only acts as a physical barrier but also has secretory, nutrient, and waste transport and sensation-related roles). In addition to these multiple functions, another aspect is the connectivity of the artificial tissue models. In our body, nearly all tissues are in contact with the circulation [41], have connections with the immune system, and also contain immune cells and innervated and being controlled by the nervous system. Thus in order to claim a strong fidelity to in vivo conditions, on-chip model systems should contain at least rudimentary components that mimic these elements [42]. For example, in a recent example of a complex in vitro skin model system, Vidal et al. used collagen-silk composite systems to culture human skin cells with macrophages (immune component) and nerve cells (for monitoring innervation) (Fig. 1.3) [43]. The system contains hypodermis using patient lipoaspirate, dermis with skin fibroblasts, and epidermis with keratinocytes. Although it takes 42 days to develop the actual model, the level of tissue complexity achieved provides a demonstration of the level of fidelity of the next generation in vitro models. Incorporating system level components is one level of complexity for the miniaturized models. The other level of complexity comes from the incorporation of dynamic components in the form of liquid flow, stress/strain application, and also particularly for the skin and respiratory epithelium, the establishment of an interface with air; these can be achieved by multicompartment on chip systems, that can provide polarization of the 3D culture systems with multiple options for perfusion, air/liquid culturing and also by the ability to introduce new cell types automatically during the temporal maturation of the artificial tissue. A recent system by Sriram et al., using a fibrin matrix
based culture system embedded on a chip that can perfuse the system during the maturation and then provide a direct contact with the air for keratinocytes for achieving keratinization of the upper layer, provides a barrier function performance similar to that of human skin, making it suitable as a testing system for the skin penetration of materials [44]. Another aspect in the modeling of barrier tissues is their topographical features that are paramount to their functions; most of the epithelial tissues have complex topographical structures such as crypt
villus of gastrointestinal epithelium that will be crucial to mimic as they are tightly linked with their functions. Microfabrication methodologies that have been well established in tissue-engineering field are being widely used in the introduction of such properties to the on-chip systems [45].

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Biomaterials for Organ and Tissue Regeneration

Figure 1.3 An in vitro immunocompetent (containing macrophages) and innervated skin model. Source: Reprinted with permission from Vidal SEL, Tamamoto KA, Nguyen H, Abbott RD, Cairns DM, Kaplan DL. 3D biomaterial matrix to support long term, full thickness, immunocompetent human skin equivalents with nervous system components. Biomaterials 2019;198:194
203.

An important future aspect of organ models is going one step beyond and also incorporating the tissue-specific microbiota into the models, which is currently being implemented by different groups [46]. Lastly, one of the bottlenecks in the biomaterials field is its highly empirical nature and the limitations to the experimental capacities due to the cost or scarce resources related to the experiments. Also most of the expected behavior in the case of regenerative medicine is mid- to long term, which also puts breaks on the speed of analysis and evaluation. One potential way of circumventing this problem is to rely more on simulations and mathematical models which are not only can provide a cheaper, accessible, and highly modular ways of biomaterial assessment and development [47] but also might provide insights that could have been lost in the data due to a specific focus on outcomes. Moreover, as in many previous cases, exploitation of the big data methodologies developed in other fields for a metaanalysis of the overall field of biomaterials to gain mechanistic insights from the shear amount of research data available over the course of last 40
50 years is another venue that needs to be pursued to push the field forward.

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15

Conclusion

The research and development in the biomaterial field seemingly still is in its exponential growth phase with a constant flow of novel biomaterials and expanded applications of the technologies developed originally for healthcare. Even though the rate of success in the replacement of tissues and organs had not been at the levels expected and can be considered disappointing in other areas such as controlled delivery and in vitro models biomaterials have clearly become a central component. By continuing importing the advances in other fields, our field can continue to thrive and can provide solutions to unmet clinical needs and new technologies with far-reaching potentials.

Acknowledgment This project has received funding from the European Union’s Horizon 2020 research and innovation program under grant agreement no. 760921 (PANBioRA).

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14. [26] Li J, Liu X, Tomaskovic-Crook E, Crook JM, Wallace GG. Smart graphene-cellulose paper for 2D or 3D “origami-inspired” human stem cell support and differentiation. Colloids Surf, B: Biointerfaces 2019;176:87
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[29] Tamay DG, Usal TD, Alagoz AS, Yucel D, Hasirci N, Hasirci V. 3D and 4D printing of polymers for tissue engineering applications. Front Bioeng Biotechnol 2019;7:164. [30] Miri AK, Khalilpour A, Cecen B, Maharjan S, Shin SR, Khademhosseini A. Multiscale bioprinting of vascularized models. Biomaterials 2019;198:204
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Physicochemical properties of biomaterials

2

Vincent Ball1,2 1 University of Strasbourg, Faculty of Dental Surgery, Strasbourg, France, 2French National Health Institute, Unit 1121, Strasbourg, France

2.1

Introduction

As all materials, biomaterials display some physicochemical properties that have to suit for the application they are intended for. Those properties are simultaneously bulk properties of the material and surface properties, because the chosen biomaterial has its own integrity and interacts with organs through its surface. The most important physicochemical properties of a material are its mechanical properties, its stability against corrosion/degradation, and its electrical
optical properties. These kinds of properties are intimately linked to the composition and the structure of the material. In the case of biomaterials the electrical
optical properties are often of minor significance, with some exceptions, however. For instance, a tooth substitute has to be white, and a contact lens has to be transparent. Even if the history of biomaterials goes back to the ancient Egyptians, a rational investigation of the structure
properties relationship for biomaterials started only roughly about 60 years ago benefiting from the progress of high-performance materials developed for the aerospace industry. In the case of biomaterials the materials requirement extends those of a classical material, because biocompatibility is required [1]. In addition, the selected biomaterial for a given application has to adhere to its environmental tissue. This chapter will hence be organized in the following subsections: 1. 2. 3. 4.

bulk properties of biomaterials surface properties of biomaterials properties of biomimetic materials real-time monitoring of an implanted biomaterial and personalized implants

For none of these parts an exhaustive description will be given, for obvious reasons of space. Instead, the general aim of this chapter is to show that the physicochemical properties of a biomaterial need to respond to specific requirements of the target tissue/organ. The technical response to such requirements can only be approached by a multiparametric optimization approach. Here as in material science, the establishment of “Ashby diagrams” [2], relationships between different materials parameters for different materials, helps in the design of the optimal biomaterial for a specific application [3]. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00002-7 © 2020 Elsevier Ltd. All rights reserved.

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The investigation of biomaterials structure
properties relationship is not yet perfectly understood, and in addition, there is a strong dependence of the physiological state of the patient to be implanted on the final biological properties (good tissue integration, absence of bacterial infection) independent of the intrinsic material properties. This is in a sense also true for “regular” engineering materials: the aging of a bridge will indeed be somewhat different in a windy country than in a country where wind is rarely blowing, the orientation of the bridge with respect to the main wind orientation also plays a role, etc. However, in the case of biomaterials the intrinsic-specific variability is much stronger: every patient has its own genetic, epigenetic, and immunological background. To overcome and to cope with such random variability, the design of internal monitoring systems of biomaterials in their biological environment is of great promise. We will finish this chapter by describing such promising perspectives.

2.2

Bulk properties of biomaterials

The full description of materials used as biomaterials is given in this chapter of the present book with special emphasis on materials used for tissue
organ regeneration. Metals are mostly used for the design of hip prosthesis, dental implants, and so on. But for the design of tissue engineering scaffolds, polymer-based materials and polymer composites are playing a central role. Ceramics are also important materials in hip prosthesis and for dental use.

2.2.1 Shape and size control It appears obvious, but it is worth recalling such an evidence, that the shape of a biomaterial, is of tremendous importance [4]. The different parts of a hip prosthesis have a given shape and size, to be suitable to the local environment of the implant, in the case of the patient’s femur. Similarly, the dental implant and the contact lens have a given size and shape. These properties are now easily and reproducibly controlled by modern manufacturing techniques. 3D-printing techniques become now efficient to produce biomaterials, essentially scaffolds, personalized for the need of a given patient [5]. The size and shape of the required material is defined after medical imaging (computed tomography for instance) of the zone being subjected to host the desired biomaterial.

2.2.2 Mechanical properties The mechanical properties such as the tensile yield stress, the Young modulus, the fatigue strength and toughness of titanium, and its main alloys used as biomaterials have been tabulated [6]. Many different experimental methods for the characterization of dental materials are available and well described in the literature [7]. Typical shapes of single tensile tests for the different classes of biomaterials are shown in Fig. 2.1.

Physicochemical properties of biomaterials

21

Yield strength Stress (Pa)

Tensile strength

Brittle material F Ductile material

F

Slope=Young’s modulus

Strain (%)

Figure 2.1 Schematic representation of the tensile stress/strain behavior of solids presenting either a brittle behavior (red curve, typical of ceramics) or a ductile behavior (green curve, typical of metals). Some typical material properties such as the Young modulus (the slope of the linear part of the stress
strain curve, corresponding to elastic deformation), the yield strength (the maximal stress the materials is able to withstand in the elastic deformation regime which is defined by the intersection point of a straight line parallel to the experimental linear elasticity domain starting at 0.2% strain, green dashed line), and the tensile strength (the maximal stress the material is able to withstand before rupture) are indicated in blue. Point F corresponds to the materials’ failure in both cases. The area between the stress
strain curves and the horizontal axis provides the toughness of materials, the energy required to fracture the material.

The plastic deformation of metals corresponds mainly to the migration of dislocations present in the metal, and engineering of the metal architecture through thermal treatment can allow to extend the plastic deformation domain. Of course, fatigue tests aimed to evaluate (in an accelerated way) the long-term stability of a material under current cyclic loads are mandatory for prosthetic materials. Such a test implies repetitive solicitation of the material up to stress level corresponding to a defined fraction of the yield stress followed by returning to a zero-stress level. The number of loading
unloading cycles before material rupture is measured. The resistance to wear conditions is also mandatory for a biomaterial as a whole: the appearance of abrasion debris may induce cytotoxicity even if the implanted biomaterial is biocompatible as a bulk material [8]. More information about the mechanics of materials and biomaterials can be found in dedicated books [2]. The mechanical properties of a biomaterial are not only important for themselves, but they can also condition the fate of cell adhesion and proliferation [9]. In all cases the mechanical properties of an implanted material should be extremely close to those of the material (i.e., tissue) it replaces. Ceramics such as Bioglass used as bone implants have a lower toughness and higher the Young modulus than bone and can thus not be used as substitutes for bones subjected to high loads such as femoral or tibial bones.

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Biomaterials for Organ and Tissue Regeneration

The mechanical properties of a biomaterial must be optimized with respect to the materials’ mass density. For a material having a given volume, and a given range of mechanical properties, the lightest material is always chosen (provided that its price is not prohibitive).

2.2.3 Corrosion and degradation in a given chemical environment Natural body fluids contain dissolved oxygen and other oxidants, and the presence of solutes such as lipopolysaccharides can enhance the oxidation rate of metals [10]. Owing to their corrosion behavior some metals, such as iron, are eliminated because corrosion will produce toxic cations. However, noble metals, such as gold used in dental crowns, are extremely stable from this point of view. Other metals, even if extremely corrodible (having a low standard redox potential), are stable in the presence of oxidants owing to the formation of an impermeable and strongly adhesive oxide layer. This is the case of titanium, the standard metal for dental implants, and other applications: it is naturally covered by a passivation TiO2 layer that affords good corrosion resistance [11]. The thickness and morphology of this passivation layer can be modified through anodization in the presence of electrolytes such as NH4F [12]. Surprisingly, it has been found recently, using a combination of scanning electrochemical microscopy and impedance spectroscopy, that the recovery of the TiO2 passivating layer takes about 15 minutes in a phosphate buffer saline solution at pH 5 7.3 [13].

2.2.4 Control of porosity, pore size, and pore connectivity Most biomaterials as bulk materials are intrinsically nonporous. This may be required for mechanical reasons as well as for the protection against corrosion. Indeed, the corrosion rate is always faster near defects and accessible pores of a structure. However, controlled porosity, pore size, and connectivity between the pores can be of fundamental importance for applications such as bone tissue reconstruction. Bone is naturally a vascularized tissue and is colonized by osteoblasts (involved in bone regeneration) and osteoclasts (involved in remodeling), hence it is a living tissue. The ideal bone substitute would be an autologous graft from the patient for immunological reasons as well as because such a graft is already colonized by cells. However, the amount of available bone is limited, and its uptake is performed at the expense of healthy bone elsewhere in the body (donor site morbidity). Hence substitute materials are required. Pure titanium or its alloys are ideal from a chemical point of view because they are naturally inducing osteointegration (which is a surface property of a material), but their elastic modulus is too high compared with that of bone. Hence many powdered bioceramics and elastomeric composites have been developed for this aim because of simultaneous chemical and mechanical compatibility with bone, and also because the deposited material is granular and has indeed some intrinsic porosity [14]. New additive manufacturing

Physicochemical properties of biomaterials

23

technologies, such as selected laser melting, allow now to produce metal- or alloybased porous structures having simultaneously the required mechanical properties and porosity with pore size in the required diameter range for bone regeneration, namely, between 400 and 1000 μm [15,16].

2.3

Surface properties of biomaterials

2.3.1 Surface energy-hydrophilicity All biomaterials are aimed to interact with biological fluids (blood, serum, interstitial fluid) or tissues, and these interactions occur through an interface. Hence understanding the physicochemical properties of the biomaterial/fluid or biomaterial/tissue interface is of fundamental importance. All interfaces are characterized by a surface tension, the amount of free energy to provide to the material to increase its surface area in contact with another material by one surface area unit. On flat and chemically homogeneous surfaces, when a tiny drop of a liquid is deposited, it can either totally spread on that surface or roll-off or take an intermediate shape. In the first case the adhesion between the liquid and the solid material is extremely strong, and in the second case, there is no adhesion at all and in the latter case the liquid
solid adhesion is intermediate. In this later case, there is a contact line at the air, liquid, solid interface. The tangent line drawn from that contact point (seen from the front as in Fig. 2.2) along the liquid/air interface makes an angle θ with the planar solid substrate. This angle is called the contact angle (Fig. 2.2). The measurement of θ [17,18] gives a mathematical relationship between the three implied surface tensions: γ LVU cos θ 5 γ SV 2 γ SL

(2.1)

Eq. (2.1) is called the Young equation, and γ LV , γ SV , and γ SV are the surface tensions for the liquid/air, solid/air, and solid/liquid interfaces, respectively. Note that Young’s equation is not valid in the case where gravitational forces become predominant with respect to surface tension forces because they will flatten the droplet with a y γLV

Vapor

Liquid Solid

θ

γSV γSL x′

x

Figure 2.2 Schematic representation of the outcome of a contact angle measurement of a liquid drop on an ideally planar and smooth solid. γLV , γ SL , and γ SV represent the interfacial tensions of the liquid/air, the solid/liquid, and the solid/air interfaces, respectively.

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Biomaterials for Organ and Tissue Regeneration

concomitant decrease in the apparent contact angle. This equation can be simply derived by consideration that a surface tension is also a force per unit length (same physical dimension as an energy per unit surface area) and projecting the three forces on the horizontal plane oriented with an x0 x axis. An obvious problem appears with Young’s equation: there is only one measurement, the value of θ, and there are three unknowns, namely, the three surface tensions. Fortunately, the surface tension of most liquid/air interfaces is tabulated and can be easily measured [17]. Since one is interested in γ SL , measurements with at least two different liquids of known γ LV are necessary. One can even get the role of the different intermolecular interactions (polar, nonpolar interactions, acid
base contributions, etc.) in γ SL [17]. Even if it may seem at first glance a trivial surface characterization method, many difficulties are associated with the characterization of liquid wetting processes [19]. Anyway, the contact angle values are influenced by chemical and physical heterogeneities. By physical heterogeneities, we mean surface shape in general and surface roughness in particular. In the case of biomaterials, most of the time one just characterizes the wetting of the materials’ surface by water. In the case, low contact angles are measured, the material is called hydrophilic, and, in the case, where the contact angle is higher than 90 degrees, the material is called hydrophobic. The critical angle for the hydrophilic/hydrophobic transition is, however, subjective and discussed [20]. Most implant materials display improved interaction with cells and biocompatibility when they have an intermittently hydrophilic surface. The surface energy of a material is controlled by a few molecular layers at the interface, meaning that the measured contact angles and hence the calculated surface tensions are extremely dependent on contamination. A nice example comes from TiO2, the natural passivation layer of titanium that should theoretically be a hydrophilic surface owing to the
TiOH groups at the surface. But in practice, it is hydrophobic owing to the presence of adsorbed hydrocarbons from the ambient air [21] which is enough to significantly reduce the bioactivity (interactions with surrounding cells) of the material. Hence great care should be taken to control not only the sterility but also the cleanliness of a biomaterial’s surface, in a reproducible manner for obvious medical practice reasons [22]. Standardized surface treatments are now available to control the surface chemistry of titanium implants as the SLActive process launched by Straumann (Basel, Switzerland). It consists in acid etching, processing under protective gas before final storage in saline medium. It is also possible to process such materials by plasma treatment or using the natural photocatalytic effect of TiO2 (in the anatase form) which is a semiconductor with a bandgap of 3.2 eV, allowing to create a charge separation under UV-visible illumination, typically below 380 nm. The created charges, electrons and holes, will allow for the decomposition of the organic molecules in the contaminating layer. All these surface treatments and conditioning steps have been reviewed, and the interested reader may find more information in Ref. [23]. In the case of titanium screws used for dental implants, the distance between the grooves allows for air bubbles to be trapped in between after the deposition of even

Physicochemical properties of biomaterials

25

tiny (in the μL range) water droplets giving the impression of a hydrophobic surface even after the appropriate cleaning to remove contaminants [24]. This last example is perfectly described in the frame of the Cassie
Baxter model for the wetting of porous surfaces [25] and shows that the real-life (bio)materials are far away from ideal flat surfaces as the one illustrated in Fig. 2.2. This is anyway the case of many surfaces exhibited by plants, insects, etc. The way a bioceramic used for bone substitution interacts with the neighboring bone is an interfacial process guided by the surface composition of the ceramic [26]. Indeed, histological investigations show that the ceramic material is surrounded by a poorly crystalline and carbonated apatite
like material very shortly after implantation. This occurs through a strong interaction between Ca21 cations and the hydroxyl groups available at the surface of the bioceramic material. Hence the surface region close to the implant surface becomes supersaturated with respect to amorphous calcium phosphate or other calcium phosphate phases. This chemical environment is favorable for osteoblast adhesion and proliferation.

2.3.2 Lack of toxicity, of unfavorable immunological response, hemocompatibility A biomaterial should, at least, not induce an unfavorable response of the environmental tissues. In any case the introduction of a biomaterial in the body induces an immunologic response and in most cases is accompanied with bleeding due to surgical intervention. Artificial materials introduced in the body induce a foreign body response [27] and may rapidly be encapsulated by a fibrous cap and insulated from the rest of the body. Both of these responses have to be controlled. In particular, the adhesion of blood platelets (responsible for blood coagulation) on the surface of a biomaterial and their aggregation leading to a thrombus formation is unwanted. The control of the surface chemistry of an implanted material allows to control or even to guide such a response.

2.3.3 Surface topography The behavior of many cell types depends not only on the rigidity of the underlying material but also on its surface topography [28]. Fortunately, the surface topography of many metals can be easily controlled by polishing, anodization, ion implantation, and other techniques. A prototypical investigation showed the sensitive response of human osteoblasts adhering to orthopedic metallic substrates (Ti6Al4V alloy) with variable surface roughness [29].

2.3.4 Protein adsorption Protein adsorption is the first event following hydration of a biomaterial after contact with biological fluids such as blood. The issue of protein adsorption will condition downstream cell adhesion events and depends markedly on the

26

Biomaterials for Organ and Tissue Regeneration

physicochemical nature of the investigated substrate as well as on the protein electrostatics (conditioned by its amino acid sequence) and their conformational stability [30]. Protein adsorption from a complex biological fluid such as plasma or blood is, however, far more complicated than from a single protein solution. Usually, on hydrophilic substrates, the first adsorbed proteins are those with the highest diffusion coefficient because diffusion through a stagnant fluid layer across a concentration gradient always occurs. These first arrived proteins are then often replaced by proteins of higher molecular weight (lower diffusion coefficient). This dynamic adsorption sequence is called the Vroman effect [31] and is somewhat hindered on hydrophobic substrates where the first adsorbed proteins undergo strong conformational changes leading to an exposure of their internally buried hydrophobic amino acids in direct contact with hydrophobic groups on the biomaterial’s surface with a concomitant release of water (and a global entropy increase of the water
protein system) [32]. The influence of tiny changes in the hydrophobicity of just one amino acid, without major structural changes in the protein conformation, can already have a large influence on the protein stability and its adsorption at a solid
liquid interface [33]. Molecular simulations, as in other fields of materials science, appear to be of great help in predicting the behavior of proteins at solid
liquid interfaces [34]. Protein adsorption, as well as contact angle values, is also strongly modified by local chemical and topographical defects on the surface [35]. The final outcome of protein adsorption on a biomaterial is of major importance for the later outcome of cell adhesion and proliferation. Some specific molecular recognition events allow for strong interactions between amino acid sequences on the adsorbed proteins and some receptors present in the cell membrane. A prototypical example is the interaction between the RGD (L-arginine
glycine
L-aspartic acid) triad of amino acids, highly abundant in proteins of the extracellular matrix of cells with cell surface receptor family, integrins [36].

2.3.5 Versatile modification of the biomaterials’ surface chemistry To simultaneously change the hydrophilicity of surfaces and to confer them with specific functionality, polymers may be grafted to surfaces. These polymers may be grafted on specific chemical groups (hydroxyl groups terminating oxide surfaces are an interesting target) or grown from grafted monomers. Of particular interest are the terminal groups pointing outward to the biological milieu, these groups may provide some specific recognition, for instance, RGD peptides, or may passivate the surface to render its antifouling. These grafted polymers allow for a vast repertoire of surface functionalization but with the drawback that the grafting moiety is dependent of the surface chemistry of the biomaterial (here rigorous surface cleaning plays again a major role). In the last 10 years, new versatile coating methods based on rather toxic monomers such as dopamine [37], aniline [38], or aminomalononitrile [39] allowed to produce robust and extremely biocompatible coatings which

Physicochemical properties of biomaterials

27

can also be efficiently postfunctionalized with nanoparticles, drugs, or specific recognition polymers. These findings show that there remains a wide-open window for the design of new biomaterials or coatings able to confer biocompatibility to a material.

2.3.6 Degradability of surface coatings Not all biomaterials need to stay for an indefinite time in the body. This is mandatory for a hip implant but not for a porous scaffold for bone surgery or for tissue engineering. In this case, one expects cells and vasculature to progressively invade the scaffold to create a new tissue. But after a given time, the scaffold is not required anymore and should be degraded. This is possible with hydrolyzable polymers leading to nontoxic degradation products. For instance, poly-α-hydroxyesters [poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(ε-caprolactone) (PCL)] are degradable upon hydrolysis in the presence of water, which allow them to be used as temporary scaffolds in tissue engineering as well as for suture materials and bone fixation. Their degradation products are natural metabolites, for instance, lactic acid in the case of PLA, and are hence nontoxic with the exception of a sudden burst that can induce a strong inflammation due to abrupt pH changes [40]. The hydrolysis rate of such polymers is higher for copolymers and decreases in the following order under equal experimental conditions: PGA . PLA . PCL and decreases also with the molecular mass of each kind of polymer. These thermoplastic materials undergo bulk induced erosion whereas poly(anhydrides) and poly (phosphazenes) undergo a heterogeneous hydrolysis process [41] that is predominantly located at the polymer/water interface. This offers many advantages with respect to bulk hydrolysis because the mechanical properties of the material stay constant during the degradation (the bulk of the polymer remaining intact), and the production of acidic monomers is slowed down as well as the related toxic effects. Polyhydroxyalkanoates are an interesting family of biocompatible polyesters not only from the point of view of their degradability but also because of the possibility to widely tune their mechanical properties from rigid and brittle (as e.g., poly(3hydroxybutyrate) with a strain at break as low as 5%) to highly elastic and stretchable (copolymers of poly(3-hydroxybutyrate) with poly(3-hexanoate) have an elongation at break of 400%
1100% making them useful for cardiovascular engineering [42]).

2.3.7 Antibacterial properties Antimicrobial activity of a biomaterial is mandatory to avoid a biofilm formation resulting in failure of the material. A typical example is peri-implantitis initiated by bacterial invasion in the lumen between the dental implant and the gingiva. This bacterial infection will lead to inflammation, bone erosion, and final failure of the implant. Nowadays, all medical devices are sterilized before implantation, but only a few bacteria present in the environment during implantation can be sufficient to initiate bacterial colonization.

28

Biomaterials for Organ and Tissue Regeneration

Antimicrobial activity is by itself not a physicochemical property of a material but is a consequence of a physicochemical modification of that material. Such a property can be acquired either by a chemical modification, grafting antimicrobials on the surface, or repellant molecules or through a topographic modification [43]. The literature in this domain is vast owing to the formidable challenge the fight against bacteria represents to reduce the occurrence of nosocomial infections. In addition, antimicrobial resistance against antibiotics has increased dramatically which drives the search for physical or new biochemical means of antimicrobial activity. Many interesting alternatives to antibiotics seem to emerge for instance the use of antimicrobial peptides disrupting the cell membrane of bacteria but without a necrotic action against other cells [44], the grafting of nitroxide groups [45,46], etc.

2.3.8 Active biomaterials Biomaterials may be asked to not only play a passive role in the body, without cytotoxicity, but also more efforts are devoted to use surface-immobilized coatings or gels for the release of cytokines able to trigger a positive immune response, namely, an optimal balance between proinflammatory and antiinflammatory effects [47].

2.4

Properties of biomimetic biomaterials

To get a larger repertoire of mechanical properties, surface functionalities, and new design possibilities, it is helpful to look at natural materials already present in living organisms and all having functional properties. Nacre, for instance, has extraordinary mechanical properties and is made from a brick-and-mortar arrangement of chitin and calcium carbonate platelets. Such a structure can be reproduced synthetically by means of a layer-by-layer deposition strategy [48]. This strategy for the design of new potentially useful biomaterials is called “biomimetism.” When tropoelastin fibers from silk are added to silk fibroin, the foreign body inflammatory response with respect to silk fibroin alone is reduced [49]. Substitution materials for bone repair display better performance in terms of toughness, when their structure and composition closely match that of natural bones. This concept is at the origin of design of bone collagen nanocomposites [50]. Surface chemistry aimed to mimic the extracellular matrix of cells has become an intensive research topic, because cells are in contact in a tissue through the polysaccharide-rich coatings [51]. Obviously, the repertoire of possible biocompatible materials and coatings and their properties is far from being fully explored. To reach such an aim, combinatorial approaches are more and more used. For instance, a whole set of structurally

Physicochemical properties of biomaterials

29

related polymers can be modified with a set of drugs [52], and the drug release rate of coatings made from these polymers was investigated. Such approaches are now greatly facilitated with the availability of synthesis robots and multiplexing of analytical platforms as in microfluidic channels.

2.5

Real-time monitoring of an implanted biomaterial and personalized implants

It is of the highest interest to monitor in situ the body response of an implanted material. Each implantation produces an immune response and an acute inflammation [27]. The inflammation process is systematically accompanied by a local pH decrease, called tissue acidosis. Following in situ the time evolution of the pH around an implant and the possible recovery of physiological pH with time is a good way to follow the success or failure of an implant. As an example, poly(Nisopropylacrylamide) nanoparticles modified with a pH sensitive and a pH insensitive dye emitting in the near-infrared region have shown to display a pH-sensitive ratiometric emission in vitro and in vivo. The subcutaneous implantation of silica particles as well as B16F10 malignant cells in C57 mice induced an acidosis and could be followed with this ratiometric near-infrared fluorescence sensor [53]. Many research projects are nowadays developed to couple such sensors, integrated on the implant’s surface with a signal emitting device to allow for real-time monitoring of the physiological state around the implant with a mobile phone. This kind of intelligent biomaterials will for sure improve the success rate of implanted materials by reducing the patients’ intrinsic variability because the medical care services will allow to follow in real time the relevant biological information of the effect of an implanted material. This will allow for a rapid medical intervention in case of appearing abnormalities and to reduce the risk of additional emergency surgery.

2.6

Conclusion and perspectives

Biomaterials have to fulfill a set of bulk properties and surface properties to adapt to their biological environment and to elicit a positive biological response, that is, cell adhesion and progressive incorporation within a tissue or an organ. A deeper understanding of the materials properties and new additive manufacturing methods as well as inspiration from living materials (having the ability to self-repair) already in the body of living organisms will allow for the development of new sets of biomaterials with an improved bioresponse.

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9.

Polymer-based composites for musculoskeletal regenerative medicine

3

Patrina S.P. Poh1,*, Maria A. Woodruff2 and Elena Garcı´a-Gareta3,* 1 Julius Wolff Institute, Charite´—Universitatsmedizin Berlin, Berlin, Germany, 2Institute of ¨ Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia, 3Regenerative Biomaterials Group, RAFT Institute, Mount Vernon Hospital, Northwood, United Kingdom

Abbreviations AA ALP BG CaP CNT ECs ECM HA HEMA HPMA MNPs Tm MSCs mHA mBG MWCNT nHA nBG NGF NGC OPF PTHrP PCL PDLLA PET P(3HB) pHEMA 

acrylic acid alkaline phosphatase bioactive glass calcium-phosphate carbon nanotube endothelial cells extracellular matrix hydroxyapatite 2-hydroxyethyl methacrylate 2-hydroxypropyl methacrylamide magnetic nanoparticles melting temperature mesenchymal stem cells micro-hydroxyapatite microscale bioactive glass multiwalled carbon nanotube nano-hydroxyapatite nanoscale bioactive glass nerve growth factors nerve guide conduit oligo(poly(ethylene glycol) fumarate) parathyroid hormone-related protein poly(ε-caprolactone) poly(D,L-lactic acid) polyethylene terephthalate poly(3hydroxybutyrate) poly(2-hydroxyethyl methacrylate)

Both authors contributed equally to this work.

Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00003-9 © 2020 Elsevier Ltd. All rights reserved.

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Biomaterials for Organ and Tissue Regeneration

PLA PLCL PLCL PRGD PLGA PLLA TMC-CL PU PVA SEM SWCNT SMCs TEVG TGF-β3 β-TCP TMC VEGF VIP WW2

3.1

poly(lactic) acid poly(lactic-acid-caprolactone) poly(L-lactide-co-ε-caprolactone) poly((lactic acid)-co-[(glycolic acid)-alt-(L-lysine)]) poly(lactide-co-glycolide) poly(L-lactic acid) poly(1,3-trimethylenecarbonate-caprolactone) polyurethane poly-vinyl alcohol scanning electron microscopy single-walled carbon nanotube smooth muscle cells tissue-engineered vascular grafts transforming growth factor β3 β-tricalcium phosphate trimethylene carbonate vascular endothelial growth factor vasoactive intestinal peptide World War 2

Introduction

Musculoskeletal disorders are the second largest contributor to disability worldwide comprising more than 150 diagnoses that affect the locomotor system, that is, muscles, bones, joints, and associated tissues such as tendons and ligaments [1]. They range from those that arise suddenly and are short-lived to lifelong conditions associated with ongoing pain and disability. Lifelong disability can be chronic, such as osteoarthritis, and rheumatoid arthritis; acute, such as tendon, ligament, meniscus, cartilage, and bone trauma; or the result other conditions, such as cancer or infection. These conditions and their surgical treatment frequently lead to clinical loss of tissue. When treating critical
sized tissue defects, clinically satisfactory outcomes are most often achieved by autograft [2,3]; however, the harvest of tissue from the patients creates new defects and the possibility of increased morbidity. Allografts are used to overcome the drawbacks of autografts but are beset by the inconsistency of quality and concerns over immunogenicity [4]. Finally, the use of endoprosthetics often requires secondary surgery for replacement as patients’ outgrowth the implant size or as the implant failed to integrate into the surrounding tissue [5]. Over the last decades, the advancement in the medical technologies industries provided the potential to repair and restore musculoskeletal impairments through tissue engineering and regenerative medicine (TERM) strategies. The basic concept of the TERM triad has developed as the field progresses: (1) scaffold to act as temporary supporting structure for tissue growth and maturation, (2) cells to produce the desired tissue, and (3) signaling molecules in the form of biochemical or biophysical cues (e.g., growth factors, mechanical stimuli, and topography) to direct cells behaviors. These concepts have been widely explored—all showing great potential in the regeneration of musculoskeletal tissues [6
8].

Polymer-based composites for musculoskeletal regenerative medicine

35

Through the years, scaffolds of various forms made with a plethora of biomaterials have been extensively studied for musculoskeletal regeneration [9,10]. Owing to the abundance of studies, in this chapter, we highlight polymer
based composite scaffolds used in the regeneration of musculoskeletal tissues: bone, cartilage and osteochondral defects, tendon/ligament/enthesis, and skeletal muscle (Fig. 3.1). Bone and skeletal muscle are highly vascularized and innervated tissues and their successful repair demands regeneration of the blood vessels and nerves that give life to both tissues. Therefore recent advances in composites intended for vascular and nerve regeneration will also be reviewed.

3.2

A brief history of composites

The modern history of composites began in 1935 with the invention of fiberglass by Owen Corning. During World War 2 (WW2), the pace of composite development was accelerated due to the high demand for composite materials for building aircraft. Post-WW2, Owen Corning in partnership with Brandt Goldsworthy extrapolated the application of fiberglass—composite into the boating and automobile industries. Alongside with the development of new polymers, fibers (glass/carbon) and ceramics glasses—the industry of composites—began to mature in the 1970s

Figure 3.1 Schematic showing the basic concept of the TERM triad using polymer-based composite scaffolds and how it can be used for musculoskeletal regeneration, highlighting the tissues discussed in this chapter. Moreover, some musculoskeletal tissues such as bone are highly vascularized and innervated so their successful regeneration involves regenerating blood vessels and nerves. TERM, Tissue engineering and regenerative medicine.

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Biomaterials for Organ and Tissue Regeneration

[11]. It is during this time that research began on composite materials for tissue regeneration.

3.3

Polymer-based composites scaffold characteristics

Composite is a combination of two or more materials, which are different in composition and morphology, to obtain specific biological, chemical, physical, and mechanical properties. Generally, composites comprise two phases: a continuous matrix (primary material) and a dispersed phase or reinforcing agent/filler (secondary material). They can be further classified into three categories according to the shape of the reinforcing agent: (1) particulate composite, (2) fibrous composite, and (3) lamellar composite. In order to achieve an ideal composite material for the intended application, one must consider (1) the composition of the primary and secondary material, (2) the chemical and/or mechanical bond between the matrix (primary material) and the reinforcing agent (secondary material), (3) the volume ratio of primary-to-secondary material, and (4) the surface topography/shape/size of reinforcing agent. The choice of primary and secondary material based on the criterion mentioned is critical for determining the ultimate properties of the composite. All the biomaterials used for scaffold development present advantages and disadvantages, as listed in Table 3.1. Composites have the combined advantages of both the primary and the secondary materials. Moreover, often tissues’ structures are hierarchical (i.e., bone or skeletal muscle) involving different structural levels and components which can be best recreated in the laboratory using composites. Scaffolds intended for musculoskeletal tissue regeneration must have four critical abilities: (1) the scaffold should have a form to act as a temporary tissue expander occupying the space it is designed to fill, (2) the capability to integrate and attach to the surrounding host tissue, (3) the ability to emulate the biomechanical function of the native tissue it replaces, and (4) the capability to drive the regeneration and remodeling of the intended tissue [12]. These abilities can be modulated by the scaffold’s mechanical and biodegradation properties which are governed by the choice of designs and biomaterials.

3.3.1 Mechanical properties Musculoskeletal tissues support body weight and transfer mechanical forces between bones and muscles through ligaments/tendons for locomotion. Thus the mechanical properties of composite scaffolds intended for musculoskeletal tissues regeneration are of critical consideration. Table 3.2 summarizes the compositional and structural factors that affect the mechanical properties of polymer-based composite scaffolds. Mechanically, composites are considered anisotropic, where the mechanical properties vary with the direction of loading. Generally, the inclusion of fillers into a polymer matrix can enhance the composite mechanical strength [13
15].

Table 3.1 Biomaterials used in musculoskeletal tissue regeneration: examples, advantages and disadvantages. Biomaterials Polymers

Examples Natural

G

G

Synthetic

G

G

G

G

G

G

Ceramics

Calciumphosphate Bioglasses and glass
ceramics Others

Metals

G

G

G

G

G

G

G

G

G

Protein: collagen, fibrin, gelatin, silk fibroin Polysaccharides: hyaluronic acid, chondroitin sulfate, cellulose, starch, alginate, agarose, chitosan, carrageenan Poly-glycolic acid Poly-lactic acid Poly-(ε-caprolactone) Poly-(lactide-co-glycolide) Poly-hydroxyethylmethacrylate Oligo(poly(ethylene glycol) fumarate) Hydroxyapatite β-Tricalcium phosphate Silicate bioactive glasses (45S5, 13
93) Borate/borosilicate bioactive glasses (1393B2, 13-93B3) Alumina ceramic Titanium and its alloys Tantalum Stainless steel Magnesium and its alloys

Advantages G

G

G

G

G

G

G

Biodegradability Biocompatibility Bioactivity Unlimited source (some of them) Biodegradability Biocompatibility Versatility

Disadvantages G

G

G

G

G

G

G

G

G

G

G

G

G

Biocompatibility Biodegradability Bioactivity Osteoconductivity Osteoinductivity (depending on structural and chemical properties) Excellent mechanical properties Biocompatibility

G

G

G

G

G

G

G

Carbon

G

G

Composites

G

Carbon nanotubes (single-walled and multiwalled) Carbon nanofibers Combination of the above

G

G

G

Excellent mechanical properties Electrical properties Combination of the above

G

G

G

Low mechanical strength High rates of degradation High batch-to-batch variations Low mechanical strength High local concentration of acidic degradation products Low cellular interaction Brittleness Low fracture strength Unpredictable degradation rates

Lack of tissue adherence Corrosion Risk of toxicity due to release of metal ions Nonbiodegradable (except magnesium and its alloys) Nonbiodegradable Cytotoxic under certain conditions Combination of the above

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Biomaterials for Organ and Tissue Regeneration

Table 3.2 Summary of composition and structural factors that affect the mechanical and degradation properties of polymer-based composites scaffold for musculoskeletal regenerative medicine. Composition

Structure

Choice of primary material (matrix) Choice of secondary material (reinforcement agent or filler) Volume ratio of primary-to-secondary material Composite material fabrication process (e.g., crosslinking of matrix) Fiber thickness of primary material (micro vs nano) Fiber orientation of primary material Reinforcement agent morphology (particle, fibers) Reinforcement agent size (micro vs nano) Reinforcement agent orientation Reinforcement agent distribution Porosity (macro, micro, nano, percentage, interconnectivity, homogenous vs gradients) Pore size (macro, micro, nano) Scaffolds’ fabrication method

However, this is only true if filler loading does not exceed the upper threshold [16
19]. Misra et al. [17] showed that inclusion of ,20 wt.% of spherical nanoscale bioactive glass (nBG) into poly(3-hydroxybutyrate) [P(3HB)] matrix increased the composite Young’s modulus compared to pure P(3HB). Conversely, .30 wt.% spherical nBG loading decreased the composite Young’s modulus but remained higher than pure P(3HB). Similar observations were reported by Vozzi et al. [18] and Mattioli-Belmonte et al. [19], where inclusion of .11 wt.% of carbon nanotubes (CNTs) into poly(L-lactic acid) (PLLA) matrix [18] and .20 wt.% of CNT into PCL matrix [19], respectively, resulted in a rapid decrease in composite elastic modulus. The filler size can also affect the mechanical properties of the composite. As reported by Misra et al. [17] and Jo et al. [20], incorporation of nBG into a polymer matrix showed enhanced mechanical strength compared to a polymer matrix incorporated with micron-sized BG. In addition, the filler distribution and interfacial bonding between fillers and polymer matrix can also affect the composite mechanical strength. Liu et al. [21] illustrated that surface modification of BG with diisocyanate (a short polymer chain) improved the mechanical strength of PLLA/ modified BG composite scaffold compared to that of unmodified PLLA/BG composite scaffold. It was observed that particle distribution was improved within the polymer matrix of modified BG/PLLA, thereby permitting efficient stress transfer between the matrix and filler, which would otherwise be compromised by the presence of particle agglomerations [17,21]. In addition, functionalization of BG surfaces with diisocyanate has enhanced the interfacial bonding between the BG and PLLA contributing to improved mechanical strength. Similarly, it has been shown that chemical crosslinking of the matrix component of nBG/collagen
hyaluronic acid
phosphatidylserine composite, greatly increased the composite elastic modulus compared to composites with noncrosslinked matrix [22].

Polymer-based composites for musculoskeletal regenerative medicine

39

For a scaffold, it has been widely accepted that increased scaffold porosity leads to decreased mechanical strength [13,18,19,23
25]. The filament orientation also plays a role in determining the scaffold mechanical properties [26]. It was shown that despite the same scaffold porosity (B61%), scaffolds with filament orientations of 0/72/144/36/108 degrees have lower stiffness and yield strength compared to scaffolds with filament orientations of 0/60/120 degrees [26]. In addition, dependent on the scaffold fabrication techniques, scaffold mechanical strength varies [27,28]. Jung et al. [27] illustrated that poly(lactic)acid (PLA)/calcium metaphosphate composite scaffolds possess higher tensile strength when prepared via their novel sintering process (0.72 MPa) compared to solvent casting methods (0.13 MPa). While Wu et al. [28] reported that 3D printed mesoporous BG (microscale bioactive glass) scaffolds have greater compressive strength (16.10 6 1.53 MPa) compared to that of scaffolds prepared by polyurethane (PU) templating (0.08 MPa). Collectively, it is possible to tailor the composite scaffolds’ mechanical properties based on the requirements of the intended application. However, one would need to carefully choose the primary and secondary materials, as well as the appropriate fabrication method to fine-tune the desired mechanical properties while taking into consideration the inherent biochemical or biophysical cues of the chosen biomaterials.

3.3.2 Biodegradation properties The degradation of synthetic polymers is through hydrolytic cleavage and enzymatic degradation facilitated by macrophages, with each polymer having a dominant degradation route. Bat et al. have shown that poly(trimethylene carbonate) (TMC) and copolymers of TMC with 10, 20, or 30 mol.% of PCL were fully degraded in vivo in 28 days [29]. In another in vivo degradation study, Pˆego et al. [30] showed that TMC was completely degraded in ,1 month. However, TMC with 90 mol.% of PCL only lost 6 wt.% of its original mass after 1 year. These different degradation rates observed by Bat et al. [29] and Pˆego et al. [30] were due to the different degradation routes undertaken by TMC-CL. TMC-CL with a higher mol.% of PCL following the degradation mode of PCL, where initial degradation is predominantly via hydrolytic cleavage of ester bonds [30]. Details of PCL degradation routes are described in Woodruff et al. [31]. Conversely, TMC-CL with low mol.% of PCL followed the degradation mode of TMC, where enzymatic degradation facilitated by macrophages predominates [29] as described by Zhu et al. [32]. Collectively, this demonstrates that the degradation rate of copolyesters can be varied depending on the blend. The addition of various types of the reinforcing agent (secondary material) into the vast variety of polyesters can further alter the composite degradation rate. Within a composite system, every material has a considerably different rate/ mode of degradation, as shown in Table 3.3. To predict these synthetic composite degradation rates, one must not only consider degradation rate of the individual material but also take into account the chemical/structural changes of the material during the compounding of the composite, interfacial bonds between materials, the

Table 3.3 Rate and mode of degradation of synthetic and natural polymers and composite fillers.

Composite matrix Synthetic polymer

Crystallinity

Poly-glycolide Poly-(L-lactide) Poly-(D,L-lactide) Poly-(L-lactide-coD,L-glycolide)

Semicrystalline Semicrystalline Amorphous Amorphous

Predominant degradation mode/ route Bulk, surface hydrolysis Bulk, surface hydrolysis Bulk, surface hydrolysis Bulk, surface hydrolysis

Poly-caprolactone Poly-anhydrides

Semicrystalline Semicrystalline

Bulk, surface hydrolysis Surface hydrolysis

Poly-phosphazene

Semicrystalline

Surface hydrolysis

Poly-(trimethylene carbonate) Polyhydroxyalkanoates

Amorphous

Surface enzymatic degradation by macrophages Enzymatic degradation by microbial

Oligo(poly (ethylene glycol) fumarate)




Semicrystalline

Bulk, surface hydrolysis

Degradation time

Degradation products

6
12 months .24 months 12
16 months 1
15 months (dependent on the blend) .24 months Days to years dependent on the blend Weeks to months dependent on the side chain structure ,1 month

Glycolic acid Lactic acid Lactic acid Lactic acid and glycolic acid

Weeks to years dependent on the environment


Caproic acid Diacid monomers (dependent on the blend) Phosphate, ammonia, and other products (depends on the side chain structure) Trimethylene carbonate and carbon dioxide Hydroxyacid

Fumaric acid and poly (ethylene glycol)

Natural polymer Collagen Gelatin Fibrin Silk fibroin Albumin Polysaccharides

Origin/definition Fibrous proteins found in bone, cartilage, skin, and other connective tissue Mix of peptides and proteins produced from partial hydrolysis of collagen Blood plasma protein derived from fibrinogen and thrombin One of the two proteins that compose the silk biocomplex excreted by Bombyx mori silkworms during cocoon production Blood plasma protein Monosaccharides joined together by glycosilic linkages

Composite filler Calcium-phosphate ceramics

Example/note (s) HA, α-/β-TCP

Silicate-based bioactive glass

45S5 Bioglass and its derivatives by addition of strontium, zinc, magnesium, etc. into the silicatenetwork Thin, tube-shaped pure carbon

Carbon nanotubes (1
3 nm) Superparamagnetic nanoparticles (,20 nm)

Intrinsically nonmagnetic but can be readily magnetized in the presence of an external magnetic field

HA, Hydroxyapatite, RES, reticuloendothelial system, TCP, tricalcium phosphate.

Degradation mode/route Dissolution of ceramics/bioactive glass by hydrolytic reaction releasing free ions into the surrounding

Removal by macrophages phagocytosis and enzymatic degradation [34] Removal by macrophages phagocytosis via the RES or by diffusion [35]

Degradation mode/rate Enzymatic degradation. Degradation rate varies

Enzymatic (e.g., chitosan, hyaluronic acid, chondroitin sulfate), or hydrolytic (e.g., carrageenan oxidized alginate [33]) degradation; nondegradable (e.g., alginate, cellulose) Degradation/dissolution rate Weeks to months, but generally, the dissolution rate of amorphous HA/TCP . α-TCP . βTCP . crystalline TCP Dependent on the composition of glass and degree of crystallinity


. 3 weeks

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Biomaterials for Organ and Tissue Regeneration

volume ratio of matrix to filler, and size and shape of filler. All these factors will affect the composite material crystallinity and/or hydrophilicity/hydrophobicity, thereby altering the composite degradation rate. Guarino et al. [36] showed that the combination of different materials in a composite system altered the material crystallinity and, subsequently, changed the composite degradation rate. It was observed that PCL crystallinity was highest in PCL/ hyaluronan, followed by PCL/hyaluronan/bioactive ceramics particles and PCL/ hyaluronan/PLA fiber composite. It was reflected in their degradation profile that decreased PCL crystallinity resulted in a faster degradation rate. It was deduced that water can more easily diffuse into the loosely packed amorphous region compare to the crystalline domain. Therefore PCL which possessed the lower degree of crystallinity were more susceptible to hydrolytic degradation, and thus, had a faster degradation rate. Generally, the increased hydrophilicity of a composite system increases the composite degradation rate. It is well reported that incorporation of hydrophilic filler [i.e., BG or calcium-phosphate (CaP) ceramics] into a hydrophobic
polymer matrix can accelerate the composite degradation rate [14,36
42]. Furthermore, increased filler content or reduction in filler size (micro vs nano size) can increase the degradation rate [16,17,43,44] because they increase the surface area available for hydrolytic degradation. Moreover, increased composite hydrophilicity increases water absorption into the interfacial space between matrix and filler via the composite bulk/surface imperfections (i.e., microcracks/microvoids) created during composite scaffold processing. The continuous degradation of the polymer matrix and filler particles enlarges the interfacial space, microcracks, and microvoids within the composite bulk. This, in turn, increases water uptake/diffusion into the composite bulk, thereby accelerating degradation. Chemical crosslinking of the matrix components can also affect the degradation rate of a composite. As shown by Wang et al. [22], chemical crosslinking of the organic matrix (i.e., collagen
hyaluronic acid
phosphatidylserine) of a composite containing inorganic fillers (i.e., BG) greatly reduced the water swelling ratio, thereby deferring degradation. For composites containing polymers that undergo hydrolytic degradation, the general trend of degradation is that increased water uptake will accelerate hydrolytic degradation, thereby accelerating composite degradation. However, this general degradation trend may not be true in all cases. For example, El-Kady et al. reported that although 50 PLLA:50 nBG wt.% composites had higher water uptake compared to 75 PLLA:25 nBG wt.% composites, the degradation rate of 50 PLLA:50 nBG wt.% was slower than that of 75 PLLA:25 nBG wt.% composites. The authors deduced that this phenomenon may be caused by the decreased porosity of 50 PLLA:50 nBG wt.% compared to 75 PLLA:25 nBG wt.% [45]. In another study, Huang et al. incorporated β-tricalcium phosphate (β-TCP)/hydroxyapatite (HA) particles into a hydrogel matrix containing different ratios of monomers, namely, acrylic acid (AA), 2-hydroxypropyl methacrylamide (HPMA), or 2hydroxyethyl methacrylate (HEMA) crosslinked with N,O-dimethacryloyl hydroxylamine. It was observed that the water uptake by AA-HEMA/HA/β-TCP composites

Polymer-based composites for musculoskeletal regenerative medicine

43

was greater than HPMA-HEMA/HA/β-TCP ones. However, it was found that the degradation rate of AA-HEMA/HA/β-TCP composite was slower than HPMAHEMA/HA/β-TCP composites. The authors believed that the strong interaction of the hydroxyl and carboxylic side-groups of AA-HEMA hydrogel with the Ca21 of HA/β-TCP played a dominant role in deferring degradation [46]. Taken together, the degradation behavior/trend of polymer
based composites is tuneable. However, the interactions between different materials are very complex. To generate a composite with controllable degradation rate, one must fully understand the complex interactions between different materials, as well as the effect of various material processing methods, composite compounding and fabrication methods. The factors affecting the degradation of composite materials were discussed based on in vitro studies. In order to better estimate the degradation time of composite material, long-term in vivo studies are essential. This is because the presence of cells, enzymes, physiological fluid, etc. in vivo can contribute to the degradation of these composite materials. For example, Niiranen et al. [47] showed that the degradation of PLA and PLA/BG scaffolds was more pronounced in vivo compared to in vitro. Similar outcomes were demonstrated by Fu et al. [48] and Liao et al. [49], where the accelerated rate of composite degradation in vivo compared to in vitro was reported. Unfortunately, in both in vivo degradation studies endpoint (12 and 10 weeks, respectively), the composites did not undergo complete degradation. Consequently, the in vivo degradation time cannot be determined. Hence, long-term in vivo studies would be recommended to determine the degradation time for any given composite.

3.4

Polymer-based composite scaffolds for specific musculoskeletal tissue regeneration

3.4.1 Bone Bone is a highly dynamic tissue that has the important intrinsic ability to spontaneously regenerate and remodel upon injury. However, above a critical size defect due to injury, disease, or malformation, this ability is lost, and bone grafting is necessary. Bone graft substitutes (BGS) derived from natural or synthetic biomaterials offer great alternatives to allografts or autografts; however, the currently available BGS offer minimal structural integrity and different degrees of efficacy on bone regeneration dependent on the healing environments [50]. Presently, despite various graft materials that mimic bone properties (reviewed in [50]), treatment of bone voids or defects, especially in patients with comorbidities, remains an unsolved clinical challenge. Therefore hierarchical composite scaffold constructs (Fig. 3.2) have been under intense development to provide a potential synthetic bone graft to this unmet clinical need.

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Figure 3.2 Variety of polymer-based composite for bone tissue engineering. (A and B) Freeze-dried scaffolds of nHA—pullulan/dextran polysaccharide composite. Magnified SEM image (B) showed a dispersion of nHA through the pullulan/dextran polysaccharide matrix [74]. (C and D) Bioactive glass and photocrosslinkable poly(ε-caprolactone) scaffold fabricated by stereolithography [89]. (E) Illustrating the formation of (F) nanocomposite materials mimicking mineralized collagen fibrils using CNTs and HA crystals [112]. (G) Electrically conductive scaffolds made from a composite of PLA and MWCNTs. The graph showed the effect of MWCNTs contents on the surface resistance of electrospun composite PLA/MWCNTs nanofiber meshes [125]. (H) Magnetized fibers made of polycaprolactone (PCL) and FeHA composite. The value of saturation magnetization (M, emu/g) increases when the PCL/FeHA mass ratio decreases [130]. CNT, Carbon nanotubes; FeHA, iron oxide hydroxyapatite; HA, hydroxyapatite; MWCNT, multiwalled carbon nanotube; nHA, nanohydroxyapatite; PLA, polylactide; SEM, scanning electron microscopy. Source: All figures were reproduced with permission.

3.4.1.1 Composite with bioactive ceramics/glasses Although synthetic polymers have tailorable mechanical and degradation properties, it has been widely accepted that synthetic polymers, when used in isolation, have limited beneficial effects on bone formation, both in vitro [16,20,21,51
54] and

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in vivo [20,51,55]. Thus for the regeneration of bone, the use of synthetic polymers in combination with bioactive materials such as CaP-ceramics and silicate-based BG is considered more favorable. This group of materials is known for their bioactivity, which is the ability to form a hydroxycarbonate layer at the material/fluid interface when exposed to biological fluid, thereby forming a strong bond to host bone [56
58]. It is widely believed that with the optimal amount of bioactive materials incorporated within a polymer bulk, the composite is capable of creating an osteogenic environment to stimulate the differentiation of cells into bone-forming cells (osteoinductive), as well as promoting the development of a mineralized interface as a natural bonding junction between living tissues and nonliving scaffolds (termed bioactive). Consequently, encouraging regeneration of the defect sites with improved osseointegration between living bone and implant (the direct structural and functional connection between living bone and implant [59]).

3.4.1.1.1 Composites with calcium phosphate
based bioceramics CaP-based bioceramics include HA (Ca10(OH)2(PO4)6), α-TCP, β-TCP, and biphasic CaP (BCP). The driving force for the research and development on these CaP-based bioceramics is that they highly resemble the inorganic component of human bone (B50% of the dry weight of human bone is a modified form of HA [60]). HA, when compared to TCP, has slower resorption rate and undergoes little conversion to a bone-like material after implantation [61]. Subsequently, BCP, a mixture of HA and α- or β-TCP, was generated, of which their bioactivity and resorption kinetics are dependent on the HA:TCP ratio. In recent years, nano-HA (nHA) has gained considerable attention for bone tissue engineering. It has been suggested that synthetic nHA, which closely resembles bone mineral composition, size, and morphology, possesses better osteoconductivity [62]. This may be contributed by the increased surface roughness and surface wettability which leads to enhanced protein adsorption and subsequently cell interaction with nHA surfaces in comparison to micro-HA (mHA) surfaces [63]. Webster et al. have shown enhanced osteoblasts adhesions [63] and osteoclast-like cell functions [64] on nHA sintered discs compared to mHA sintered discs. Similarly, Balasundaram et al. [65] illustrated that osteoblasts adhesion on nHA is comparable to that of RGD (Arg-Gly-Asp)-functionalized mHA. Although nHA possesses great potential for bone regeneration, its application in bone tissue engineering is limited due to its inherent brittleness and poor fatigue resistance. Thus they are commonly used as injectable cement paste at nonload bearing sites and coatings for implants. To explore the potential use for repairing large osseous defects, nHA is often incorporated within a polymeric matrix to form composite bone scaffolds [66,67]. Natural polymers, for example, chitosan, collagen, and polysaccharides, are commonly used as the matrix material for composite containing nHA fillers [68
74]. Chitosan is a biodegradable polysaccharide which shares a number of chemical and structural similarities with collagen. It has been reported that the addition of chitosan during coprecipitation of nHA can reduce the crystallinity and the crystallites of nHA, which in turn increase the bioresorption of nHA [68,69].

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Through freeze-drying technique, Thein-Han and Misra [70] fabricated scaffolds consisting chitosan with 1 wt.% nHA particles (50 nm) that greatly improved the cell attachment and proliferation compared to that of pure chitosan scaffolds. Using the same scaffold fabrication method, Kong et al. [71,72,75] incorporated 12 wt.% of nHA (70
100 nm) into a chitosan scaffold, which showed enhanced bone formation in vivo when combined with autologous bone morrow. Similarly, Zhang et al. [73] showed healing of femoral defect filled with chitosan-nHA paste. Using freeze-dried polysaccharide-based scaffolds composed of pellulan, dextran, and nHA (50 nm), Fricain et al. [74] were able to show bone regeneration in critical-size orthopedic defects in preclinical small and large animal models, in three different bony sites, that is, the femoral condyle of rat, the mandibular defect, and tibial osteotomy in goat. Although natural polymers have a beneficial effect on the bioresorbability of nHA, they lack mechanical stability and shape fidelity after implantation. To circumvent this shortcoming, Wang et al. [76] encapsulated nHA (20
40 nm) into chitosan microspheres that were further dispersed through a matrix of poly(lactideco-glycolide) (PLGA) with (10:10:80 wt.%). In vitro studies showed maturation and mineralization of cell extracellular matrix (ECM) on the PLGA/chitosan/nHA scaffolds. However, no mineralized tissue was formed at 14 days postimplantation using an ectopic nude mice model. This may be due to the relatively short implantation period or the low nHA content within the scaffolds. With increasing nHA content, Wang et al. [55] showed that polyamide/nHA (40:60 wt.%) composite scaffolds can achieve complete osteointegration of rabbit mandibular defects after 12 weeks. Similar biological responses were shown by Qian et al. [54] in their PCL-nHA scaffold.

3.4.1.1.2 Composites with bioactive glasses The first generation of BG with a silicate composition was developed in 1969 by Hench et al. and is commonly known as 45S5 BG or Bioglass (46.13 SiO2
2.60 P2O5
24.35 Na2O
26.91 CaO mol.%). Over the years, various types of BG with different chemical compositions have been developed. These BG have additional elements incorporated into the silicate network such as magnesium, strontium, iron, boron, zinc, and copper [77
79]. Generally, all BG are bioactive. The bioactivity of BG is driven by the continuous dissolution of the BG when in contact with physiological fluid, followed by condensation and polymerization of the dissolution ions on the surface of BG resulting in the formation of a crystalline carbonated-substitute HA-like layer on the surface of the glass [57,58,80]. However, different glass composition can result in different extents of bioactivity due to the change in BG dissolution rate. The BG dissolution rate is governed by the amount of glass network former or modifier within the silicate-glass network. Consequently, this will affect cell behavior. It is also well documented that selected dissolution ions of BG have a beneficial effect on bone cells [77,81]. Similar to CaP-based ceramics, BG are brittle. Thus they are often incorporated into the polymeric matrix. It has been well reported that with the right amount of bioactive material, polymer composites containing BG are able to

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display inherent bioactivity in vitro, as evidenced by the formation of CaP precipitates on the material surfaces when immersed in stimulated body fluid or culture media under physiological conditions [44,45,52,53,82
91]. However, the extent of bioactivity (rate of CaP formation) of composites is dependent on the amount of bioactive material incorporated into the polymer bulk. Generally, composites containing higher amounts of bioactive material showed greater bioactivity and exhibited a faster rate of CaP precipitation [44,45,82
84,89,92,93]. In addition, it was also reported that the inclusion of nano-sized bioactive material instead of micron-sized bioactive material further accelerated the CaP formation rate [17]. The high content bioactive material with smaller bioactive material size, provide a larger surface area for the dissolution of bioactive materials when in contact with fluid at physiological conditions. This, in turn, accelerated the condensation and polymerization of the dissolution ions on the surface of bioactive material, resulting in the faster rate of crystalline carbonated
substitute HA-like layer formation on the surface of the composite. Generally, the incorporation of BG into the polymer bulk has been shown to accelerate the cell proliferation rate, especially at early time-points; as well as enhancing osteoblast differentiation and matrix mineralization when compared to polymer-only scaffolds [20,27,52,53,82
84,86,89,94,95]. Most studies show that increased BG loading can further enhance cell adhesion, cell proliferation, osteoblast differentiation, and ECM mineralization [20,52,83]. For example, Kim et al. [52] incorporated 5 or 25 wt.% of nBG (45S5 Bioglass) into PLA and molded the composite into a film by thermal compression. Cell proliferation and matrix mineralization were greater in PLA/25BG . PLA/5BG . PLA [52]. Conversely, in some studies, no significant difference in terms of cell proliferation and matrix mineralization were found between polymer-based composites loaded with different amounts of BG [84,86]. Blaker et al. [84] incorporated 5 or 40 wt.% of 45S5 BG into poly(D,L-lactic acid) (PDLLA) bulk and porous scaffolds were fabricated using thermally induced phase separation. Cell proliferation and matrix mineralization were greater on composites compared to neat PDLLA. However, no significant differences were observed between the two composites. Similarly, Meretoja et al. [86] found no significant differences in osteoblast differentiation and mineralization between poly(caprolactone-co-DL-lactide) (P(CL-DLLA)) composites containing 5 or 10 wt.% of S53P4 granules (53% SiO2, 23% Na2O, 20% CaO, and 4% P2O5). In one of the studies, it was shown that cell proliferation and osteoblast differentiation were lower in PDLLA/40 wt.%-45S5 scaffolds group compared to PDLLA/5 wt.%-45S5 scaffold group [96]. The authors suggested that prolonged pretreatment of the PDLLA/40 wt.%-45S5 scaffolds may prevent cytotoxic ion release, thereby improving the experimental outcomes. This is in line with the findings of Verrier et al. [97] who showed increased cellular adhesion to PDLLA/40 wt.%-45S5 scaffolds pretreated in culture medium for 3 days prior to cell seeding compare to PDLLA or PDLLA/5 wt.%-45S5 scaffolds. Generally, it is accepted that BG can promote the formation of a strong bond with bone due to its inherent bioactivity [58,98] and this property has been detailed in [99]. Likewise, for polymer/BG composite, it was illustrated by

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Marcolongo et al. [100] that interfacial bond strength of polysulfone/BG composites (80:20 wt.%) with host bone tissue was significantly higher than polysulfone control groups (12.4 vs 5.2 MPa). However, no significant difference was found between composite and polysulfone groups in terms of bone contact at the implant surface. To date, to our best knowledge, there is a lack of quantitative studies on interfacial bonding strength between polymer/BG composite and host bone. Soft tissue animal models can be used to assess the osteoinductivity of any given biomaterial. Under the soft tissue environment, any mineralized tissue/mature bone formation observed can be directly attributed to the biomaterial’s inherent osteoinductive capability. Several studies on soft tissue implantation of BG-only paste/ scaffolds showed the ability of these scaffolds to form mineralized tissue and bone [101,102]. Conversely, studies on subcutaneous implantation of polymer/BG composite film/scaffolds into small animal models showed the limited mineral formation and, to our best knowledge, no mature bone formation was reported [94,96,103]. In one study, tissue in-growth was observed in the composite disk (P (CL-DLLA)/S52P4 BG composite—50 or 70 wt.% of BG) but not in neat polymer matrices. However, after 6 months, no apparent bone formation was detected in any of the groups [103]. Using porous scaffolds, similar outcomes were also reported by Gomide et al. [94] and Yang et al. [96] with poly-vinyl alcohol (PVA)/BG [50 wt. % BG, 70SiO2
30CaO (wt.%)] and PDLLA/45S5 (5 or 40 wt.% BG), respectively. Notably, 8 weeks postimplantation, Gomide et al. [94] showed that composite scaffolds seeded with differentiated mesenchymal stem cells (MSCs) showed some degree of osteoid tissue formation which was not observed in cell-free scaffolds and scaffolds seeded with nondifferentiated MSCs. Therefore differentiated MSCs may be the major contributing factors to the observed osteoid formation. The authors also showed that in the cell-seeded composite scaffold constructs, necrosis of cells was observed in the central region of the scaffolds. Unfortunately, no indepth discussion was presented by the authors with regards to the cell necrosis observed. In both studies [94,96], the extent of vascularization, which is a crucial factor in determining the rate of tissue remodeling, matrix mineralization and bone formation within the scaffold [104,105], was also not investigated. In contrast to in vivo animal studies conducted using soft tissue models for the assessment of polymer/BG composite osteoinductivity, a greater consensus was generally observed in animal studies conducted using a bone defect model for the polymer/BG composite osteoconductivity. Generally, it was found that polymer/BG composites had an accelerated bone regeneration rate compared to the respective polymer-only film/scaffold [20,106,107]. Using a porous polyamide/BG (25 wt.% BG) scaffold obtained via solvent casting/porogen leaching technique, Su et al. [106] showed that polyamide/BG composite scaffolds had greater newly formed bone compared to polyamide-only scaffolds. In studies using nonporous polymer/BG composite [20,107], Jo et al. [20] revealed that defects covered with PCL/BG composite membrane (20 wt.% of BG) have thicker and more bone growth compared to untreated defects and PCL membrane groups. Furthermore, the histomorphometric analysis showed that the regenerated bone area of the composite group was significantly higher than the untreated

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defect and PCL membrane group. No significant difference between the untreated defect group and PCL membrane group was detected. Thus the study indicated that incorporation of BG into PCL bulk could enhance the rate of bone regeneration [20]. In another study, where commercially available S53P4 BG [53 SiO2, 23 Na2O, 20 CaO, and 4 P2O5 (wt.%)] was incorporated into fiber-reinforced composite (FRC) implants, it was shown that the biological behavior of FRC and FRC/BG composite implant was comparable to that of a titanium implant after 4 and 12 weeks of implantation. Furthermore, FRC/BG implants showed increased periimplant osteogenesis and bone maturation compared to FRC and titanium implant groups [107]. Collectively, to date, soft tissue implantation of cell-free polymer/BG composite scaffolds tested have not shown mineralized tissue/mature bone formation. This indicates that the implanted PCL/BG composites are not sufficiently osteoinductive. However, when implanted in close vicinity to host bone (bone defect animal model), polymer/BG composites have demonstrated significant improvements in bone formation and regeneration rate at the defect site when compared to polymeronly constructs. This indicates that polymer/BG composites possess improved bone regeneration capability compared to polymer-only constructs. This may be because in a bone-forming environment, the free ions released from the dissolution of BG has a synergistic effect with the surrounding host bone microenvironment, thus accelerating cell proliferation, differentiation, tissue remodeling, matrix mineralization, and bone regeneration. All in all, it was found that composites containing BG can better promote bone regeneration as compared to pure polymer. However, the optimal amount of BG may vary between different studies and are not comparable as BG composition and size, amount of BG incorporated into the polymer bulk, and scaffold fabrication techniques can all affect the outcomes.

3.4.1.2 Nanocomposite materials mimicking mineralized collagen fibrils In natural bone, collagen fibrils consist of self-assembled collagen triple helices with a quarterly staggered pattern, producing a 67 nm long periodic structure with 40 nm gap and 27 nm folded part [108]. The gaps provide sites for mineral nucleation and growth, where HA crystals are assembled onto the collagen fibrils with their c-axes aligned along the longitudinal axes of collagen fibrils, giving mineralized collagen fibrils [108]. Based upon this nanostructure of bone, CNTs
HA composite has been developed; where CNT acts as the collagen fibrils and HA represent the mineralized phase. Such a composite can be achieved through chemical deposition of HA onto CNT surfaces or chemical vapor deposition (CVD) process to facilitate the growth of CNTs on HA particles. For the former method, due to the relatively strong π
π interaction between CNT molecules, the limited number of nucleation sites of pristine CNTs, and the similarity of the surface negative charge of HA and pristine CNTs [109], various surface functionalization on CNTs surfaces have been employed to achieve a homogeneous deposition of HA on

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CNTs surfaces when immersed in solution containing CaP precursors [109
112]. On the other hand, the CVD process utilizes metallic catalyst particles (i.e., Fe, Ni or Co) to facilitate homogeneous growth of CNTs on HA surfaces [113
116]. Although some breakthroughs on nano-HA/CNT composite synthesis have been achieved, this biomimetic nanocomposite material is still at its infancy with debatable biocompatibility. In vitro, HA/CNT composites have been showed to be noncytotoxic [116,117], support osteoblasts attachment and proliferation [111,118]. However, to date, there is only a handful of in vivo biocompatibility studies on HA/ CNT composites. Sintered HA/CNT composites implanted intramuscularly in rats showed pronounced inflammation in the first postoperative week which subsided after 2 weeks [119]. In another study, sintered rod-shaped HA/CNTs were implanted into defects created on the diaphysis of rats [120]. After 4 weeks, no fibrous tissues were observed between the newly formed bone and composite implants, indicating good biocompatibility. Nonetheless, long-term in vivo studies for biocompatibility for nano-HA/CNT composite are lacking.

3.4.1.3 Electrically conductive composites In the late 1950s it was discovered that electrical potential is present on the mechanically loaded bone [121]. Since then, the application of electrical stimuli through a metal electrode implanted at the defect sites to promote bone regeneration has been widely explored with positive outcomes reported in numerous in vivo studies [122,123]. This initial method involved secondary surgical procedures for the removal of the electrode which possesses the potential risk of infection at the implantation site and damage to the newly formed bone tissue. Furthermore, bone formation was limited to the periphery of the electrode tip and does not encompass the extent of the defect. To overcome this, Supronowicz et al. [124] incorporated CNT into PLA bulk to form an electrically conductive composite. When external electrical stimuli were applied on the PLA/CNTs composite, osteoblast proliferation increased by 46% after 2 days, and the extracellular calcium concentration and collagen type 1 gene expression were significantly enhanced. Using electrospun scaffolds made of PLA incorporated with acid oxidized CNT, Shao et al. [125] illustrated that 100 μA electrical stimuli enhance osteoblasts proliferation, but 200 μA electrical stimuli hindered the growth and proliferation of osteoblasts.

3.4.1.4 Magnetized composites Magnetic nanoparticles (MNPs) less than 20 nm in size are known as superparamagnetic, that is, they are nonmagnetic without an exterior magnetic field. However, in a nanoscale, each MNP itself is considered a single magnetic domain with a magnetic field [126]. Therapies involving the application of MNPs are already well known in the nanomedicine field, with the main application as drug delivery vehicle, treatment for cancer hyperthermia, and MRI signal enhancement agent [127]. The application of MNPs has also penetrated the field of bone TERM

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[126,128]. In earlier publications, MNPs were incorporated into prefabricated HA/ collagen [127] or HA/TCP [126] scaffolds by means of dip-coating. Advancement in additive manufacturing (AM) technologies has enabled the fabrication of magnetic composite scaffolds with MNPs incorporated into the bulk of PCL [129,130], PLA [131], or PLGA [128]. In vitro studies have shown that magnetized composites scaffolds were able to support osteoblasts adhesion and accelerate osteoblasts proliferation rate [126
128,131] when compared to nonmagnetized scaffolds. However, the concept of using such magnetized scaffolds to enable the loading/reloading of growth factors or bioagent to the scaffold sites at any desired time-point thereby aiding the regeneration of the defect is yet to be proven [127,129].

3.4.2 Cartilage and osteochondral regeneration Cartilage is a stiff, dense, and mostly inflexible connective tissue found in the joints between bones as well as in other parts of the human body such as the ears or nose. In this chapter, we will focus on repair of articular cartilage and the cartilage
bone interface (i.e., osteochondral defects). Chondrocytes, the main cell type found in cartilage tissue, produce large amounts of ECM composed of collagen and elastin fibers and proteoglycans. Four different areas can be distinguished in cartilage: (1) superficial zone, where flattened chondrocytes are observed and collagen fibers are parallel to the articular surface; (2) middle zone, where rounded chondrocytes and less parallel fibers are observed; (3) deep zone, with chondrocytes ordered in vertical columns and collagen fibers perpendicularly aligned to the articular surface; and (4) calcified cartilage, composed of hypertrophic chondrocytes [132]. Underlying the cartilage is subchondral bone [132]. Unlike the other connective tissues in the body, cartilage is avascular whereby maintenance of chondrocytes functions is through waste/nutrients diffusion. Therefore cartilage and osteochondral defects have very limited capacity for regeneration [133]. Tissue engineering approaches have been adopted for ex vivo construction of cartilage and subchondral construct using scaffolds made of various biomaterials [134]. The ECM that chondrocytes reside is commonly recapitulated using hydrogels [135]. In recent years, hydrogel design has advanced to improve their utility in AM of scaffolds that act as a carrier for cells or biological signaling molecules [135]. To mimic the cartilage/subchondral defect structure, composites are widely utilized. This includes composites of natural/natural, natural/synthetic, synthetic/ synthetic polymers, and combination of natural or synthetic polymers with bioceramics/BG.

3.4.2.1 Composites of natural polymers To this end, there exist products for cartilage repair made of a composite of natural/ natural polymers approved for clinical used or under clinical trials, for example, Cartipatch (Tissue Bank of France, Lyon, France); an agarose/alginate hydrogel with autologous chondrocytes and Novocart 3D (TETEC Tissue Engineering

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Technologies AG, Reutlingen, Germany); and a 3D collagen/chondroitin sulfate matrix with autologous chondrocytes [136]. Using the concept of the hydrogel as a delivery vehicle for biological molecules, Bian et al. [137] embedded alginate microspheres encapsulating transforming growth factor β3 (TGF-β3) and/or parathyroid hormone-related protein (PTHrP) and human MSCs within a hyaluronic acid hydrogels. In vivo results showed that constructs containing TGF-β3 microspheres resulted in superior cartilage matrix formation compared to the codelivery of PTHrP with TGF-β3. To improve cartilage regeneration outcome the release kinetics of PTHrP and TGF-β3 need to be finetuned.

3.4.2.2 Composites of synthetic and natural polymers 3.4.2.2.1 Gelatin/PCL Composites comprising gelatin and PCL arise to combine the favorable mechanical properties of PCL [31] with the superior biological performance of gelatin due to the natural presence of cell recognition amino acid sequences, that is, RGD [138]. Several studies have used gelatin/PCL composites for the repair of cartilage defects. Feng et al. [139] found that the phase separation of gelatin and PCL (50:50 wt. %) during the process of electrospinning can be overcome by the addition of acetic acid. This yielded scaffolds with thinner, smoother, and homogeneous nanofibers. Later, Xue et al. [140] sandwiched 5 3 105 of chondrocytes between two gelatin/ PCL membranes and showed that preculture of such constructs leads to cartilage formation and was able to retain their shape. Later, Zheng et al. [141] investigated the cartilage regeneration potential of scaffolds comprising of different ratios of gelatin and PCL (70:30—G7P3; 50:50—G5P5; and 30:70—G3P7 wt.%). Using the same chondrocytes/scaffold construct assembling technique as Xue et al. [140], chondrocytes/scaffold constructs were implanted subcutaneously into nude rats. Good cartilage regeneration was observed for groups G7P3 and G5P5 just after 3 weeks in vivo. It was shown that G3P7 scaffolds disrupted the cartilage regeneration due to the presence of a large volume of the nondegraded scaffold, indicating that the high content of PCL was unfavorable for 3D cartilage regeneration [141].

3.4.2.2.2 Oligo(poly(ethylene glycol)fumarate) gel/gelatin microparticles Oligo(poly(ethylene glycol) fumarate) (OPF) is a linear unsaturated macromer, which consists of two repeating units of the poly(ethylene glycol) and fumaric acid that are alternately linked by ester bonds. OPF is water soluble and the degradation of the polymer occurs by hydrolysis of the ester bond [142]. OPFs have been used in the form of hydrogels for cartilage/osteochondral defects repair [143,144]. Several studies have loaded gelatin microparticles containing growth factors into the matrix of OPF hydrogel creating novel biphasic composite scaffolds for cartilage and osteochondral repair [143
145]. Using this approach, the Mikos’ lab showed the possibility of controlling the release kinetics of growth factors by adjusting parameters such as OPF formulation and microparticle loading that

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affected the composite swelling behavior [143
148]. Through the variation of scaffold compositions tested, it was indicated that OPG/gelatin scaffolds alone were not sufficient to promote subchondral or hyaline cartilage regeneration. The incorporation of growth factors (TGF-β3 and insulin-like growth factor-1) had a beneficial effect on hyaline cartilage regeneration. However, the addition of TGF-β3 may have hampered the subchondral bone regeneration due to its’ inhibitory effect on osteogenic differentiation of MSCs and osteoblasts.

3.4.2.2.3 Poly(lactide-co-glycolide)-gelatin/chondroitin sulfate/hyaluronic acid scaffold Scaffolds made of gelatin/chondroitin sulfate/hyaluronic acid lack mechanical stability. Hence, Fan et al. [149] developed a PLGA-gelatin/chondroitin sulfate/hyaluronic acid scaffold, whereby microsponges made of gelatin/chondroitin sulfate/ hyaluronic acid were formed within the pores of mechanically stable PLGA macrostructure. In addition, the amine groups of TGF-β3 (1 ng/μL) were crosslinked with the carboxyl groups of the hybrid scaffolds to mimic the natural ECM of cartilage. In vivo implantation of the preseeded scaffolds into rabbit osteochondral defects showed better cartilage regeneration in scaffolds containing TGF-β3 compared to control scaffolds. These results suggested that the scaffold alone did not support cartilage regeneration and that growth factors are the key to cartilage regeneration. However, it is unclear if the host cells played a major role, as acellular scaffolds were not tested.

3.4.2.3 Composites of synthetic polymers Very few examples of composites combining only synthetic polymers for cartilage and osteochondral regeneration can be found in the literature compared to the numerous studies of composites incorporating natural polymers. This highlights the preference of biomaterials scientists for incorporating components of biological origin to mimic the natural ECM of tissues. In one of the very few examples, the potential of PVA/PCL nanofiber scaffolds seeded with MSCs from rabbit bone marrow for cartilage regeneration was investigated. In vivo showed improved healing of full-thickness cartilage defects compared with untreated control and cell-free composite scaffolds. This indicated that MSCs seeded on the PVA/PCL composite scaffolds is the driving force for articular cartilage repair [150]. Interestingly, a composite of poly-glycolic/poly-lactic acid and poly-dioxanone with autologous chondrocytes has been commercialized by Biotissue Technologies, Freiburg, Germany under the name of BioSeed-C [151]. Currently, this product is only commercially available in the EU for the treatment of traumatic and focal chondral and osteochondral defects (Grade III
V according to Outerbridge).

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3.4.2.4 Composites of synthetic or natural polymers with bioceramics Bioceramic chemical composition is similar to that of the biomineral present in bone. Hence, the strategies of incorporating bioceramics into synthetic or natural polymers have also been explored for the creation of composites for osteochondral repair. For example, Xue et al. [152] showed that PLGA/HA (90:10 wt.%) scaffolds with B88% porosity preseeded and precultured with 1 3 105 MSCs for 12 days can fully regenerate a full-thickness osteochondral defect in rat after 12 weeks, while PLGA scaffold group exhibited mainly fibrocartilage characteristics. This indicated that the addition of nHA into PLGA matrix has a beneficial effect in promoting hyaline cartilage regeneration in the presence of MSCs. In another study, Fan et al. [153] investigated the potential of HA-TCP and PLGA/HA-TCP scaffolds to act as a supporting matrix for autologous articular cartilage transplant in a sheep model for osteochondral regeneration. After 3 months, it was revealed that transplanted cartilage layer supported by HA
TCP scaffolds had better cartilage quality, structural integration and higher mechanical strength than that supported by PLGA/HA
TCP scaffolds. The results indicated that the physicochemical properties, including the inherent mechanical strength and material chemistry of the scaffolds, play important roles in influencing the repair of osteochondral defects driven by autologous cartilage transplants.

3.4.3 Tendon, ligament, and enthesis regeneration Tendons and ligaments are fibrous connective tissues composed of mostly collagen I fiber bundles that are densely packed and parallelly aligned in the direction of tensile force. Tendons connect skeletal muscle to bone, while ligaments connect bone to bone. The main function of both tissues is mechanical. Ligaments guide joints through their normal range of movements when a tensile load is applied, while tendons conduct forces from skeletal muscles to bones, resulting in joint movement [154]. Similar to cartilage, tendons and ligaments are avascular, hence have limited regenerative capabilities. When tendon/ligament replacement surgery is necessary, besides auto-/allografts, a synthetic graft of polyethylene terephthalate (PET) fibers is the most commonly used synthetic tendon/ligament with encouraging short-term results [155] but lacks long-term outcomes. Moreover, to date, the regeneration of the enthesis, that is, the tendon/ligament-to-bone interface or insertion, remains an unmet clinical challenge due to its graded nature transitioning from unmineralized soft tissue to mineralized bone through an unmineralized fibro-cartilage intermediate [156]. Therefore TERM strategies using a variety of biomaterials in the form of scaffolds with the desired mechanical properties have been extensively studied to develop tissue substitutes for the repair of tendon and ligament. To date, collagen I fibers remain the most popular biomaterial used for the repair of tendon and ligament as they mimic the fibrous structure of these tissues [154,156]. To overcome

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the inherently weak mechanical stability and tensile strength of collagen I fibers, composites of collagen I with natural (e.g., silk fibroin, cellulose) or synthetic (e.g., PCL, PLA, and PLGA) polymers have been developed and evaluated for their potential as tendon/ligament substitutes.

3.4.3.1 Natural composites Bombyx mori silkworms excrete a silk biocomplex containing two proteins during cocoon production: silk fibroin and sericin. Silk fibroin is 70%
80% by mass of the silk biocomplex and is commonly used as a biomaterial in TERM after the separation from sericin using sodium carbonate, urea, detergents at near-boiling temperatures [157]. The superior tensile strength and toughness of silk fibroin, that is, elastic modulus is 5
9 GPa, tensile strength is 250
400 MPa and failure strain is 23%
26% [158], and the ease to process it into any form, that is, gel, films, fibers, braided fibers, make silk fibroin a good biomaterial candidate for tendon and ligament repair. Altman et al. were the first group to develop a braided silk fibroin scaffold and showed its’ potential for the repair of the ligament [159]. Later, Chen et al. [160] embedded the braided silk fibroin scaffold within a collagen matrix and showed better ligament regeneration and stronger scaffold
ligament interface compared to that of silk fibroin-only scaffold within a rabbit medial collateral ligament defect model.

3.4.3.2 Synthetic composites Poly-α-hydroxyl esters, such as PGA, PLLA, and PLGA, are also commonly used biomaterials for tendon and ligament repair. Due to the hydrophobic nature of polyα-hydroxyl esters, Chen et al. incorporated freeze-dried collagen I onto a knitted mesh of polyglactin 910 (90:10 copolymer of glycolic acid and lactic acid) [161] to better support cells attachment. In an attempt to engineer ligament that includes the enthesis region, Spalazzi et al. [162,163] developed a triphasic composite scaffold. The scaffold comprises a ligament region made of poly-L-lactide-co-glycolide knitted mesh seeded with primary bovine anterior cruciate ligament fibroblasts; a nonmineralized fibro-cartilage intermediate region composed of poly(D-L-lactide-co-glycolide) microspheres formed by water/oil/water emulsion; and the enthesis region comprises poly(D-Llactide-co-glycolide) and 45S5 Bioglass microspheres seeded with primary bovine trabecular bone osteoblasts. These studies successfully demonstrated the feasibility of regenerating the ligament and enthesis using a single construct.

3.4.4 Skeletal muscle regeneration Skeletal muscle, a specialized tissue that is involved in dynamic events like locomotion or mastication, is composed of muscle fibers that are formed by dense and oriented myoblasts into long, cylindrical, and multinucleated myotubes. Muscle fibers bundle into fascicles which further bundle together into muscles. Clinical

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regeneration of skeletal muscle is necessary after myopathy, trauma, tumor extraction, or muscle denervation. Tissue engineering of skeletal muscle was first demonstrated by Strohman et al. in 1990 [164], who grew a monolayer of myoblasts on a membrane of Saran Wrap (The Dow Chemical Company, United States), trademark for polyvinylidene chloride polymers. Myoblasts vigorously contracted and deformed the Saran Wrap membrane, from which they ultimately detached. Since this first skeletal muscle tissue construct, scientists have developed new strategies using a variety of composite biomaterials as scaffolds for skeletal muscle regeneration. To date, the main challenge in engineering scaffolds for skeletal muscle regeneration is to guide the prealignment of myoblasts into muscle fibers, which ultimately bundle together into muscles, providing the mechanical properties that can sustain the repeated contraction and relaxation movements of natural skeletal muscle [165]. Scaffolds made of composite materials are commonly used as temporary supporting and guiding structure to achieve skeletal muscle regeneration.

3.4.4.1 Composites of synthetic and natural polymers Mimicking the skeletal muscle ECM made of predominantly collagen is a popular strategy for the engineering of skeletal muscle [166,167]. Besides that, gelatin, the natural polymer produced from partial hydrolysis of collagen, has also been employed within a composite system for skeletal muscle regeneration. Kim et al. [168] showed that PCL/gelatine nanofibers crosslinked with genipin can support the proliferation of myoblasts and the formation of myotubes. Another study showed that methacrylated gelatin hydrogels containing vertically aligned CNT can better support skeletal muscle cells growth and yielded a higher number of functional myofibers than cells that were cultured on hydrogels with randomly distributed CNTs and horizontally aligned CNTs [169].

3.4.4.2 Synthetic composites Composite of synthetic polymers has also been studied for their potential for regeneration of skeletal muscle. Bandyopadhyay et al. [170] showed that scaffolds made of 70/30 L-lactide/ε-caprolactone (PLC) seeded with human myoblasts can form multinucleated myotubes expressing human muscle-specific markers when implanted ectopically under the skin in SCID mice. This suggests that PLC scaffolds loaded with myoblasts can be used for skeletal muscle engineering or for inducing muscle repair. Leveraging on electrospinning technique, McKeon-Fisher et al. have developed a variety of synthetic composite scaffolds comprise of a PCL or PLLA matrix filled with poly(3,4-ethylenedioxythiophene) nanoparticles or multiwalled carbon nanotubes (MWCNTs) or gold nanoparticles [171
174] and tested them for their potential for skeletal muscle repair. Among which, only the PCL-MWCNT scaffolds with an exterior sheath of polyAA/PVA hydrogel were evaluated in vivo and showed good biocompatibility and supported skeletal muscle regeneration.

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The necessity for nerve and vascular regeneration

Tissue regeneration is followed by tissue remodeling to ensure restoration of the tissue functions. Hence, the continuous supply of nutrients and the sensing of appropriate stimuli through the vasculature and nervous system respectively are of utmost importance to guarantee successful tissue regeneration and remodeling.

3.5.1 Importance of vasculature and innervation for skeletal muscle regeneration Skeletal muscles are highly vascularized to enable nutrients/waste transport for the maintenance of muscle functions. Anatomically, skeletal muscles are made up of basic contractile units of muscle fibers. Each muscle fiber, also term myofiber, upon maturation, is innervated by a single motor neuron. The innervation of motor neuron into myofiber has been implicated to play a role in specifying the myofiber contractile properties [165]. Each of the innervated myofibers is surrounded by a layer of connective tissue which assembles into fascicles and further bundle together into muscles. Muscle is a functional unit in which individual myofibers contract upon neuron activation, and the cumulative effect of myofibers contractions is transformed into movement via the myotendinous junctions, the sites where myofibers attach to skeletal through tendons [165]. Thus it is apparent that the functional properties of skeletal muscles are dependent on the maintenance of the complex framework of myofibers, vasculature, and nerve innervation.

3.5.2 Importance of vasculature and innervation for bone regeneration Bone vasculature is essential for bone development, growth, remodeling, and homeostasis. As with skeletal muscle, the primary role of vasculature is to supply bone with nutrients as well as growth factors, cytokines, hormones, or chemokines and remove waste products. Moreover, vasculature in bone plays a communicative role between bone and surrounding tissues [175]. On the other hand, the function of nerve innervation in bone is less understood. However, it is widely accepted that the central nervous system regulates bone development and skeletal remodeling via direct innervations and indirect hormonal pathways. Recently, accumulating evidence has indicated that nerve reinnervation plays a major role in fracture healing besides transmitting sensations such as pain and proprioception. In numerous in vivo studies, it was shown that sensory nerve reinnervation into the fracture sites promoted bone regeneration [176
178]. Collectively, bone regeneration should par up with the regeneration of nerves and vasculature. Nonetheless, nerve and vascular regeneration are very challenging research areas. In the following sections, the current strategies for nerve and vascular regeneration using polymer-based composites will be reviewed.

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Nerve regeneration

Clinically, nerve regeneration can be achieved through the use of nerve guide conduits (NGCs). The concept of using NGC for nerve regeneration has existed for more than a decade and has evolved over time to clinical reality as an alternative to autologous nerve grafting [179]. The first generation of NGC is of tubular shape, with a single hollow lumen that may lead to polyinnervation or inappropriate target reinnervation due to random dispersion of regenerating axons through the NGC lumen [179,180]. In addition, such single lumen hollow NGC is incapable of supporting the formation of fibrin cable between the proximal and distal stumps during the initial stages of regeneration. Therefore their application is typically limited to nerve gaps of ,3 cm with small diameter and noncrucial nerves (e.g., digital and radial sensory nerves) [179,180]. Therefore different approaches have been undertaken by researchers to further improve the efficacy of NGC, in order to achieve an ideal NGC which is capable of bridging larger nerve gap (Fig. 3.3). An ideal NGC should be (1) biodegradable at a controlled rate whereby, initially, the NGC should remain intact to allow axon regeneration across the nerve gap, then degrade gradually to prevent axonal compression, (2) biocompatible to avoid excessive unresolved immunological responses, (3) sufficiently permeable to allow nutrients
waste diffusion but limiting scar infiltration, (4) mechanically robust to support axonal regeneration and sufficiently flexible to enable bridging of nerve gap which span across different plane, and (5) electrically conductive to promote neurite outgrowth, as electrical synapses is the key component of neural communication [179,180]. To develop an ideal functional artificial NGC, scientists often employed a biomaterial approach in combination with structural adaptation.

3.6.1 Composite biomaterial approach for nerve regeneration At present, a wide variety of biomaterials have been attempted for the fabrication of NGC, either of natural origin or synthetically derived materials. In this chapter

Figure 3.3 Schematic of designs for an NGC to promote directional axon growth and regeneration. NGC, Nerve guide conduit.

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the focus will be on NGC made of synthetic polymer
based composite. An in-depth review of NGCs made of natural biomaterials and composites have been published by Khaing et al. [181]. Biodegradable synthetic polymers that have been explored as NGCs include, but not limited to, PGA, PLA, and PCL, using various fabrication techniques as detailed by Sarker et al. [182]. Although the use of biodegradable synthetic polymers as NGC can have some merits (e.g., biodegradable and biocompatible), it often lacks the structural/chemical cues and physical/mechanical properties needed to direct neuronal regeneration over a large nerve gap. Thus biodegradable synthetic polymers are often used in the form of composites, either as polymer blends (copolymers) or in combination with other biomaterials or bioactive factors, to harvest the desirable properties of each material which is favorable for promoting nerve regeneration.

3.6.1.1 Composites of synthetic polymers (copolymers) Synthetic polymer blends (copolymers) commonly utilized for nerve regeneration include poly(lactic-acid-caprolactone) (PLCL) [183,184], PLGA [185,186], and poly(1,3-trimethylenecarbonate-caprolactone) (TMC-CL) [187,188]. In an early study, a crystalline copolymer of PLCL (50/50 PLLA/PCL) was utilized for the long-term (2 years) nerve regeneration study [189]. Although the 50/ 50 PLLA/PCL conduit showed the ability to promote nerve regeneration, the degradation of biomaterials was not completed after 2 years. This can be a disadvantage as nerve fibers can grow across an 8 mm gap over 3 weeks, and maturation of a sciatic nerve of a rat takes place within 16 weeks. Therefore slow degrading biomaterials can potentially impede the process of nerve regeneration and functional recovery. Subsequently, an amorphous 50:50 PDLLA/PCL copolymer (50/50 (85/ 15L/D)PLA/PCL) was synthesized, fabricated into an NGC, and implanted into a rat sciatic nerve defect model [184]. The authors showed that the 50:50 PDLLA/ PCL conduit was able to support nerve regeneration in rats sciatic nerve defect [184], and the performance was superior to autologous nerve graft [190]. However, due to the 50:50 PDLLA/PCL inherent swelling properties, a conduit with a thick wall can cause blockage of the conduit lumen and hence impede nerve regeneration. Conversely, when the conduit wall is thin, the conduit was found to collapse easily and consequently disrupt nerve regeneration. Furthermore, it was showed that the conduit losses its strength after 2 months of immersion in PBS [184]. Therefore in a further study, it was noted that the 50:50 PDLLA/PCL conduit degradation rate was too fast and is not an optimal conduit for the repair of long nerve defects in humans as the rate of nerve regeneration in the human body is slower than in rats [183]. Subsequently, a different copolymer composition of amorphous 65:35 PDLLA/PCL (65/35 (85/15L/D)PLA/PCL) was investigated. The 65:35 PDLLA/ PCL conduits maintained their mechanical strength and flexibility for up to 10 weeks and only started losing mass after 10 weeks of immersion in PBS. In the same study, it was also showed that the 65:35 PDLLA/PCL conduits were noncytotoxic and hemocompatible [191]. Long-term (2 years) in vivo degradation studies on 65:35 PDLLA/PCL conduits showed that the conduits do resorb. However, the

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resorption was not complete even after 24 months of implantation in a rat sciatic nerve defect model, with biomaterial fragments with multinucleated giant cells and macrophages found along the regenerated nerve tissue [192]. Therefore the authors put forward that the long-term safety use of the new amorphous PLCL is questionable [192]. In the meantime, the 65:35 PDLLA/PCL conduit was US Food and Drug Administration approved and commercialized as Nuerolac. To date, clinical outcome has been variable [193]. A randomized, multicenter trial demonstrated that patients with up to 20 mm hand nerve lesion treated with Nuerolac achieved recovery of sensibility compared to that of patients who received end-to-end nerve suture treatment [194]. In another clinical follow-up study with average 11.03 mm nerve defect, of the 28 implanted Neurolac conduits, 8 complications were reported, which included protrusion of the conduit and difficulty in wound healing. In addition, of the eight reconstructions with .11 months follow-up, 75% showed poor sensory recovery [195]. To date, Neurolac is the only commercially available copolymer NGC. Due to the tuneable properties (i.e., degradation rate, swelling ratio, and mechanical properties) of synthetic polymer composites, intense research into different synthetic polymer blends intended for nerve regeneration is currently in progress. A group from Twente University (Netherlands) developed a series of copolymers consisting of trimethylene carbonate and ε-caprolactone (TMC-CL). As the TMC content increased, the biomaterial became more elastic with Young’s modulus ranging from 6.3
141 MPa and stress at break ranging from 0.03
12 MPa [187]. Later, it was demonstrated that TMC-CL can support human Schwann cells attachment and proliferation, with the copolymer having a higher content of TMC performing better [30]. However, at TMC content .70% the TMC-CL completely degraded within a month in vivo [29]. With reference to various requirements of a nerve conduit and consideration of processability, the authors suggested that TML-CL of 10:90 mol.% is the most suitable candidate to be developed into nerve graft. However, in vivo implantation of cell-free TMC-CL conduit into 20 mm did not show any sign of nerve regeneration [196]. In contrast, implantation of cell-seeded TMC-CL conduits into 40 mm medial nerve defects in rats resulted in partial nerve regeneration as compared to autograft group [197].

3.6.1.2 Composites containing carbon nanostructures It has been suggested that electrically conductive biomaterials can promote nerve regeneration as neurons transmit signals through an electrical potential. One attractive biomaterial, known for its electrical conductive properties and widely used in the field of nerve regeneration, is carbon nanostructures, that is, single-walled CNTs (SWCNTs), MWCNTs, and carbon nanofibers. It has been shown that carbon nanostructures can maintain and promote electrical activity in networks of cultured neurons [198], potentially through the discontinuous and tight interaction between carbon nanostructures and neuronal membrane surfaces, which can in turn induce specific changes in the cell membrane electrical behavior [198]. However, in recent

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years, due to reported cytotoxicity of carbon nanostructures [199,200], effort has been moving toward the generation of composites where carbon nanostructures are often incorporated within a polymer matrix. Such composites are able to retain the electrically conductive property of carbon nanostructures [201,202], while hindering the direct contact of cells with carbon nanostructures, thereby minimizing its cytotoxic effect [77]. A study by Arslantunali et al. demonstrated that the presence of MWCNTs may contribute toward neural cell survival during the conduction of electrical potential [201]. In the study, it was found that neuroblastoma cells maintained their viability on poly(2-hydroxyethyl methacrylate) (pHEMA)/MWCNTs composite upon application of electrical potential, while cells on pure pHEMA did not survive [201]. Using an electrospun composite made of PLLA and SWCNTs or MWCNTs, Kabiri et al. [203] demonstrated that the composite upregulated mature neuronal markers expression as compared to pure PLLA scaffolds. This finding coincides with Kam et al. [204], where SWNT/laminin composite thin film demonstrated the ability to promote neural stem cells’ growth, proliferation, and differentiation. Therefore it was suggested that the conductive material may have added value in neural regeneration as it can provide a substrate for the delivery of programmed and spatially controlled electrical stimulation to the repaired tissue. The potential of carbon nanostructures as a component within a composite for nerve regeneration is enormous. However, many challenges remain for the generation of polymer/carbon nanostructure composites which possess the superb properties of carbon nanostructures. A review by Spitalsky et al. [205] gives a comprehensive overview of current challenges for the production of polymer/carbon nanostructure composites and strategies employed by researchers to enhance the beneficial properties of carbon nanostructures within a composite system.

3.6.1.3 Polymer
ceramics composites A tendon-chitosan (t-chitosan) conduit coated with HA (t-chitosan/HAp) demonstrated the ability to support nerve regeneration superior to that of t-chitosan conduit [206]. Furthermore, the addition of laminin-1 (a basement membrane protein that can promote cell adhesion, spreading, migration, growth, and neurite outgrowth) or laminin peptides onto the t-chitosan-HAp conduit further enhanced the nerve regeneration capability, with comparable histological outcomes as isografts. However, it was noted that the functional recovery of the regenerated nerve was delayed in the conduit groups as compared to isografts [206]. Recently, Qiu et al. [207] constructed a composite NGC comprising PDLLA, poly((lactic acid)-co-[(glycolic acid)-alt-(L-lysine)]) (PRGD) and β-TCP (PDLLA/ PRGD/βTCP). The composite constructs were implanted into a rat sciatic nerve gap model for up to 35 days, with groups of PDLLA and PDLLA/PRGD conduits used for comparison. Postimplantation analysis showed that PDLLA/PRGD/βTCP conduits achieved the best nerve regeneration in terms of myelinated axon regeneration and functional recovery. Notably, it was shown that PDLLA/PRGD/βTCP conduits can suppress oxidative stress after injury by enhancing the expression of antioxidant

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enzymes, hence promoting the expression of cytoskeletal proteins, resulting in better nerve regeneration outcomes compared to PDLLA and PDLLA/PRGD conduits [207].

3.6.1.4 Synthetic polymer
natural polymer composites A bilayer PU/collagen conduit fabricated via rapid prototyping was developed in the Zhang lab [208,209]. The conduit has an outer PU layer with micropores (15
20 μm) to prevent fibrous tissue infiltration into the conduit lumen, and an inner collagen layer with oriented nano-size filaments with pore sizes of 20
100 μm to permits nutrients infiltration. In vitro, the PU/collagen conduit showed compatibility with Schwann cells [209] and was later implanted into a 10 mm peroneal nerve defect of rat [208]. The bilayer conduit demonstrated superior nerve regeneration compared to animals implanted with pure PU conduit [208]. However, in the study, a positive control group of nerve autograft was not included; thus no baseline was provided for the evaluation of the efficacy of the bilayered PU/collagen conduit.

3.6.1.5 Composite nerve guide conduit with structural adaptation To achieve the ideal NGC, besides the selection of biomaterials, the structural design of an NGC is also a critical factor in determining the efficacy of the NGC [182]. Thus in more recent years, research has been focusing on improving the internal lumen structure of NGC to enable the bridging of larger nerve gaps. The modifications on the internal structure of the NGC lumen aim to provide a better scaffolding to support and enable fibrin matrix formation across a larger nerve gap which is a crucial step in the initial phase of nerve regeneration, as well as serving as topographical guidance to the regenerating axons and migrating Schwann cells. Currently, structural modifications to the NGC lumen include (A) the incorporation of one or more intraluminal channels, giving a multichannel NGC; (B) application of electrospinning to produce fibrous NGC; and (C) the incorporation of intraluminal guidance. These structural modifications are often used in isolation or in combination with the aim of achieving an optimum NGC design which is capable of bridging a large nerve gap ( . 4 mm) [182]. Due to the demanding requirement of NGC, optimum structural design of NGC may be complemented by the right choice of biomaterials. And often, to achieve such complex structural design of NGC, the combination of biomaterials is utilized to make up the different structure of the NGC. Therefore in the following section, the different NGC structures made from composites of biomaterials will be reviewed.

3.6.1.5.1 Multichannel nerve guide conduit Multichannel NGC was developed with the intention of mimicking the architecture of nerve fascicles [182,210]. The use of multichannel NGCs seem to be able to reduce the dispersion of regenerating axons but display no significant benefits over

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single lumen NGCs in term of promoting nerve regeneration in vivo [182,210]. These findings suggest that further improvement on the multichannel structure is needed to achieve a construct that more closely resembles the nerve structure, that is, the band of Bu¨ngner (proliferating Schwann cell lined basal lamina tubes). However, the fabrication of such a complex structure is technologically challenging. Consequently, to our best knowledge, no composite materials have been attempted for the fabrication of such multichannel NGC.

3.6.1.5.2 Electrospun fibrous nerve guide conduit The advantages of using electrospun tubes as NGC over a continuous tube is that the tube fabricated can be highly flexible and porous, therefore can be well adapted for use within biological systems [211]. It has been demonstrated that electrospun fibrous NGC can support nerve regeneration in vivo [212,213]. In an in vivo study, graded electrospun tubes comprise an outer layer of PCL/PLGA tight-knit mesh and an inner layer of large fiber PCL mesh was implanted into a 10 mm rat sciatic nerve gap [213]. After 4 months, sciatic nerve failed to reconnect in the untreated group, while most animals that received treatment with the electrospun NGC showed nerve regeneration and functional reconnection of the nerve gap with neurite outgrowth oriented along the longitudinal conduit axis. However, with this approach, 40% of the electrospun NGCs were collapsed and displaced from the defect site and consequently excluded from the study. In addition, after 4 months, the lumen of the electrospun NGC showed shrinkage at the mid-point of the 10 mm gaps, most likely caused by muscle compression exerted during pacing, as well as cellular infiltration within the tube wall [213]. Therefore to prevent electrospun NGC from collapsing and displacement from the defect site, further reinforcement is necessary while not compromising the flexibility of the NGC. As with the regeneration of other parts of musculoskeletal parts, the rate of nerve regeneration can also be further enhanced with the applications of bioactive factors such as glial cell
derived neurotrophic factor, nerve growth factors (NGFs), neurotrophin, and many others [211]. Utilizing coaxial electrospinning, a composite of electrospun NGC with graded architecture was fabricated by Liu et al. [214]. The generated NGC was composed of an outer shell of P(LLA-CL) (50:50) and a core containing the recombinant rat β-NGF/bovine serum albumin dissolved in PBS. The composite NGC was implanted into 10 mm rat sciatic nerve gap model for 3 months. Postimplantation analysis showed that the P(LLA-CL)/NGF composite performed as well as autograft with comparable nerve function recovery and nerve regeneration and performed significantly better compared to empty P(LLA-CL) and P(LLA-CL) with injected NGF. In another study, laminin peptides with varying degrees of glycine spacer were covalently bound onto an electrospun chitosan mesh, then the construct was inserted into a solvent-casted chitosan tube, resulting in a bilayered NGC. In vivo evaluation of nerve, regeneration capability showed that conduits with immobilized laminin peptides performed better than nontreated conduits and were comparable to the isograft group. Furthermore, by increasing the space between the active laminin peptide and the electrospun chitosan mesh (increased glycine spacer length), the

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tendency of fusion of minifascicles to larger fascicles increased, resulting in better nerve regeneration outcomes [215].

3.6.1.5.3 Nerve guide conduit with intraluminal guidance To overcome the limited structural support of hollow NGC, intraluminal guidance is often incorporated within the lumen of NGC. However, to date, no standard has been established for the optimal configuration of such intraluminal guidance for NGC [182,211]. This may be because the efficacy of such intraluminal guidance is dependent upon several other considerations. Among the consideration for intraluminal guidance is the orientation of the fibers within the lumen of NGC. In numerous in vivo studies, the incorporation of aligned fibers within the lumen of NGC has shown good nerve regeneration efficacy in a rat sciatic nerve gap model [216,217]. Studies have demonstrated that the packing density of fibers within the lumen of NGC can affect axonal regeneration [218,219]. In a study by Huang et al., silkbased NGC (lumen diameter: 1.6 mm) containing 0 (PN0), 100 (PN100), 200 (PN200), or 300 (PN300) intraluminal Spridrex fibers (fiber diameter: 10 3 20 μm) were implanted into a rat sciatic nerve model. After 4 weeks, PN200 possessed significantly highest axon density and supported 60% as much axon growth when compared to autologous nerve graft control. Therefore the authors suggested that a too low packing density within the NGC lumen provided insufficient support for axonal regeneration, while high packing density may lead to physical obstruction of the NGC lumen, thereby preventing axonal regeneration [218]. Another consideration for intraluminal guidance is biomaterial selection, as it may play a role in determining its efficacy. For example, Schnell et al. [220] demonstrated that in vitro collagen/PCL (25:75) composite scaffolds were able to better support Schwann cell migration, neurite orientation, and process formation of Schwann cells, fibroblasts, and olfactory ensheathing cells when compared to PCL electrospun fibers. Furthermore, for optimal nerve regeneration, the intraluminal guidance should have an appropriate degradation rate as not to obstruct the process of axonal growth; hence selection of biomaterial with an appropriate degradation rate is crucial. Using a multimodel approach, Quigley et al. [219] fabricated an NGC consisting of intraluminal guidance made of melt-extruded PLGA fibers (30
40 μm) enclosed within an outer membrane of knitted PLA sheet with PLA fibers electrospun on the outer wall, with a pore size of , 2 μm. Prior to implantation into a rat sciatic nerve gap model, an alginate hydrogel encapsulated with a cocktail of neuro factors were casted within the lumen of the NGC. Over 4 weeks of implantation, it was reported that the effect of self-mutilation was reduced in the group implanted with the fullconfiguration conduit compared to the group implanted with conduits containing only an alginate hydrogel. The histological analysis also showed that conduits containing intraluminal guidance can better promote nerve regeneration over conduit without intraluminal guidance. Another common combination of biomaterials for the assembly of NGC with intraluminal guidance is of synthetically derived polymer (i.e., PGA, PLGA) and naturally derived polymer (i.e., collagen, chitosan). For example, conduits made of

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collagen-coated PGA mesh containing intraluminal guidance of collagen fibers coated with laminin were implanted into dogs with 80 mm nerve gap. Post analysis demonstrated good nerve regeneration and sensory nerve function restoration [221]. In another study, Wang et al. implanted a chitosan conduit containing intraluminal PGA fibers (chitosan/PGA) into a 30 mm sciatic nerve defect in the dog. After 6 months, the NGC completely degraded and the dog sciatic nerve trunk had been reconstructed with the restoration of nerve continuity and functional recovery [222]. Later, the chitosan/PGA NGC was used to repair a 35 mm long median nerve defect in a male patient. During the 36 months follow-up period, the patient showed gradual functional recovery of the sensory nerve with no postoperative discomfort reported [223]. Collectively, the use of composite biomaterials continues to show progress in the field of nerve regeneration. With composites, researchers are able to achieve NGC with comparable mechanical properties of native nerve. In combination with the strategic design of NGC, composite-based NGC demonstrated great potential for nerve regeneration.

3.7

Vascular regeneration

During the process of musculoskeletal tissue regeneration, insufficient nutrient and oxygen supply, as well as waste removal from the regenerating/remodeling site, limit the successful repair of tissues [176,224]. It has been reported that different cells survive at different distances from the blood vessels (from 100 μm to even 1 mm), thus showing various sensitivities to oxygen [225]. These are important considerations to take into account when designing scaffolds for tissue regeneration. Several strategies are used today to achieve the goal of vascularized scaffolds upon implantation in vivo. Incorporation of cells [mature and precursor endothelial cells (ECs) as well as in coculture with stem cells, progenitor, or fully differentiated cells] into the scaffolds and/or addition of pro-angiogenic factors such as vascular endothelial growth factor (VEGF) are popular ones [226,227]. Microsurgery strategies such as flap fabrication and arteriovenous loop are also used [228,229]. However, in this chapter, we will focus on polymer-based composite scaffolds that were designed to incorporate vascular networks or constructs that were engineered to regenerate small blood vessels (,6 mm).

3.7.1 Fabrication of polymer-based composite scaffolds that incorporate a vascular network Scaffolds’ porosity is important for cellular infiltration and growth, ECM deposition, and vascularization [230]. The effect of pore size on the growth of ECs on polymeric scaffolds was investigated by Narayan and Venkatraman [231]. It was observed that cell growth was higher with a smaller pore size of 5
20 μm, as similarly reported for osteoblasts [232]. Contrary to the observations that a higher

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porosity and pore size favor greater bone ingrowth and vascularization in vivo [232], smaller porosity and pore sizes induced the creation of a hypoxic environment, which promotes chondrogenesis [228]. Hence, scaffold architecture can be used to control vascularization. For example, patterning of vascular networks in scaffolds is a promising strategy to achieve vascularized tissue-engineered constructs. Muller et al. [233] incorporated vascular network design into PCL/HA scaffolds that were fabricated by fused deposition modeling. Implantation of the cell-seeded scaffold into a rat model alongside an arteriovenous bundle showed vascular ingrowth along with the patterned vascular network. Capillary and connective tissues throughout the constructs were observed [233]. In another study, Zhao et al. [234] fabricated PLGA scaffolds with interconnected channels by a lowtemperature 3D printing system. The inner channels of the PLGA scaffolds were loaded with a fibrin/collagen hydrogel containing adipose-derived stem cells that differentiated to ECs and smooth muscle cells (SMCs), both of which are found in blood vessels, and formed vascular-like structures. Thus the authors produced a scaffold with incorporated vasculature with potential in complex organ manufacture.

3.7.2 Engineering of small diameter blood vessels (,6 mm) with polymer-based composites Synthetic materials such as PET (Dacron) or polytetrafluoroethylene are currently used to replace medium to large diameter blood vessels ($6 mm) with good success [235,236]. However, these materials are unsuitable for the replacement of small diameter blood vessels (,6 mm) where there is a risk of thrombosis, calcification, restenosis, and lack of growth in pediatric applications [236]. Therefore the investigation into other biomaterials for small diameter tissue-engineered vascular grafts (TEVGs) is urged. Catto et al. summarized the requirements for the ideal small diameter TEVG in their comprehensive review article [236] with biocompatibility, mechanical properties, and processability being the three main categories to comply with.

3.7.2.1 Synthetic polymer
based composites containing collagen and gelatin Collagen and elastin are the primary ECM structural components of blood vessels. Therefore numerous examples across the literature can be found using collagen and its derivative gelatin in combination with synthetic polymers [237
239]. Among the many studies, Fu et al. [239] fabricated electrospun gelatin/PCL and collagen/ PLCL composite scaffolds as small diameter TEVG. The collagen/PLCL scaffolds possessed greater Young’s modulus (1.77 6 0.09 MPa compared to 1.49 6 0.06 MPa for gelatin/PCL) and formed homogeneous vessel-like structures when subcutaneously implanted into nude mice [239]. Using gelatin but in combination with Tecophilic, Vatankhah et al. fabricated nanofibers as potential small diameter TEVG [240,241]. Studies showed that the

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composite nanofibers served as an adhesive substrate for vascular SMCs when compared to Tecophilic alone. Moreover, the phenotype of SMCs was preserved on the Tecophilic/gelatin scaffolds thus suggesting the potential of these scaffolds as small diameter TEVG. Similarly, Ino et al. [242] and Atlan et al. [243] investigated the potential of PVA/gelatin tubular scaffold as graft material for small diameter vessel. Unfortunately, long-term in vivo study revealed graft failure as graft patency reduced over time, possibly due to changes in the mechanical properties [243]. In another study, a composite scaffold consisting of an inner layer of poly(ethylene glycol)-b-poly(L-lactide-co-ε-caprolactone)/gelatin loaded with platelet-derived growth factor, a middle layer of PLGA/gelatin loaded with VEGF, and an outer layer of PCL/gelatin was fabricated [244]. The addition of gelatin to the composite multilayered scaffold provided an initial improvement in vascular ECs adhesion followed by facilitated ingrowth of vascular SMCs as it degraded. More importantly, it was shown that the composite scaffold was able to maintain patency in the rabbit left common carotid artery for up to 8 weeks [244].

3.8

Conclusion and future prospects

Musculoskeletal disorders are the second largest contributor to disability worldwide affecting muscles, bones, joints, and associated tissues such as tendons and ligaments. TERM strategies offer the potential to repair and restore musculoskeletal impairments by using scaffolds that can include cells and/or biochemical or biophysical cues. Scaffolds of various forms are made with a plethora of biomaterials, among which polymers are widely used. However, perfect biomaterial does not exist. By combining different biomaterials into composites, we can improve their properties and performance. Moreover, the intricate and often hierarchical structures of body tissues such as bone or skeletal muscle can be best recreated with the use of composites where different materials and morphologies can be combined. To find the “perfect” composite for each tissue requires a thorough understanding of the individual biomaterials’ properties, as well as the composite mechanical and biodegradable features. In terms of “off-the-shelf” availability, it would be of interest to find the “perfect” composite for each type of tissue. Looking into the future, research into new biomaterials, especially synthetic polymers, is of key importance for the development of this research field. Additive manufacturing, or more commonly known as 3D printing, will be important for fine-tuning complex 3D morphologies, as well as mechanical properties of composites for personalized custom-fit implants (for more information, please see Chapter 14: BMP-assisted bone regeneration and applications in biofabrication, section Advanced Biomanufacturing).

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Acknowledgments This work was funded by the Restoration of Appearance and Function Trust (UK, registered charity number 299811) charitable funds and ARC Linkage Grant LP130100461.

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13.

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4

Rabah Gahoual1, Yannis-Nicolas Franc¸ois2, Nathalie Mignet1 and Pascal Houze´1,3 1 Department of Chemical and Biological Technologies for Health (UTCBS), CNRS UMR8258 - Inserm U1022, Faculty of Pharmacy, Paris Descartes University, Paris, France, 2 Laboratory of mass spectrometry of interactions and systems (LSMIS), CNRS UMR7140, University of Strasbourg, Strasbourg, France, 3Biochemistry department, University Hospital Necker-Enfants Malades, Public assistance - Paris Hospitals (AP-HP), Paris, France

4.1

Introduction

The development of biomaterials has demonstrated an unprecedented progress which has been emphasized by the introduction of a significant number of innovative materials. As such, the advent of biomaterials represents a formidable opportunity for the implementation of therapeutic treatments, able to induce tissue or organ regeneration [1]. Meanwhile, improved understanding of complex biological mechanisms and the enhancement of the production process of biological (macro)molecules using state-of-the-art biotechnologies have also experienced some major breakthroughs [2
4]. These progresses have led to their wide adoption for therapeutic purpose. Indeed, different types of proteins, such as monoclonal antibodies (mAbs) or fusion proteins, are currently approved for therapeutic use and the first biopharmaceutical product based on siRNA has been approved by the FDA on the third semester 2018 [5
8]. As a consequence, recent research articles describe the incorporation of biomolecules such as proteins to the structure of biomaterials in order, for instance, to deliver locally biological agents capable of favoring the regeneration of the tissue or improve the integration of the biomaterial [9,10]. Due to their inherent structural complexity and crucial implications, the incorporation of biomolecules to the material requires a dedicated characterization. The analysis of highly complex macromolecules, especially in complex matrices such as biological samples, is used to represent a challenge for the analytical sciences. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00017-9 © 2020 Elsevier Ltd. All rights reserved.

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A deeper understanding of the structure of biological macromolecules and the introduction of innovative instrumentation and methodologies, such as, for instance, -omics analysis, have allowed to alleviate the limitations of conventional analytical methodologies. Still, the characterization of biomolecules requires to be based on a multilevel analysis relying on complementary analytical methodologies, which provide different types of information. This chapter is focusing on providing insights on various cutting-edge analytical methodologies, such as high-resolution mass spectrometry (MS). These different analytical techniques can be implemented in order to study, on one hand, the influence of biomaterials incorporation to the structure of the biomolecules and on the other hand, to monitor the outcome of the therapeutic agent.

4.2

Analytical methodologies for protein identification and monitoring

In the context of the development of innovative biomaterials for tissue regeneration, the need for protein identification and/or monitoring arises from several aspects. Because of its exogenous origin, the in vivo presence of a biomaterial involves some major consequences especially on a systemic level [11]. As an illustration, the impact on the immune system is most prominent mainly due to its natural protecting function against foreign bodies [12]. A significant number of research articles are discussing the immunological reactions triggered by the exposure of biomaterials [13,14]. The biological processes involved in the immune response are regulated by the expression of different proteins such as tumor necrosis factor-α [15,16]. As a consequence, the possibility to identify in a specific manner this type of proteins and monitor their respective levels of expression can provide decisive information regarding the systemic impact of a biomaterial. The monitoring of various protein markers has been used to understand in a comprehensive way the implication of the biomaterials on the immune response in order to develop materials which could generate a reduced response and minimize the occurrence of side effects [17]. From another perspective a significant number of biomaterials recently developed for tissue regeneration rely on the incorporation of proteins to the structure of the biomaterial. Among applications the incorporation of proteins to the structure of biomaterials can be urged by further functionalization of the material [18], improved compatibility with the microenvironment [19], or controlled modulation of the immune response [10]. Thus in the most advanced applications, the biomaterial is gradually releasing the protein in order to generate the desired effect. It is therefore necessary to have analytical techniques capable of providing specific detection and absolute quantification of the incorporated proteins, in order to assess the material for the protein release capacity in addition to providing consistent data on the release profile.

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Regardless of the investigated aspect, the analytical methodologies require to provide an outstanding specificity in order to give the detection of a single protein present in complex biological samples for which available volume are commonly low. Similarly, the used techniques have to be particularly sensitive due to the low concentration of proteins generally present or released, which are analyzed in mixture with a large number of proteins, generally exhibiting significantly higher concentrations. Finally, in such a demanding context, the analytical strategies employed need to provide the analysis of several protein markers simultaneously. Over the last three decades a few analytical techniques have emerged and demonstrated to be particularly adapted to this type of analysis. As such, the major instrumental developments of flow cytometry (FC) have demonstrated its relevance to investigate in a comprehensive manner the immune response. MS is an analytical technique that enables the detection, the identification, and the quantification of the compounds constituting a sample by measuring their m/z ratio. The fundamental principle of MS consists of generating, from the studied molecule, ions in the gas phase by the intermediate of an ionization source. The ions produced are then separated based on their mass-to-charge (m/z) ratio and detected independently. Because the relation between the physical phenomenon used to separate the ions and the m/z ratio is known; it is possible to deduce the m/z ratio of an unknown molecule and, therefore, determine with an excellent accuracy its molecular mass. MS instruments can be virtually partitioned in four distinctive parts: the ionization source participates to the ionization and emissions of the compounds in the gas phase, the transfer line ensures the transport of the formed ions to the mass analyzer, the mass analyzer separates the compound in order to determine their m/z ratio, and the detector which produces an electrical signal handled by the data-treatment system (Fig. 4.1). MS is characterized by an outstanding specificity which allows to distinguish compounds exhibiting minor m/z ratio differences. In addition, MS is a particularly sensitive technique, with contemporary MS instruments able to detect compounds at subpicomolar (10212 mol/L) level. At the same time, the different types of MS instruments available make it compatible with the analysis of a large variety of molecules. Finally, the use of dedicated approaches, such as tandem MS (MS/MS), can allow to obtain to some extend further structural information regarding the analytes. Thereby, more than a simple detection method, MS is considered today as one of the major analytical techniques. Although is a centenary technique, MS analysis of polar molecules such as biomolecules was made possible relatively recently by the development of compatible ionization sources, electrospray ionization (ESI), and matrix-assisted laser desorption

Figure 4.1 Schematic representation of the different sections constituting MS instruments.

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ionization (MALDI). However, only introduced at the end of the 1980s, MS has demonstrated to be an outstanding tool for the analysis of a wide variety of biomolecules such as peptides and proteins. Commonly, a separation is often implemented prior to the MS analysis. Even though MS instruments are capable of detecting several compounds simultaneously and determine their respective m/z ratio if significantly different, the use of a separation method possesses different positive implications. Indeed, the preliminary fractionation of the compounds constituting a sample and their progressive transfer to the MS instrument is particularly relevant for the analysis of complex samples, giving the possibility to improve the number of compounds effectively detected and lowering the competition effect potentially occurring during the ionization process [20]. Liquid chromatography (LC) coupled with MS (LC
MS) is therefore used in most of the cases for the analysis of biological molecules such as nucleic acids, peptides, and proteins. When performing LC separation, the sample is carried inside a stationary phase by the intermediate of a constant flow of liquid phase referred to as mobile phase. The flow of mobile phase is created by a set of pumps able to deliver a flow rate accurately adjusted and extremely constant, even at high pressure in order to achieve efficient and reproducible separations. The stationary phase is constituted by a porous solid, densely packed in order to form a chromatographic column. Consequently to their introduction to the stationary phase, the different compounds constituting the sample are experiencing retention to the stationary phase. Because the retention is conditioned by the physicochemical characteristics of the compounds, they will exhibit significantly different retention times. Thus the components of a mixture will be eluted from the chromatographic column at different times and then gradually transferred to the MS instrument. The most common type of chromatographic stationary phase, referred as reverse phase, is composed of silica particles modified with alkyls groups (e.g. C18H37, C8H17) in order to make the media hydrophobic and apolar. However, various types of stationary phases are available in order to adapt the retention, which offers the possibility to change the selectivity of the separation. From an instrumental standing point, LC separation is realized on a routine basis mainly using high-performance LC (HPLC) instruments. Typically, HPLC is implementing stationary phases with particle size from 2 to 10 μm and a flow rate commonly comprised between 0.5 and 1.5 mL/min. More recently, miniaturized LC instruments, referred to as nanoLC, have been developed for routine analysis. Regardless of the format, LC
MS hyphenation can be quite straightforward, especially for ESI
MS instruments. Indeed, the mobile phase outlet can be directly connected to the sample introduction component of the MS system. Therefore the principal requirement is that the mobile phase is compatible with MS analysis and sufficiently volatile to be eliminated in the ESI source. Especially, it is the case for the mobile phases composed of H2O/ACN or H2O/ MeOH mixtures, typically used in reverse-phase chromatography. The advent of MS analysis applied to biological molecules has opened new horizons for the (bio)analytical chemistry field. Consequently, LC and MS instruments have benefited from major technical improvements which have allowed to constantly improve the sensitivity and level of characterization achieved using LC
MS

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analysis. The performances of contemporary LC
MS instrumentations are now capable of providing a tremendous amount of data in a single LC
MS analysis which gives the possibility to achieve in-depth characterization of highly complex biological samples enabling the emergence of MS-based proteomic analysis.

4.3

Mass spectrometry
based proteomic analysis

The terminology proteome designates a complete set of proteins present in an organism or in a compartment (e.g., mitochondria and membrane), at a given time and for a given physiological state. The analysis of the proteome, commonly mentioned as proteomic analysis, consists of the systematic identification composing a biological sample which can be completed with quantitative information. This analytical workflow has been developed consequently to large-scale genomics in order to provide a detailed analysis regarding the diversity of the proteins [21] in an extended manner. Indeed, proteins are resulting from the transcription of the DNA-coded information. However, the evolution of the protein expression levels, depending for instance on the state, is not enclosed in the genetic heritage. It is then necessary to perform the analysis on the protein level in order to attribute minute changes in the expression levels. Similarly, proteins are commonly experiencing endogenous chemical modifications described as posttranslational modifications (PTMs) [22]. Notably, some modifications have been attributed to be important in biological signaling like deamidation [23,24] or phosphorylation [25]. In this case as well, genomic analysis does not provide information regarding that aspect therefore requiring a direct characterization of the protein. Originally, proteomic analyses were mainly performed using bidimensional differential gel electrophoresis (2D-DIGE). Due to their chemical nature, proteins are widely diverse in terms of size and they also bear a significant number of charged functions in solution. As a consequence, the separation provided by gel electrophoresis demonstrated to be particularly adapted to proteins with the possibility to achieve high resolution even for complex biological samples [26]. However, 2DDIGE requires a specific experimental expertise. Also, this analytical approach does not provide unambiguous identification of the proteins and additional experiments are needed in order to obtain a structural identification of the separated proteins. The introduction of MS instrumentation adapted to the analysis of biomolecules has modified in a drastic manner the proteomic analysis [27]. Indeed, MS possesses the interest to provide an outstanding sensitivity as well as specificity, which is particularly relevant for the analysis of complex samples. In addition, it can be coupled in a straightforward manner to LC in order to implement a separation prior to the detection. In state-of-the-art proteomic analysis or bottom-up proteomics, the protein mixture composing the sample is undergoing proteolytic digestion using endoprotease enzymes. The complex mixture of peptides is then separated and analyzed by LC hyphenated to MS/MS (LC
MS/MS) as emphasized in Fig. 4.2. The peptides are

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Figure 4.2 Schematic representation of a generic MS-based proteomics experiments, including proteolytic digestion, to generate the peptide mixture which is then analyzed by LC
MS/MS to obtain high-resolution MS measurement of the peptide in conjunction to the fragmentation spectrum. LC
MS/MS, Liquid chromatography hyphenated to tandem mass spectrometry; MS, mass spectrometry. Source: Adapted with permission from Switzar L, Giera M, Niessen WMA. Protein digestion: an overview of the available techniques and recent developments. J Proteome Res 2013;12(3):1067
77 [28]. ©2013 American Chemical Society.

then gradually transferred to the MS instrument where their m/z ratio is measured followed by their fragmentation in the gas phase in order to measure the m/z of the fragments (Fig. 4.2). The conjunction of these two information is used to identify unambiguously the different peptides by direct homology against a protein sequence database used to generate theoretical peptides in silico by the intermediate of a search algorithm [29
31]. MS instruments have experienced some major improvements over the last two decades, especially regarding to the MS accuracy, resolution, sensitivity, and acquisition frequencies. These improvements have allowed to push further the level of characterization achieved using proteomic analysis. As a direct consequence, it is possible to perform the proteomic analysis on complete protein extracts without any type of prefractionation, which is commonly referred to as shotgun proteomics, just by performing directly a complete digestion of the entire sample content [32]. Therefore the proteomic analysis has over the recent years demonstrated to be particularly powerful with the possibility using state-ofthe-art MS instrumentation to identify in a single analysis more than 10,000 proteins using an injected quantity of sample generally corresponding inferior to 1 μg of protein extract [33,34]. Also, this analytical workflow is compatible with the incorporation of quantification elements in order to monitor the evolution of the expression level of the identified proteins [35
37]. The proteomic analysis workflow is partitioned into three consecutive phases which requires critical optimization depending on the sample nature and complexity to enable a relevant characterization: sample preparation, (nano)LC
MS/MS analysis of the peptide mixture, and finally data treatment.

4.3.1 Sample preparation for proteomics experiments Proteomic analysis is based on peptide characterization in order to lead to the identification of the protein composing the sample. From an MS standing point the analysis of peptides is more favorable due to significantly higher ionization efficiencies of the ESI source interfacing the LC system and the MS system, compared to intact proteins. Also, because of their reduced size, the fragmentation of peptides requires

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less energy therefore yielding an important number of specific fragments, which enables a systematic and high-fidelity peptide identification from fragmentation spectra. To generate the peptide mixture from samples mainly composed of proteins, the sample preparation designed for proteomic sample relies on proteolytic digestion using different types of proteases. Proteases are enzymes whose biological activity involves the cleavage of the peptide backbone of proteins [38]. Prior to proteolytic digestion, proteins are commonly undergoing a reduction process in order to cleave disulfide bridges between cysteine residues and therefore destabilize the tertiary/ quaternary structure of the protein. Different reagents can be used to perform the disruption of disulfide bridges such as ß-mercaptoethanol or Tris (2-carboxyethyl) phosphine
HCl [39] however dithiothreitol (DTT) represents by far the most commonly adopted alternative (Fig. 4.3A) [40]. That observation is explained by the efficiency of the reduction, the favorable kinetics allowing the reduction to be realized rapidly in addition to minor interferences with the downstream (nano)LC
MS/ MS analysis. Because the reduction is a reversible process, the free thiols are submitted to alkylation in order prevent the reformation of the disulfide bonds [41]. Alkylation can be realized using various types of alkylation agents which covalently bond the thiols present on reduced cysteine residues. One can cite iodoacetic acid or N-(2-aminoethyl)maleimide, still the alkylation agent most widely implemented is iodoacetamide (IAM). IAM is a highly reactive alkylation reagent capable of performing the reaction in less than 1 hour in a reproducible manner (Fig. 4.3B) [42,43].

Figure 4.3 Schematic representation of (A) the reaction of the reduction of disulfide bridges between two cysteine residues using DTT and (B) alkylation by the intermediate of IAM. DTT, Dithiothreitol; IAM, iodoacetamide.

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As mentioned, the principal objective of the reduction
alkylation process is to destabilize the structure of the protein for the enzyme to have access to all the cleavage sites present on the protein backbone and therefore maximize the yield of the digestion reaction. The proteolytic enzymes used for digestion in proteomic experiments have the particularity to provide specific digestion of the protein meaning that cleavages occur on specific amino acid residues. That aspect is particularly important to predict in silico the set of peptides generated from the digestion of proteins which enables correlation with the proteins listed in the database. To achieve that purpose, different types of proteases are available exhibiting their respective specificities [43]. Still, trypsin represents the enzyme most commonly used. Trypsin is cleaving specifically the peptide backbone in C-termini of lysine (Lys) and arginine (Arg) residues [44]. Due to the occurrence probabilities of these types of amino acid residues, peptides generated exhibit in the typical case, molecular masses compatible with high resolution MS measurements and concomitant fragmentation in the gas phase to enable MS/MS characterization. In order to provide an optimal activity of the enzyme, the digestion is performed in a dedicated buffer, commonly (NH4)HCO3 at a concentration of 50 mM (pH 8.0), whereas trypsin is also active at physiological pH but irreversibly inactivated in pH conditions below 4 [45]. The digestion process is then performed by incubation at room temperature for up to 12 hours depending on the nature of the proteins and the complexity of the sample. The reaction can be quenched by addition of trifluoroacetic acid (TFA) or preferably formic acid (FA). Indeed, TFA was determined to be responsible of additional background noise due to its ionization during the ESI process [46]. The sample is finally diluted to the desired concentration using the LC mobile phase.

4.3.2 Peptide mixture analysis by liquid chromatography coupled to tandem mass spectrometry Consequently, to the digestion of the sample, the mixture of peptide obtained is separated and analyzed by (nano)LC hyphenated to high-resolution MS/MS (nanoLC
MS/MS). Indubitably, the development of nanoLC has been urged by the emergence of proteomic applications. The implementation of LC in a nanometric format allows, on one hand, to achieve high separation efficiencies compared to their analytical counter parts by the intermediate of stationary phases composed of smaller particle sizes [47]. This aspect is interesting with respect to the tremendous complexity of proteomic samples. On the other hand, nanoLC involves the use of lower mobile phase flow rates, typically below 500 nL/min which is particularly favorable to the ESI process, especially due to reduced interferences originating from the mobile phase such as ion suppression effects [48]. It is important to note that nanoLC requires the use of dedicated instrumentation able to deliver accurately controlled and robust nano flow rates. Especially, the pumps used for the circulation of mobile phase require to deliver a flow rate inferior to 500 nL/min in an accurate and constant manner. Therefore conventional HPLC pumps cannot be used anymore. Thereby, higher sensitivity is achieved using nano flow rates [49],

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which offers the possibility to achieve a relevant analysis even if the quantity of sample generally available is fairly limited. In addition, this characteristic, supported by the excellent sensitivity of MS instrumentation, allows the identification of lowly abundant proteins [50]. The separation of the peptides composing the digested sample is performed using reverse phase nanoLC basically using a C18-based stationary phase and a mobile phase composed of a mixture of water and acetonitrile containing generally 0.1% FA operated using a conventional gradient [43]. Because of the volatility of its components, this mobile phase system is rather compatible with the ESI process by allowing an efficient elimination of the solvent. The gradient time is adapted to the complexity of the sample; thus gradient time is usually 60 minutes long however it can be extended to several hours in the case of highly complex samples (Fig. 4.4). An extended gradient allows a gradual transfer to the MS of the different peptides composing the sample in order to enhance the number of peptide effectively detected [52]. Following their separation, the peptides are directly transferred to the MS instrument commonly by the intermediate of an ESI source. The instrument is automatically operated in a sequential manner in order to provide the measurement of the m/z ratio of the digested peptides (MS analysis) followed by fragmentation in the gas phase, using collision-induced dissociation (CID). The m/z values corresponding to the fragments generated from the selected peptide are also measured in a fashion conventionally referred to as MS/MS. CID process relies on the collision of the selected parent ion with inert gas molecules such as N2 or He to generate the fragmentation. This fragmentation technique has proven to be particularly relevant for the sequencing of peptide, yielding mainly b- and y-ions, in a consistent manner (Fig. 4.5) [54].

Figure 4.4 Base peak chromatogram achieved from nanoLC
MS/MS analysis of 10 ng tryptic digest of Shewanella oneidensis. nanoLC
MS/MS, (Nano)liquid chromatography hyphenated to high-resolution tandem mass spectrometry. Source: Adapted with permission from Zhu Y, Zhao R, Piehowski PD, Moore RJ, Lim S, Orphan VJ, et al. Subnanogram proteomics: impact of LC column selection, MS instrumentation and data analysis strategy on proteome coverage for trace samples. Int J Mass Spectrom 2018;427:4
10 [51]. ©2018 Elsevier.

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(A)

(B)

Collision gas

X3 Y3 Z3 X2 Y2 Z2 X1 Y1 Z1

Guiding quadrupole

R1 O Fragment ions

Parent ions

Ion fragmentation

Fragment ion Neutral loss

H2N

C C

R2 O N H

C H

C

R3 O N H

C H

C

R4 N H

C H

COOH

A1 B1 C1 A2 B2 C2 A3 B3 C3

Figure 4.5 (A) schematic representation of a CID collision cell showing the fragmentation performed in MS/MS analysis. (B) Representation of the nomenclature conventionally used to describe peptide fragmentation in MS. Each amino acid is susceptible to generate different fragment ions depending on the cleaved bond mainly driven by the fragmentation mode implemented. CID, Collision-induced dissociation; MS, mass spectrometry; MS/MS, tandem mass spectrometry. Source: Partially adapted with permission from Roepstorff P, Fohlman J. Letter to the editors. Biomed Mass Spectrom 1984;11(11):601 [53] (panel B). ©1984 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

Partially because of the complexity of the sample, the instrument is operated in data-dependent acquisition. Using this mode of acquisition, the instrument automatically detects the elution of a compound based on the intensity of the MS signal, regardless of the m/z value measured. Then, the parent ion is selected to undergo subsequent fragmentation in an automatic fashion. This strategy allows to perform the analysis without any a priori knowledge regarding its composition. In addition, the initial m/z ratio measured in MS for the parent ion is completely linked to the fragmentation spectra obtained. The analysis is performed using high resolution hybrid MS instrumentation mainly based on quadrupole time-of-flight or orbitrap mass analyzers. These types of instruments are providing high resolution and high mass accuracies and offer the possibility to perform MS and MS/MS experiments (Fig. 4.6). Thus thanks to their high resolution power they are capable of resolving the isotopic profile of complex macromolecules even in the case of peptides of a few kDa (Fig. 4.7A). The plethora of peptides which can be generated from the digestion of the proteins contained in the sample, high resolution MS is required in order to prevent any misidentification. As emphasized in Fig. 4.7, the correlation between the m/z measurements of the entire peptide in concomitance with the fragmentation requires to be in complete agreement in order to enable the positive identification of the peptide.

4.3.3 Protein identification Because of the large amount of data generated from a single experiment, data analysis tends to represent the major bottleneck of the proteomic analysis. To tackle this limitation, different research groups have strengthened their expertise with bioinformatics solutions capable of providing automated identification of proteins

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Figure 4.6 Schematic diagram of (A) MS/MS instrumentation based on orbitrap mass analyzer (LTQ orbitrap, Thermo Fischer Scientific) and (B) MS/MS instrumentation based on hybrid quadrupole time-of-flight mass analyzer (Agilent 6560 IM Q-TOF, Agilent Technologies). MS/MS, Tandem mass spectrometry. Source: Adapted with permissions from Scigelova M, Makarov A. Orbitrap mass analyzer
overview and applications in proteomics. Proteomics 2006;6(S2):16
21 [55] and Kurulugama RT, Darland E, Kuhlmann F, Stafford G, Fjeldsted J. Evaluation of drift gas selection in complex sample analyses using a high performance drift tube ion mobilityQTOF mass spectrometer. Analyst 2015;140(20):6834
44 [56]. ©2006 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim and ©2015 Royal Society of Chemistry, respectively.

using MS/MS datasets and integrate them in a user-friendly workflow. Different solutions are therefore available, either commercial or provided as open-source software [57
60]. Basically, software identification relies in the use of a search algorithm. Using a database composed of proteins amino acid sequence, the algorithm generated in silico a set of theoretical fragmentation spectra depending on the proteins present in the database and also the type of proteolytic enzyme used. This set of spectra is compared to the experimental data obtained from nanoLC
MS/MS experiments. The matching spectra enable the identification of the peptides by homology with the theoretical MS/MS generated from the database. The identification of the

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Figure 4.7 Spectra obtained from the Ultra-High Pressure liquid chromatography coupled to tandem mass spectrometry (UHPLC-MS/MS) analysis of a tryptic peptide generated from human serum albumin illustrating (A) MS measurement of the peptide correlated with (B) the corresponding MS/MS spectra enabling the identification of the peptide exhibiting the amino acid sequence AWAVAR. MS, Mass spectrometry; MS/MS, tandem mass spectrometry.

peptides allows to backtrack their origin in the database which ultimately provides the identification of the proteins present in the sample. Thanks to the resolution and mass accuracy provided by high resolution MS, the search algorithm is able to apply harsh restriction, such as the implementation of decoy protein sequences, in order to minimize as much as possible false positives. The search algorithms available also systematically include a scoring system based on different criteria such as the number of fragment identified [61]. The specificity of the peptide identified for the protein is also a key criterion to avoid misinterpretations. Still, the user is able to manually interpret MS/MS data if necessary. The use of bioinformatics tools has enabled high-throughput and confident identification of proteins using nanoLC
MS/MS large datasets obtained from proteomic experiments. Their large adoption, in conjunction to major instrumental developments, has contributed to explore further extensively complex protein samples. They clearly participated in the establishment and application on a larger scale of proteomic analysis.

4.3.4 Applications of proteomic analysis to the development of biomaterials for tissue regeneration Proteomic analysis can supply a large amount of information regarding the protein content of samples which may be exploited in different manners. Primarily, data obtained from proteomic experiments can be used to identify the protein diversity composing a biological sample [62,63]. Also, by adaptation of the sample purification, it gives the possibility to focus the investigation on a specific compartment

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such as mitochondria or exosomes [64
66]. Historically, this analytical methodology has been used for biomarker research by comparison of different individuals and conditions [67,68]. Quantitative information may also be obtained from experiments in order to interpret changes in protein expression levels [69,70]. The quantitative proteomic approach has been applied for instance to evaluate the influence of environment on various types of organism [71
73]. Conventional MS analysis has been previously applied to the characterization of biomaterials in order to obtain important information regarding their intrinsic structure. As an example, time-of-flight secondary ion MS (TOF-SIMS) has been implemented to perform surface analysis in the form of imaging experiments. Using TOF-SIMS imaging, Belu et al. achieve the precise characterization of the drug metoprolol encapsulated in ethylcellulose. The TOF-SIMS imaging experiments demonstrated the correct encapsulation of the drug, the formation of the ethylcellulose coating in addition to show the silica core of the particles [74]. To provide this type of characterization, the different compounds are previously analyzed independently in order to determine specific m/z signal they are respectively exhibiting. The biomaterial sample characterized by multiple TOF-SIMS experiments in a scanning fashion in order to determine the localization of the different compounds in presence and form an image of the object. The same group used a similar analytical approach in order to study the structure of drug bead systems of different encapsulated in amylose or Eudragit L30-D [75]. However, proteomic analysis is currently emerging as a complementary analytical methodology capable of providing information regarding the impact of the contact of the biomaterial with the organism or after implantation. Thereby, proteomic analysis gives the opportunity to attribute subtle changes in protein expression levels consequently to the exposition to the biomaterial or to monitor the proteins involved in tissue regeneration. Thus Kuo et al. used systematic proteomic analysis in order to improve the mesenchymal stem cells (MSCs) expansion and differentiation in bone marrow and Wharton jelly. In order to identify the proteins involved in these biological process, bone marrow and Wharton jelly samples were analyzed using differential 2D-DIGE to determine the proteins exhibiting expression levels significantly different between bone marrow and Wharton jelly. In total, 32 proteins demonstrated significant differences in expression levels could be selected using this approach. The proteins were then further characterized by MALDI
MS/MS analysis to obtain their unambiguous identification. Finally, the expression of the identified was knockdown in order to study their influence of the tissue regeneration. Results showed that lowering the expression of transgelin-2 allowed to increase the growth rate of MSCs. Also, knockdown of HSP90β had the effect to increase bone nodule formation and stopped cell growth [76]. MS-based proteomic analysis was also extensively used to characterize the extracellular matrices (ECM) [77]. Therefore Naba et al. studied ECM proteins originating from pancreatic islets, during angiogenic switch and subsequent insulinoma development. In this study the pancreatic islets samples were first undergoing decellularization in order to fractionate the proteins

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originating from ECM. The samples were then prepared by performing reduction/ alkylation followed by tryptic digestion. The peptide mixtures generated from the sample preparation were then separated and analyzed by nanoLC
MS/MS analysis. Because the acquisition rate of typical high resolution MS instrument is important enough, the nanoLC
MS/MS experiments allowed to provide a sensitive analysis of this sample composed of a mixture of several thousands of peptides. MS/MS data interpretation enabled the confident identification of the protein composing the ECM samples. Thereby, 120 ECM proteins could be successfully identified using nanoLC
MS/MS data. Consequently to the identification the pancreatic ECM proteins, the MS signal intensity corresponding to the specific m/z ratio of their peptides was monitored throughout their development (before/after angiogenic switch and as insulinomas). The results exposed show that 35 proteins were exhibiting significantly different expression levels during the development. This study clearly demonstrates the possibility to use MS-based proteomic analysis to monitor the expression levels of proteins in different conditions. Also, the authors suggest the expression levels of the proteins highlighted from the nanoLC
MS/MS proteomic analysis may be influenced to improve tissue regeneration in the context of diabetes for instance [78]. In a recent research article, Garcia-Puig et al. implemented a similar analytical workflow to investigate the protein content and expression levels of zebrafish ECM during heart regeneration. To enable the proteomic analysis of ECM, the heart samples were first undergoing decellularization with the aim of preventing contamination by protein originating from the intracellular medium. Consequently to this preparation, the samples were analyzed by nanoLC
MS/MS using a high-resolution MS instrument. MS/MS data were used to identify unambiguously the proteins composing the sample, and the intensity of the MS signal was considered to estimate the expression levels on the different proteins. The nanoLC
MS/MS experiments performed on control samples containing both ECM and intracellular medium, enable to successfully identify 447 different proteins. After sample decellularization, 36 proteins strictly belonging to ECM were identified. By considering the MS signal, they could attribute during tissue regeneration, the increase of the expression levels of several types fibrinogen, fibronectin 1b and periostin b. At the same time, collagen and fibrillin 2b levels were significantly lowered [79]. Also, MS/MS spectra providing the fragmentation of the identified peptides can enable the specific identification of peptide backbone PTMs which are involved in biological regulation and signaling [80
82]. Therefore proteomic experiments can be implemented in conjunction with complementary analytical approaches such as genomics and FC in order to obtain a detailed understanding regarding the implication underlying the implantation of biomaterials [83]. The information provided by proteomic experiments can provide decisive information regarding implantation outcome which can help to tailor the development of innovative biomaterials or reject early materials with potential to generate important adverse effects [84]. Note, conventional MS analysis was previously used to investigate the nature of the protein naturally adsorbing onto the surface of biomaterials upon exposure.

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Thus Castner et al. have developed a method using TOF-SIMS analysis in order to identify adsorbed proteins after incubation of 2 hours in 1% bovine serum in the case of mica, polytetrafluoroethylene (PTFE) and silicon wafer. Results obtained have demonstrated the identification of several types of proteins, including serum albumin, fibronectin, fibrinogens, immunoglobulin G, and γ-globulins. As expected, they correspond to the most abundant proteins in the serums; nevertheless, the article shows the possibility using this analytical technique to identify unambiguously the adsorbed proteins [85,86]. Still, the literature regarding natural protein adsorption on biomaterials is scarce most likely due to the frequent introduction of novel media. However, the characteristics and performances concerning sensitivity and specificity of state-of-the-art proteomic analysis open up a new pathway for the comprehensive identification of proteins adsorbed to the biomaterial surface. In parallel, because recent research articles describe the development of biomaterials incorporating proteins into their structures, like growth factors or cytokines able to modulate the immune response, for their release after implantation [9]. In this case as well, recent researches show that the implementation of proteomic analysis is particularly interesting to monitor the release of the incorporated proteins. Indeed, because the identification of the protein is strictly based on high-resolution MS analysis of the generated proteolytic peptides, the identification is extremely specific. Therefore the proteomic analytical workflow is capable of identifying without ambiguity a single protein in a mixture containing several hundreds. As an illustration, Barthes et al. developed a biomaterial composed of gelatin cross-linked using transglutaminase which can be used to mimic ECM and enable cell proliferation. In order to enable cell development even in the absence of serum, the medium was loaded with several serum components such as albumin, serotransferrin in addition to different growth factors like epidermal growth factor (EGF) and fibroblast growth factor (FGF). To investigate the gradual release of the proteins incorporated to the media, gel samples were incubated in phosphate buffered saline (PBS) solution. The supernatants were then collected and submitted to conventional proteomic sample preparation—reduction/alkylation followed by overnight tryptic digestion. The peptide mixtures obtained from the sample preparation were then characterized by bottom-up proteomic analysis. Thus the interpretation of the MS/MS spectra allowed the identification of the proteins initially incorporated to the gelatin-based media. Also the results achieved from the proteomic analysis suggest their rapid release from the medium [87]. The proteomic analytical methodology based on high-resolution MS appears as a major technique to study the nature and expression levels or proteins composing biological samples. As a consequence, the large amount of information obtained from this type of analysis appears particularly attracting to improve the knowledge regarding biomaterials impact. Because proteomic is a demanding analysis from an experimental point of view which needs extensive optimization especially in regard to the matrix, further improvements can be expected in the upcoming future to improve the suitability between samples obtained from biomaterial implantation and the proteomic characterization.

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Biomaterials for Organ and Tissue Regeneration

Analytical methodologies adapted to protein structural characterization

Consequently to their introduction, the first generation of biomaterials was designed and produced with the conceptualization to represent inert supports which could sustain prolonged contact with organic tissues while having minimum biological implications on the host organism. In the recent years the research concerning the development of innovative biomaterials has demonstrated to be particularly prolific which progressively led to a shifting of that definition. As the knowledge regarding biomaterials progresses, they nowadays tend to represent complex and organized media which may incorporate into their structure various components, including proteins, implemented to be released in the direct surrounding environment. The incorporation of proteins to the structure of biomaterials is primarily driven to influence the interaction between the biomaterial and tissues [18]. The addition of proteins to biomaterials gives the opportunity for further functionalization. As an illustration, the integration of different types of cytokine appears as an elegant means to modulate the immune response after implantation, in order to limit inflammation or other adverse effects [88]. In some applications the proteins can be used on the contrary to stimulate the immune response in a localized region [89]. Different types of proteins have been successfully integrated to the structure of biomaterials, including growth factor such as vascular endothelial growth factor [90] and bone morphogenetic protein 2 [91], heparin [92], mAb [93], and granulocyte
macrophage colony-stimulating factor [94]. In the context of additional functionalization of biomaterials by protein integration, the prolonged contact between the protein and the biomaterial may influence in a dramatic manner the structure of the protein. Modification of the structure of the protein can result in various effects such as lowered or inhibited biological activity to, in the worst case, triggering of the immune response [95,96]. Thereby, it is crucial to assess the structural integrity of the protein upon contact with the biomaterial to ensure an unaltered biological activity. The stability of proteins can be compromised through a variety of processes. For instance, protein aggregation is attributed to partial or complete denaturation of the protein, exposing large part of the peptide backbone to the environment. Aggregation is then observed due to electrostatic and hydrophobic interactions, which primarily induces premature clearing [97]. Therefore protein aggregation due to the incorporation inside the structure of the biomaterial could result in elimination and subsequent altered biological activity. In the case of biopharmaceutical products for example, aggregation has been described as a critical quality attribute (CQA) by the regulation agencies like the US Food and Drug Administration (FDA) or the European Medicines Agency (EMA) Therefore, the approval of this category of therapeutics is conditioned by a detailed characterization concerning that aspect is mandatory [98]. Proteins incorporated to biomaterials may also undergo PTMs. PTMs describe chemical modifications of amino acid residues which may be endogenous like environmental conditions or triggered from the organism to induce signaling [99
101]. Proteins

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may exhibit a wide variety of PTMs such as methionine oxidation or asparagine deamidation (Fig. 4.8), principally depending on the type and localization of amino acid residue on the peptide backbone [102]. In the case of incorporation to the structure of biomaterials, the occurrence of PTMs may influence significantly the biological activity of the released protein potentially exhibiting altered efficacy or rapid elimination [103]. As an illustration, for biopharmaceutical products currently approved for therapeutic use such as mAbs, regulation agencies have positioned as CQA a complete set of PTMs hotspots which differs depending on the nature of the mAbs. Therefore these PTMs require mandatory characterization due to their known impact on the activity and pharmacological activity of the mAb, in order to envisage approval [104
106]. As a consequence, there is an increasing need for structural characterization of proteins incorporated to biomaterials to demonstrate in a comprehensive way the biological functionality upon release from the media to the direct environment of implants. In addition, investigations regarding the conservation of the structure of proteins released from biomaterials can provide complementary data supporting the beneficial effect of this type of approach. Protein represents highly complex macromolecules potentially with a wide range of micro heterogeneities. Due to their inherent complexity, the structural characterization of proteins requires dedicated analytical methodologies. Thus some major improvements both from technical

Figure 4.8 Schematic representation of posttranslational modifications (A) methionine oxidation and (B) asparagine deamidation.

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developments and methodological strategies have been recently described in the literature in order to address the complexity of the structural characterization of biological macromolecule-like peptides, proteins, or carbohydrates [107,108]. Consequently to the introduction of ionization techniques enabling the analysis of biological macromolecules [27], MS has rapidly demonstrated its ability to play a pivotal role in the characterization of proteins over the different levels composing their structure [30]. Thereby, MS can be used to analyze the primary structure of proteins [109] as well as secondary [110] and higher order structure of proteins [111]. This is explained by the powerful specificity delivered by contemporary MS instrumentation as well as relevant sensitivity, in addition to the possibility to obtain detailed structural information using dedicated workflows. The PTMs potentially affecting proteins are concerning specific amino acid residues leading to their respective outcome. The analysis requires then to focus on the primary structure of proteins in order to finely localize the position and specifically monitor the occurrence of PTMs over the entire peptide backbone. Therefore the characterization of PTMs is performed by the intermediate of a peptide-centric MS/ MS analysis. The analytical workflow used for the characterization of PTMs is mainly derived from bottom-up proteomics analysis with partially differing constraints and objectives. Prior to the analysis, the sample containing the protein of interest is undergoing digestion using proteolytic enzymes in order to generate a mixture of peptides. Afterward, the peptide mixture obtained is separated and characterized by different types of MS/MS based experiments.

4.4.1 Sample preparation Similarly to bottom-up proteomic experiments, sample preparation represents a key step to achieve a highly detailed characterization of the primary structure of proteins. This is primarily achieved by an optimal digestion of the peptide backbone. Therefore prior to the enzymatic digestion, protein disulfide bridges are reduced, conventionally using DTT and the reactivity of free thiols is inhibited through alkylation mainly using IAM. The use of these chemical reagents proved through numerous research article the possibility to perform the reduction
alkylation process of disulfide bridge in a rapid and robust manner. Afterward the protein is subjected to proteolytic digestion which is commonly performed using trypsin. However, unlike bottom-up proteomic experiments, the structural characterization requires to achieve complete sequence coverage of the peptide backbone in order to have access to the different sites potentially present on the protein. As a consequence, the use of different proteolytic enzymes such as serine proteases (GluC or LysC), metalloprotease (AspN), or chymotrypsin can be implemented due to their respective digestion specificity in order to generate peptide mixtures favorable with the subsequent MS/MS experiment and extended sequence coverage [43]. When the buffer conditions used for the digestion are similar, several proteolytic enzymes can be used in conjunction to induce further digestion of the protein backbone [112
114]. Alternatively, the sample can be submitted in parallel to different proteolytic digestion which MS/MS analysis results can be crossed in order to

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maximize the coverage of the peptide backbone. Nevertheless, in the context of structural characterization, sample preparation should be thoroughly designed to maximize coverage depending on the respective amino acid sequence of the studied protein. In addition, because the sample preparation required is particularly extensive, it is of utter importance to optimize the reaction conditions in order to limit as much as possible the occurrence of PTMs originating from the sample preparation. Indeed as an illustration, the alkaline conditions used to commonly perform tryptic digestion are described to induce endogenous deamidation of the protein [115]. Still, careful optimization of the sample preparation protocol allows to limit that effect [116]. Also, the inclusion of a control enables to easily alleviate the production of endogenous modifications due to the sample preparation.

4.4.2 Tandem mass spectrometry
based analysis of posttranslational modifications The protein mixture generated from proteolytic digestion sample preparation is afterward separated and characterized using MS. The analysis is conventionally performed using nanoLC
MS/MS instrumentation implementing C18 reverse stationary phase which porosity has been optimized for the separation of peptides. In the case of sample available in readily quantities, the analysis can be performed using ultra-high pressure liquid chromatography coupled to tandem mass spectrometry (UHPLC-MS/MS) instrumentation. Due to the nature of the stationary phase and the direct hyphenation with the mass spectrometer, the mobile phase used to perform the LC is generally composed of a mixture of water and acetonitrile containing 0.1% FA [117]. Compared to shotgun proteomic experiments described previously, the initial protein complexity regarding the number of proteins is expected to be reduced when performing structural characterization which explains that short-time gradient can be implemented without compromising on the MS/MS characterization efficiency. However, the separation conditions should be carefully optimized in order to enable estimation in a reliable manner of the level of modification for each PTMs site. Thereby, the chromatographic separation should enable a systematic separation of the unmodified peptide from its counterpart experiencing the modification of one amino acid residue. That condition allows to prevent misinterpretation due to coelutions in addition to cancel the occurrence of competition effects during the ESI which could result in biased estimation of the levels of modification [118]. More recently, capillary zone electrophoresis coupled with MS/MS (CZE
MS/ MS) was proposed as an alternative for the characterization of the primary structure of proteins, including highly comprehensive analysis of PTMs [119]. Indeed, the separation mechanism provided by the electrokinetic separation demonstrated to be particularly suitable for this type of analysis, with the possibility to achieve systematically a complete sequence coverage from tryptic digests in a reproducible manner. This characteristic is attributed to the ability of the electrophoretic separation to separate and transfer peptides to the MS instrument in a robust way regardless of

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their chemical composition [109]. The selectivity delivered by CZE demonstrated the ability to systematically separate peptides exhibiting different types of PTMs from their intact homologues therefore enabling a site-specific and confident evaluation of the level of the protein modification. Thereby, that characteristic could be demonstrated in the case of N-terminal glutamic acid cyclization (pyroGlu), asparagine deamidation (deaN), methionine oxidation (oxiMet), and aspartic acid isomerization (isoD) (Fig. 4.9) which are described as naturally occurring PTMs [22,121]. For the later one, note that isomerization of aspartic acid does not imply a mass difference between the modified peptide and the corresponding intact one which without separation could then not be attributed based on the MS analysis alone. However, because CZE
MS/MS demonstrated the systematic separation of the peptides experiencing aspartic acid isomerization, the two respective forms could be successfully identified meaning the electrophoretic separation implemented in this particular case allowed to further enrich the level of characterization achieved regarding the analysis of this challenging PTM [120]. Finally, the separation of peptides exhibiting PTMs from their intact counterpart was maintained regardless of the global amino acid sequence of the peptide. The performance of CZE
MS/MS analysis appears particularly interesting for the characterization of PTMs because, in order to provide an optimal estimation of the levels of modification, separation of the modified PTMs is mandatory [122]. Consequently to their chromatographic or electrophoretic separation, peptides are gradually transferred to the MS instrument. The instrument performs the

Figure 4.9 Extracted ion electropherograms of m/z values corresponding to HT-22 synthetic tryptic peptide (seq. FNWYVDGVEVHNAK) showing separation of (peak 1) unmodified HT-22, (peak 2) HT-22 exhibiting an asparagine deamidation in position 12 (D6/deaN12), (peak 3) HT-22 with isomerization of the aspartic acid residue in position 6 (isoD6/N12), and (peak 4) HT-22 affected by both PTMs (isoD6/deaN12). Source: Adapted with permission from [120]. ©2016 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

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measurement of the m/z ratio of the compounds eluting from the separation system (MS analysis) which is directly followed by the selection and fragmentation of the parent ion corresponding to the detected peptide. The m/z ratios of the fragment ions are then measured (MS/MS analysis). MS/ MS analysis is required in order to precisely position PTMs on the peptide backbone. To induce the fragmentation, CID is principally used. This is due to its wide availability in MS instrumentation in addition to the robust and relevant fragmentation obtained in the case of peptides using CID. However, electron transfer dissociation, which in particular has the capacity to reach higher fragmentation energies than CID, demonstrated to be interesting for the characterization of PTMs as well [123]. The identification of the peptide is achieved on the concomitant correlation between the high-resolution MS measurement of the peptide provided by MS analysis in addition to the fragmentation spectra obtained from MS/MS analysis (Fig. 4.7). Thus because CID fragmentation yields mainly b- and y-ions (Fig. 4.7B), therefore, the amino acid sequencing of the peptides can be precisely characterized by tracing the both series of ions (Fig. 4.10A). In most cases the occurrence of PTMs is involving a mass shift compared to the equivalent peptide without modification like C-terminal glutamic acid cyclization (217.0265 Da), methionine

Figure 4.10 Example of MS/MS spectra of a tryptic peptide (IYPTNGYTR) showing the identification (A) without modification and (B) with a deamidation on Asn55 (N ) illustrating the partial modification of the protein. MS/MS, Tandem mass spectrometry. Source: Adapted with permission from [109]. ©2013 Taylor and Francis publishing.

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oxidation (115.9995 Da), or asparagine deamidation (10.9840 Da). Because CID provides a gradual fragmentation of the peptide backbone, the mass increment corresponding to a modified amino acid residue present will be equal to the mass of the concerned amino acid plus/minus the mass shift of the PTMs as emphasized in Fig. 4.10. Therefore the interpretation of MS/MS spectra provides unambiguous information regarding the nature of the PTMs alongside to the precise attribution of the amino acid residue affected even if several amino acids of the same type are present on the peptide. Complementarily, the identification of the peptide allows to localize precisely the modification on the peptide backbone of the protein which enables to correlate its impact on the biological activity of the protein. Sequence coverage is therefore primordial in that aspect in order to map in an extensive manner the occurrence of PTMs over the entirety of the peptide backbone of the protein. In particular, sample preparation and analytical conditions, including MS parameters, require a complete optimization to enable a maximized detection of the peptides composing the mixture. For each PTMs site identified through MS/MS spectra interpretation, the level of modification of the protein is estimated by comparing the intensity of the signal on the MS analysis between the modified peptide and the intact homologous. For that matter, it is important that both peptides are eluting at different times prior to their introduction to the MS instrument in order to prevent any biases in the estimation. Due to the robustness of this analytical approach, supported by high resolution MS analysis, the characterization of PTMs has been extensively used to map the occurrence of PTMs on proteins in vivo [124]. This methodology is also at the moment extensively used for the characterization of biopharmaceutical products to monitor PTMs described to have an adverse influence on the activity/functionality of proteins used for therapeutic purpose [125]. Therefore it is possible to envisage the study of the occurrence of PTMs on proteins which may be generated by their incorporation to a biomaterial. Indeed, the MS/MS spectra recorded from the analysis of the protein collected from the release of the biomaterial can be used to localize eventually the presence of PTMs on the peptide backbone. In parallel a similar analysis realized as a control experiment on a sample of the protein which was not incorporated to the biomaterial could further demonstrate the origin of the PTMs (Fig. 4.10). Considering the performance of the current instrumentation, this type of characterization is particularly simple to implement and should be developed in the near future, especially in the latter stage of innovative biomaterial development.

4.5

Conclusion

The research regarding innovative biomaterials especially targeting tissue regeneration involves a synergy of disciplines and expertise. The primary purpose of biomaterials is involving long-term interaction with highly complex biological matrices. In the meantime, major breakthroughs have been achieved in biology to develop advanced analytical methodologies able to provide extensive information regarding

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biological systems and their evolution over time. Proteomics is a prominent example of this category of analytical strategies as it provides the protein content of complex biological samples in unprecedented details with the possibility to monitor their respective expression level. Therefore it even allows to access information regarding biological signaling which cannot be obtained from genetic analysis delivering a complementary insight regarding the studied organism. The conception of innovative biomaterials can profit from the implementation of this type of advanced methodologies, once strictly reserved to fundamental biology investigation. In conjunction with the integration of other types of analytical strategies such as genomics or FC, the use of those tools can help to understand the complex biological implications of the integration of biomaterials for tissue regeneration. Thus because of their technical maturity, such techniques can be employed in various context. For instance, proteomic analysis can be implemented to scale subtle changes in the protein profile upon biomaterial exposure or implantation in order to interpret biological reactions. The analytical information provided can help to tailor biomaterials with improved compatibility with the biological environment. MS tends to play a key role in the comprehensive characterization of protein structure. This technique can also be used to deliver information regarding the structural integrity of protein that has been incorporated to biomaterials. That aspect appears particularly important to ensure maintained activity of proteins integrated to the structure of biomaterials which is a crucial requirement to retain the functionalization sought for the biomaterial. In this case as well, dedicated characterization can help to point endogenous structural instability originating from exposition to the biomaterial in order to design compatible media. With the constant improvement of MS-based analytical methodologies such as the introduction of glycomics and lipidomics, knowledge regarding biomaterial outcome are expected to be further improved in the future.

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507.

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Milad Fathi-Achachelouei1, Dilek Keskin1,2,3 and Aysen Tezcaner1,2,3 1 Department of Biomedical Engineering, Middle East Technical University, Ankara, Turkey, 2 Center of Excellence in Biomaterials and Tissue Engineering (BIOMATEN), Middle East Technical University, Ankara, Turkey, 3Department of Engineering Sciences, Middle East Technical University, Ankara, Turkey

5.1

Introduction

Limitations and drawbacks of conventional methods to restore damaged, diseased, or malfunctioned tissues or organs such as tissue transplantation have led to emergence of tissue engineering (TE) approach for regeneration, which seeks to devise alternative approaches to restore fully functionalized tissues/organs. TE approaches promise the regeneration of desired tissues but still require precise and detailed inspections for mimicking native tissue/organ constructs. Consequently, elements of TE (cells, scaffolds, and bioactive agents) should be chosen carefully for the desired tissue/organ. Heterogeneous nature of tissues/organs creates more complexity in the design criteria for TE applications and requires utilizing vast range of technologies for mimicking the extracellular matrix (ECM). Cellular microenvironment plays a crucial role in dictating the cellular activities that mediate them through ECM components such as structural components and soluble factors such as growth factors (GFs). In the TE point of view, delivery of soluble agents plays an important part in dictating the fate of cells; therefore controlled delivery of desired agents is necessary for the success of desired TE applications. Spatiotemporal and controlled delivery of bioactive agents plays critical role in cellular responses, as the signaling molecules should be targeted on specific site, at specific time and desired dosage to have its maximum efficiency over desired cellular activities [1]. Various platforms have been developed to fulfill the desired control over the delivery of bioactive agents. Chemical conjugation (either immobilization or covalent incorporation) or encapsulation of bioactive agents through suitable carrier could provide versatile toolbox for controlled delivery purposes [2]. Nanotechnology that has gained attention since the 1980s in many fields, such as electronics, space, and mechanics, also has recruited its elements for biomedical purposes especially for controlled delivery purposes. In this chapter, various types of nanoscale-delivery systems with focus on TE applications will be covered. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00026-X © 2020 Elsevier Ltd. All rights reserved.

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Properties and application areas of nanoscaledelivery systems in biomedical field

Nanotechnology field deals with the utilization of materials in the nanometer scale ranging from B10 to 1000 nm [3]. The main urge to shift to nanoscale systems arises from the altered material characteristic that cannot be observed in the macroor even microscale materials. Various properties of materials, including magnetic, electrical, optical, mechanical, and chemical could be altered. These nanoscale features have created approaches in particle and fiber production and surface patterning, which could open new horizons in biomedical field that was impossible in macro and microscale systems. Taking the advantages of altered material properties can provide a versatile platform to design nanoscale-delivery systems that can exert their effect through providing mechanical enhancement to scaffolds or delivery of highly stabilized agent with favorable solubility, enhanced biocompatibility, and targeted delivery features.

5.2.1 Delivery systems for therapeutic purpose Delivery systems for therapeutic purposes have been focused on eliminating tumors or treating other diseases. Precise delivery of the therapeutic agents to the disease site plays crucial role in the healing process. Direct administration of therapeutic agents as a conventional method of treatment has a couple of drawbacks (poor solubility, constrained efficiency, rapid clearance and degradation, poor distribution with lack of selectivity, dose-dependent, and site-dependent side effects and costintensive) which limits the success of treatments [4]. To deal with such drawbacks, various delivery systems have been developed. Hydrogels with different crosslinking mechanism and stimuli-responsive characteristic [5], electrospun fibers [6], sponges [7], films [8], implantable microchips [9], microneedles [10], and micro/ nanoparticle (NP) delivery systems [11,12] are among the various delivery systems which have been utilized to compensate the drawbacks of direct administrations. In the field of nanoscale-delivery systems, various nanostructures such as nanovesicles, nanotubes (NTs), rings, NPs, nanospheres, micelles, dendrimers, and NTs prepared with natural, synthetic, or composite materials have been widely used for the therapeutic applications [11]. Diversity of applications and methods for the delivery of therapeutic agents has provided the platform for seeking the suitable approaches for TE purposes and optimization of these systems independent or in parallel to the therapeutic delivery systems. As the delivery systems for therapeutic agents are beyond the scope of this chapter, cited review papers can be read for comprehensive information.

5.2.2 Delivery systems for tissue and organ regeneration Similar to the delivery systems purposed for therapeutic agent delivery, various bioactive agents have been incorporated into delivery platforms not only to cure

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diseases but also to regenerate damaged or diseased tissues and organs. Bioactive agents with regenerative capacity are composed of various classes of compounds that exert their functions through different mechanisms. These include GFs, cytokines, drugs, inhibitors, and genes, which can be introduced with different carriers and various administration routes [13
15]. Combination of these compounds with delivery systems has provided versatile toolbox to design tissue and organ specific delivery systems for TE purposes. In the nanoscale-delivery systems, various platforms, including fibers, hydrogels, patterned surfaces, particles, and composite materials, have been used for regeneration purposes [16
18].

5.3

Nanoscale-delivery systems for regeneration purposes

Nanoscale can enhance the delivery of therapeutics agents through spatiotemporal control over the action of agents at the diseased site. Such an approach will minimize the cytotoxic effect of therapeutics agents while keeping the required dosage within the milieu of targeted tissue for effective treatment. Thanks to their high surface
area-to volume ratio, nanoscale-delivery systems compared to other systems, such as microparticles, can provide higher adsorption of bioactive molecules and, consequently, resulting in higher interactions with targeted moieties [12]. Size of the particles plays crucial role in immune response, crossing barriers, endocytosis, and circulation [12]. In general, large particles in microscale cannot pass most of the biological barriers such as blood
brain barrier or cannot circulate within the bloodstream and could cause embolism due to aggregation [19]. Large particles also could provide immune response through their engulfment by foreign body cells [12]. Nanoscale particles can deliver their cargos via endocytosis by the cells and using targeted delivery systems can minimize cytotoxic or any undesired effect on other cells. Designing delivery systems either through conjugation or encapsulation of therapeutics agents in nanoscale-delivery platforms has been replaced by more complex systems. Nowadays, delivery systems are transforming from conventional delivery approaches with an aim of only cargo delivery to the multimodal platforms, which can not only deliver their cargo in respond to specific stimuli but also can provide the monitoring features to track the healing processes through interaction between material and cell, tissue, and organ. Such multimodal platforms have emerged as theranostics field that combines diagnosis and therapeutics approaches together for personalized medicine [20]. Various nanoscale-delivery platforms try to use novel biomaterials or intelligent designs to facilitate the spatiotemporal control of delivery systems within the embedded scaffolds, modulate the release of multiple compounds simultaneously or in sequential manner, take the advantage of tissue or cell microenvironment stimuli factors to trigger the reactions, modulate the mechanical properties of the carriers, minimize the cytotoxicity effect, and track the fate of cells in vivo and in vitro [1,3].

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Table 5.1 Morphological classification of nanoparticles in tissue engineering. Classification

Type

Example of applications

References

Spherical

Nanospheres Nanocapsules Polymeric micelles Polymersomes Nanogels Liposomes

Bone TE Bone TE Bone TE Bone TE Wound healing Dental and neural TE and angiogenesis Wound healing

[21,22] [23] [24] [25] [26,27] [28
31]

Bone TE, peptide delivery Cardiac, muscle, bone, cartilage, and neural TE Neural TE Cartilage and bone TE

[34,35] [36
41]

Nonspherical

Nanostructured lipid carriers Dendrimers Nanotubes Nanorods Nanocages

[32,33]

[42] [43
45]

TE, Tissue engineering.

5.3.1 Morphological classification of nanoscale-delivery systems Various classifications of the delivery systems are used, including morphology (Table 5.1) [46,47], material properties [48], manufacturing processes [49], administration routes [47], and tissue-specific applications [48]. Material composition with different shape and size can determine the fate of the delivery system and its success rate. Therefore in designing the delivery systems, limitations of each system should be considered through analyzing its elements. For example, for the delivery systems which will be administrated into blood, size of the particle is a determining factor that is critical for its circulation time in blood and penetration to the cells [50]. Shape or morphology of NPs plays a critical role in the radial drift toward the blood vessel walls. The circulation of oblate-shaped NPs are subjected to torques resulting in tumbling and rotation, which increase the lateral drift of NPs toward the blood vessel walls, which effects their uptake by cells [50]. Shape of NPs also determines their interaction with cell membranes which is very important for endocytosis by cells. NPs encounter immune cells and it is shown that macrophages have shown morphology-dependent attachment and internalization, which plays a crucial role in NP clearance from the body [51]. In this chapter, various forms of nanoscale-delivery systems with emphasis on the shape and composition will be covered to provide a comprehensive knowledge of impact of such systems over tissue and organ regeneration.

5.3.1.1 Nanoparticles NPs are colloidal, solid, and spherical particles to which bioactive agents can be conjugated, adsorbed, or encapsulated within. They have been widely utilized in TE context, thanks to various materials and manufacturing processes, which provide

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the diverse form of NPs. By choosing suitable manufacturing parameters and suitable compounds, it is possible to obtain nanospheres, nanocapsules, polymeric micelles, dendrimers, polymersomes, nanogels from polymeric NPs (PNPs) [19,47], liposomes and micelles from lipids [47], and inorganic NPs from various metals and ceramics [52]. Formation of composite NPs through the conjugation of polymeric materials with metallic materials increases the diversity of NPs even more. These compositions facilitate the conjugation of various bioactive agents onto NPs or their recognition by immune system [53].

5.3.1.1.1 Polymeric nanoparticles PNPs are among most widely used materials for designing delivery systems due to abundance of compounds which can be used, including natural compounds such as proteins and carbohydrates and synthetic compounds. Polymeric materials have tailorable physical, chemical, and mechanical features, which can be modulated through adjusting their shape, size, monomer elements, surface chemistry and charge, solubility, mechanical strength, biodegradation, etc. They generally have high biocompatibility, low cytotoxicity with the ability to carry both hydrophilic and hydrophobic bioactive molecules such as GFs, and protect their biological activities. Release from PNPs can be modulated through physical and chemical processes that can respond to environmental changes that enable to have a control over delivery process. PNPs with sensitivity to different physicochemical stimuli such as heat, pH, light, and temperature can be designed to release their cargo [54]. Most widely used production methods of PNPs are emulsification (single and double emulsion/solvent evaporation, emulsion polymerization, emulsion/solvent diffusion, and spontaneous emulsification), nanoprecipitation, supercritical fluid technology, self-assembly, dialysis, salting-out, and microfluidic systems [4,55,56]. Nanoparticles of natural polymers Natural polymers are biopolymers that are obtained from natural resources such as proteins (gelatin, albumin, silk, elastin, and zein) and polysaccharides [pullulan, alginate, chitosan, starch, heparin, dextran, and hyaluronic acid (HA)] [57
59]. These materials are mostly biocompatible and biodegradable and find wide applications for nanoscale delivery of bioactive agents for tissue and organ regeneration purposes and can modulate properties of tissueengineered scaffolds such as mechanical properties for optimum cellular activities. Due to their abundance in nature, their products are cost-effective; however, their batch to batch variations, some material-dependent immunogenic responses, complex extraction methods and complex structures can limit their utilization [57
59]. Carbohydrate-based nanoparticles Carbohydrates are one of the diverse and abundant natural compounds, which alone or in combination with other materials such as proteins and synthetic polymers have been used in the preparation of delivery systems. Through modification of these compounds, it is possible to modulate the desired delivery system for tissue and organ regeneration. Chitosan as one of the most abundant polysaccharides is obtained from partial deacetylation of chitin. Due to its low immunogenicity, high biocompatibility, biodegradability, and antimicrobial activity, it has been widely used for TE

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applications. Due to the presence of positive charge in chitosan, it can undergo selfassembly with negatively charged molecules to form NPs. Heparin which is natural glycosaminoglycan (GAG) with its negatively charged characteristics and GF binding domain has been widely used to form NPs with chitosan which would facilitate the cytokine delivery for TE applications [60]. Stromal cell
derived factor-1α (SDF-1α) and vascular endothelial GF (VEGF) which were loaded separately into chitosan/heparin NPs have been shown to enhance the mesenchymal stem cells (MSCs) migration and proliferation compared to control groups, thanks to NPbased delivery system that exerts its effect on prolonging their release and protecting their bioactivity [60]. Similar composition of chitosan/heparin NPs have been utilized for delivery of GFs for other tissues and organs. Chitosan/heparin NPs were immobilized to decellularized bovine jugular vein scaffolds [61]. It has been reported that VEGF has been localized more in NPs through physical adsorption and modification of scaffolds with 1-(3 dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride and N-hydroxysulfosuccinimide. Burst release of VEGF was lower compared to control groups, and VEGF release was concentration-dependent. In vitro and in vivo studies have shown higher endothelial cell proliferation and massive new capillary formation, respectively. Recently, two works have been conducted using chitosan/heparin NPs for bone TE. In one of the reported works, SDF1 and bone morphogenetic protein 2 (BMP-2) were loaded in NPs which were immobilized on chitosan
agarose
gelatin scaffolds [62]. SDF-1 and BMP-2 release has provided migration and osteogenic differentiation of MSCs, respectively. In another study, dual delivery of placental GF-2 (PlGF-2) and BMP-2 in chitosan/heparin NPs resulted in better bone tissue regeneration through providing higher alkaline phosphatase (ALP) activity of cells and mineral deposition by cells compared to the delivery of either GFs alone [63]. In another study, PlGF as an angiogenesis factor has been encapsulated in chitosan/alginate NPs to provide an intramyocardial injection platform for acute myocardial infarction treatment [64]. Eight weeks after coronary ligation, enhancement in vascular density, leftventricular function, and antiinflammatory cytokine interleukin-10 (IL-10) serum level and decrease in scar area formation and level of proinflammatory cytokines tumor necrosis factor-alpha (TNF-α) and IL-6 were observed in sustained release chitosan/alginate NPs group. Potential usage of chitosan NPs for the encapsulation of nerve GF (NGF) has been reported. In this study a sustained release of NGF for transdifferentiation of MSCs to neurons was aimed [65]. MSCs that were treated with NGF-loaded NPs have shown higher expression level for tubulin βIII and microtubule-associated protein 2 (MAP-2) as neuronal cell markers compared to control groups. Polysaccharide-based NPs not only can deliver cytokines for the regeneration of tissue and organs but also can used to protect them through releasing antiinflammatory agents to modulate the immune responses. Butyrate as short-chain fatty acid as antiinflammatory agent has been encapsulated in chitosan/ hyaluronan NPs and this delivery system was designed to provide a sustained release of the agent to limit reactive oxygen species (ROS)
dependent tissue injury during inflammation [66]. Butyrate-loaded NPs were reported to inhibit H2O2 release by TNF-α exposed neutrophils without a significant decrease in the inhibitory effect

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compared to free butyrate. In another study, diclofenac as antiinflammatory drug was encapsulated in chitosan/poly(γ-glutamic acid) (γ-PGA) NPs for intradiscal injection to reduce proinflammatory mediators. They observed upregulation of collagen type II and aggrecan. This approach can be used for preventing intervertebral disk (IVD) degeneration and mediating its regeneration (Fig. 5.1) [67]. Polysaccharide-based NPs not only were used for direct delivery of cytokines but were also used for gene therapy applications. An antiinflammatory IL-10 encoding plasmid DNA has been encapsulated in alginate NPs which was decorated with tuftsin, a four amino acid peptide on its surface to interact with macrophages. It was reported that macrophages with NPs internalization acted such as Trojan horse vectors for these NPs and carried them to the site of inflammation (arthritic joints) leading to alleviation in the levels of proinflammatory cytokines and preventing joint damage [68]. Nanogels which are also named as hydrogel NPs are nonfluid colloidal/polymer networks and gained significant attention due to their adjustable hydrophilic and stimuli-responsive characteristics providing exclusive physicochemical advantages for encapsulation and release of bioactive agents [69]. Various polysaccharides have been utilized for the production of nanogels [69]. In a study by Aslan et al.,

Figure 5.1 Antiinflammatory chitosan/poly(γ-glutamic acid) NPs control inflammation while remodeling extracellular matrix in degenerated intervertebral disk. NP, Nanoparticle. Source: Reproduced from Teixeira GQ, Leite Pereira C, Castro F, Ferreira JR, GomezLazaro M, Aguiar P, et al. Anti-inflammatory chitosan/poly-γ-glutamic acid nanoparticles control inflammation while remodeling extracellular matrix in degenerated intervertebral disc. Acta Biomater 2016;42:168
79 with permission from Elsevier.

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chitosan-penta-basic sodium triphosphate nanogels have been used as delivery system for IL-2 for accelerating wound healing in rats. In vitro studies have shown burst release of IL-2 in first 2 hours [26]. For in vivo experiments, nanogels were embedded within Natrosol gels and applied in the back of rats. Application of IL-2 bearing nanogels decreased the level of malondialdehyde as lipid peroxidation marker and increased the level of antioxidant glutathione compared to control groups. For the wound-healing applications, other nanogel systems such as cholesterol bearing pullulan (CHP) nanogel cross-linked in the silicone membrane has shown superior results compared to collagen membranes [27]. HA-based nanogels are good protein carriers. Sustained release of recombinant human growth hormone (rhGH) from nanogels could be an alternative administration route for daily injection that is required for several years to treat short stature in children. Cholesterol group bearing HA nanogels with high degree of substitution that were prepared by salt-induced association of tetrabutylammonium salts of HA have been shown to bound the proteins without causing denaturation. Sustained release of rhGH has been observed over 1 week in rats [70]. Nanogels not only can be utilized in TE applications but also can be used as nanoprobes. In vivo tracking of adipose-derived stem cells (ADSC) through internalization of fluorescent molecule conjugated dextran nanogels with low cytotoxicity has been reported which can provide versatile tool for tracking the fate of stem cells in the tissue and organ regeneration [71]. Protein-based nanoparticles Protein-based PNPs have been widely used for the delivery of bioactive agents either for the treatment of diseases or TE applications. Growing interest for using polymeric-based materials arise from their biocompatibility, nonantigenicity, abundant source, scale-up capacity for manufacturing, and high binding capacity to the bioactive agents [72]. Less opsonization by the reticuloendothelial system and versatile functional groups provides opportunities for the utilization of protein-based materials to both eliminate their fast clearance from the body and modify their structure for nanoscale-delivery systems of bioactive agents for TE applications. Gelatin is a natural protein that can be derived from collagen through hydrolysis. It has high biocompatibility and biodegradability under physiological conditions. It is a cheap and abundant protein that bears Arg
Gly
Asp (RGD) cell attaching sequence, and it has been widely used in TE applications [73]. Dexamethasoneloaded, pH sensitive gelatin micelles grafted with lactic acid oligomers were used for osteogenic differentiation of bone marrow MSCs (BM-MSCs) through internalization of micelles [24]. BM-MSCs were precultured with micelles for 24 hours and seeded on gelatin hydrogels which resulted in superior in vivo bone formation due to the intracellular release of dexamethasone. This enhancement of osteogenic differentiation of BM-MSCs was reported to be dose-dependent. Usage of gelatin and chitosan/gelatin NPs have also been reported for TE applications [21,74]. For better targeting of gelatin NPs, it is possible to conjugate bone-targeting alendronate to NPs for increasing the affinity of NPs to mineralized tissue [75]. Gelatin-based NPs have been used for neuroprotection in the postischemic conditions [76], peripheral nervous system repairing [77], and neural TE model [78]. In the peripheral nerve repairing model, multichannel scaffolds were prepared by utilizing aligned

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electrospun nanofibers to mimic the nerve fascicular structure with neurotrophic gradient to enhance axonal regrowth [77]. Brain-derived neurotrophic factor (BDNF) was encapsulated in gelatin NPs, while NGF was added into the scaffold freely. Gradient maker was used to establish the gradient for both BDNF-loaded NPs and free NGF. Fast delivery of NGF and delayed delivery of BDNF promoted the initial stage of axon regeneration and late stage of myelination process, respectively. Differentiated neural stem cells effectively extended their neurites along the aligned nanofibers, and higher cell density was observed in regions with higher NGF concentration. Elastin protein as another ECM component has shown promising results in the form of thermo-responsive self-assembled NPs, and they were reported to be good protein encapsulation platform for TE applications [79]. More than 100,000 proteins exist in human plasma which have nontoxic and biodegradable characteristic with long in vivo half-lives. Just a few of these proteins such as albumin and fibrin [80] have been used as nanocarrier platform for drug delivery, imaging, and tissue-regeneration purposes [81]. Albumin as the most abundant plasma protein has been used as nanocarrier formulation alone or in combination with other polymers for bone formation [82] and IVD regeneration [83]. SDF-1α encapsulated albumin/heparin NPs have been designed as an injectable carrier for migration of BM-MSCs to the defect site for promoting IVD regeneration [83]. Stem cell migration has shown to be dose dependent in vitro (41% and 64% BMMSCs migration after 24 hour at 50 and 100 ng/mL SDF-1α concentrations, respectively). Similar to control groups, cellular proliferation remained high. Accelerated IDV regeneration was also reported at gene expression and protein levels for SOX9, collagen type II, and aggrecan, and with histological analysis compared to free SDF-1α (control group). Silk fibroin has gained significant attention in TE applications due to its biocompatibility, excellent and adjustable mechanical strength, but it has been rarely utilized for NPs production. However, it has been widely used in other forms such as hydrogel, porous scaffold, and fiber. Use of silk fibroin NPs for bone TE [84
86] and angiogenesis [87] was reported. Protein-based NPs are not limited to only animal origin proteins, and plant-based proteins can also be utilized for TE applications. Zein that can be obtained from corn have hydrophobic characteristic. Diclofenac-loaded zein NPs produced by antisolvent precipitation method were embedded within PVA nanofibers through single-nozzle electrospinning method for preparing a wound dressing material [88]. Lower burst release of the antiinflammatory agent was observed from NPs embedded in fibers compared to free NPs. In vitro results have shown significantly higher spreading and proliferation of L929 fibroblast cells on hydrogels containing zeinloaded NPs compared to control groups. Nanoparticles of synthetic polymers Polymers of synthetic origin such as poly (anhydride)s, poly(ester)s, poly(urethane)s, poly(amidoamine)s (PAMAMs), and poly(acrylate)s have been used alone or in conjugation with other materials such as metals, ceramics, lipids, and natural polymers for nanoscale TE applications [89], imaging, and therapy [90]. They provide high purity, reproducibility, minimum

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batch to batch variation, and tailorable properties. Poly(ε-caprolactone) (PCL), poly (glycolic acid) (PGA), and poly(lactic acid) (PLA) and their copolymers such as most widely used poly(lactic-co-glycolic) acid (PLGA) are members of poly(ester) family of polymers which have been used widely in TE applications due to adjustable biodegradability and biocompatibility [91]. PLGA NPs could be manufactured through various procedures such as electrospraying [91] and emulsification-solvent evaporation [92]. TE applications of PLGA NPs alone or in combination with other materials include regeneration of periodontal tissue [93,94], bone [95
97], brain [98], neural protection and repair [99,100], angiogenesis [101], vessel repair [102], cardioprotection [103
105], lung [106], pancreas [107], and bladder [108]. Gene and RNA delivery have been conducted using PLGA NPs such as small interfering RNA (siRNA) delivery to inhibit the translation of abnormal gene expression in cells for bone tissue applications [109], gene delivery to prevent macular degeneration [110], and gene delivery for chondrogenesis [111]. To make release system sensitive to stimuli or to target specific sites PLGA could be modified either through incorporation of functional groups or various polymers for TE applications. PLGA-modified chitosan has been developed to introduce pH sensitive stimuli response feature for the release of diclofenac [112]. It is possible to manipulate the polymer-bioactive agent interaction through either coating or copolymerizing the PLGA NPs with more hydrophilic polymers such poly(ethylene glycol) (PEG) which could increase the hydrophilic agent adsorption or encapsulation, prolong NPs residence time in blood, and improve tissue-specific targeting [113,114]. PCL as another famous poly(ester) family member has been widely used in TE of different tissues. It is biocompatible, cheap, but more hydrophobic and less degradable compared to PLGA [115]. It is generally used in manufacturing of scaffolds for TE applications rather than as nanoscale-delivery systems. PCL NPs have been utilized for the delivery of VEGF [116]. PCL NPs incorporated poly(L-lysine)/ HA polyelectrolyte multilayer film system has provided long-term bioactive retention capacity of NPs. In another study, VEGF encapsulated NPs were incorporated into an injectable gelatin and silicate-based shear-thinning hydrogels for irregularshaped defects in bone [117]. Cocultured osteogenic and endothelial cells that were encapsulated in the hydrogel showed enhanced growth and differentiation in the presence of VEGF-loaded PCL NPs. Polymersomes that are vesicles formed by the self-assembly of amphiphilic block copolymers have a hydrophilic core and a hydrophilic shell which can accommodate a variety of hydrophilic and hydrophobic molecules. 6-bromoindirubin-30 -oxime
loaded PEG-PCL block polymersomes were used as delivery system for activating Wnt pathway for bone regeneration [25]. Wnt-signaling activation in sustained manner resulted in the promotion of early osteogenic differentiation in BM-MSCs. Poly(N-isopropylacrylamide) (PNIPAM) is an acrylic polymer that can be used for the production of stimuli-responsive NPs. pH and temperature sensitive PNIPAM hydrogel NPs were used as delivery system for bio/chemosensors, biological imaging, and drug-delivery applications [118]. PNIPAM NPs have been used for delivery of antiinflammatory peptides to prevent cartilage degeneration in osteoarthritis [119,120]. Retinoic acid
loaded PNIPAM-co-acrylamide NPs have been

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used for neuronal differentiation of human-induced pluripotent stem cell (hiPSC)
derived neuronal precursor cells [121]. VEGF-loaded thermosensitive PNIPAM NPs in the collagen gel matrix have been used for bone TE applications [122]. NPs loaded in hydrogel were added to primed endothelial-osteoblast directed BMMSCs. Better angiogenic/osteogenic differentiation was observed in NPs containing hydrogels compared to free NPs group as the gelatin matrix decreased the burst release of VEGF from PNIPAM NPs. Positively charged N-(3-aminopropyl) methacrylamide, neutral monomer acrylamide, and glycerol dimethacrylate as degradable cross-linker have been used for the encapsulation of BMP-2 through in situ polymerization [123]. Enrichment of the monomers and cross-linkers around GFs occurs due to hydrogen bonding and electrostatic interactions which could provide tunable sustained release of BMP-2 for bone TE. Applications of poly(anhydride)-based NPs have been reported in drug delivery such as mucus surface penetration that can be found in eyes, gastrointestinal tract, airways, and nasopharynx surface. Surface and epithelium cells of mucus can be targeted for both drug and gene delivery in TE applications. Thiamine-coated poly (anhydride) NPs [124] and biodegradable NPs composed of poly(sebacic acid) and PEG [125] have been reported for these purposes. Poly(anhydride) can also form block copolymers with other polymers for preparing NPs. Poly(styrene-alt-maleic anhydride)-b-poly(styrene)
based NPs have been utilized for targeted delivery of β-catenin agonists for bone fracture healing [126]. Poly(urethane)s have been utilized in TE applications due to their outstanding mechanical properties and acceptable biocompatibility. Poly(urethane) NPs have been used as coating material to increase the anticoagulation properties of the carriers and development of more hemocompatible vascular patches [127,128]. Poly (urethane)-based nanomicelles can also be used as siRNA/microRNA delivery systems. Osteoblast targeting peptide, serine-aspartic acid-serine-serine-aspartic acid (SDSSD)
modified poly(urethane) nanomicelles were used to deliver anti-miR-214 microRNA to osteoblasts which resulted in increased bone formation, improved bone microarchitecture, and increased bone mass in an ovariectomized osteoporosis mouse model [129]. Dendrimers are spheroidal hyperbranched molecules with branches outgrowth symmetric around the core and can form spherical morphology depending on the molecular weight. PAMAM is the most widely used polymer for dendrimer production in TE applications due to low cytotoxicity, controllable size, and ease of modification. PAMAM has been used incorporated in gelatin scaffold to enhance angiogenesis for skin TE [130]. PAMAM was reported to enhance tensile strength of scaffold by acting as cross-linking agent. It also increased cellular adhesion, proliferation, and gene expression for hypoxia inducible factor (HIF1α), VEGF, and collagen type I [130]. In another study, enhancement in adhesion, proliferation, and functionality of HepG2 cells using paper matrix (cellulosic filter paper) functionalized PAMAM were reported as an in vitro 3D liver model [131]. Dendrimer applications also include the delivery of peptides and genes. Poly(ε-lysine)-based dendrimers have been used for the delivery of VEGF blocker aptamer sequence (WHLPFKC) as antiangiogenic factor for tissues such as cartilage and cornea [34].

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Complex of BMP-2 gene with PAMAM was reported to be used for functionalizing titanium surface which resulted in enhancement in osteogenic protein expression in subcutaneous introduction and also in bone tissue (femur) in rats [35]. Electrically conductive polymers can provide a cell differentiation platform for dictating the fate of stem cells as an alternative approach for GF delivery systems. 3D PLA nanofibrous scaffolds bearing conductive poly(aniline) NPs have shown better cell proliferation, osteogenic differentiation, and calcium mineralization compared to pure PLA scaffolds [132].

5.3.1.1.2 Lipid nanoparticles Nanostructured lipid carriers (NLCs) are another important material family for the delivery of diverse hydrophobic bioactive agents. Encapsulating lipophilic compounds require colloidal dispersions suitable in aqueous environments such as oil in water emulsions. Lipid-based nanoscale-delivery systems have better biocompatibility, lower cytotoxicity, and biodegradability [133]. Digestible lipids can provide sustained and targeted release for controlled delivery systems [133]. Lipid-based carriers can take various self-assembled structures such as cubosomes, hexosomes, tubules, ribbons, cochleates, and lipoplexes, which have been reviewed in literature [134]. Lipid-based nanocarriers can be in the form of nanoemulsions, liposomes, solid lipid nanoparticles (SLNs), and NLCs [133,135]. Lipid-based nanoemulsions have been used as therapeutic delivery platform such as for the prevention of neuroinflammation. Miglyol 812 is a medium-chain triglyceride, which was utilized to form lipid-based nanoemulsions. Celecoxib and near infrared (NIR) dye incorporated nanoemulsions were used to track and cure neuroinflammation of the sciatic nerve in rats [136]. Nanoemulsions are incorporated into monocytes after intravenous injection and are accumulated in injury side. Significant reduction in the visualized inflammation, infiltration of macrophages, Prostaglandin E2, and cyclooxygenase-2 secretion were observed. As a new generation of nanoparticulate active-substance vehicle, SLNs consist of pure solid lipids, while NLCs are made of a solid matrix entrapping variable liquid lipid nanocompartments [137]. NLC-based delivery systems have been used for wound-healing applications. Incorporation of Aloe vera containing NLCs was reported to increase the ultimate tensile strength of electrospun nanofibrous PLGA dressings and, consequently, better handling of them [32]. In another study, recombinant human epidermal GF (rhEGF)
loaded NLCs were prepared and topically administrated to porcine full-thickness excisional wound model. rhEGF-NLCs administration was reported to improve the wound-healing quality expressed in terms of number of arranged microvasculature, fibroblast migration and proliferation, collagen deposition, and evolution of the inflammatory response compared to control groups [33]. The application of SLNs is generally concentrated on drug/ gene targeting to tumors and other diseased organs such as ocular tissue [138]. Comparison between andrographolide-loaded SLN and NLC delivery systems has been reported for wound healing applications [139]. NLCs had higher encapsulation and release profile for andrographolide. NGF-loaded heparinized cationic SLNs conjugated with cationic lipid esterquat 1 or stearylamine were reported for

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regulating membrane charge of induced pluripotent stem cells (iPSCs) during differentiation. Zeta potential and fluorescence staining studies showed that NGFloaded esterquat 1 conjugated SLNs induced the differentiation of more iPSCs toward neuron-like cells than stearylamine [140]. Liposomes are amphiphilic lipid vesicles, similar to cell membrane, which selfassembles in aqueous environment and have been used widely in TE applications for the delivery of various bioactive agents. Using different lipid chromophores, purpurin
phospholipid and pyropheophorbide
phospholipid, suspension with multiple payloads could selectively release their bioactive agents upon photoactivation at specific wavelength irradiation [141]. Liposomes can be coated by various materials such as chitosan to improve their biocompatibility for TE applications [142]. Incorporation of liposomes in the porous chitosan thermosensitive gels using β-glycerophosphate cross-linker which solidifies above the temperature of 37 C has been recently reported which could provide nanoscale-delivery system in the form of injectable hydrogels for TE applications [143]. Liposomes have been used for dental tissue regeneration through the encapsulation of demineralized dentin matrix that contains various GFs such as VEGF and transforming GF-β1 (TGF-β1) [28]. Liposomes promoted chemotactic recruitment of progenitor cells in a dosedependent manner. Upregulation of osteodentin markers, osteocalcin and runtrelated transcription factor 2, and enhancement of biomineralization were observed in the dental pulp stem cells (DPSCs). Liposomes were used for neural differentiation such as delivery of serotonin receptor agonist (RS67506) in pluronic F-127, a thermosensitive hydrogel, to differentiate postnatal gut-derived enteric neural stem/ progenitor cells [29], and delivery of paclitaxel for induction of neuronal differentiation of neural stem cells through Wnt/β-catenin signaling for spinal cord injury repair [30]. Also, liposomes loaded with glypican-1 were used to improve the effectiveness and responsiveness of GF to therapies in the context of ischemia in the diabetic disease state through promoting angiogenesis [31].

5.3.1.1.3 Inorganic nanoparticles Inorganic materials such as metals and ceramics and their composites prepared with other materials have been widely utilized for TE applications. Main applications of inorganic materials are focused on preparing the implants for hard tissue applications, but due to their unique features, they also have been applied for the preparation of nanoscale-delivery systems. Ceramic nanoparticles Ceramics are solid inorganic materials with brittleness, high hardness, thermal and electrical insulation, and corrosion-resistance characteristics. They can be in the crystalline or amorphous form [144]. There are various subclasses of ceramic materials that have been applied in nanoscale for TE applications with material types such as hydroxyapatites (HAp), silicates, calcium phosphates (CaPs), and metal oxides [145]. Bioceramics can be categorized as resorbable and nonresorbable, or as bioactive and bioinert based on their interaction with the host tissue [144]. Recent advances in application of various nanoscale bioceramics in tissue and organ regeneration will be covered.

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Oxides of various metals such as zinc, titanium, aluminum, cerium, and iron have been used widely in TE applications. Titanium dioxide (TiO2) NPs as were used for enhancing the elastic modulus of cellulose-graft-poly(acrylamide)/HAp nanocomposite scaffolds to levels similar to mechanical properties of trabecular bone [146]. TiO2 NPs were also reported to have antibacterial characteristics, and they were applied for preventing biofilm formation after implantation [147] and wound dressing [148]. One concern regarding the TiO2 NPs utilization is the accumulation of these NPs in other organs such as liver that can create toxic effects at organ and systemic levels. In a study regarding this issue, it was reported that TiO2 NPs induced a strong oxidative stress in primary hepatocytes and morphological changes in mitochondria, which lead to damage to primary hepatocytes [149]. TiO2 NPs are reported to induce an abnormal state of macrophages characterized by suppressed innate immune function and high inflammation which make their utilization controversial for TE applications [150]. Cerium oxides (CeO2/Ce2O4) are another type of metal oxides, and their NP form has shown positive effects over angiogenesis [151] and bone formation [152]. It has been reported that enhancement in the proliferation of BM-MSCs was proportional to the increase in the ratio of Ce41 to Ce31 which were coated over titanium implants. Such increase in ratio also enhanced osseointegration and osteogenic differentiation in vivo, and enhanced M2 polarization of macrophages in vitro [152]. Cerium oxide NPs also showed antibacterial properties [153], antioxidant activity in brain [154], and they are also used for wound-healing applications [155]. In vitro results showed that these NPs caused a decrease in calcium dysregulation and neural cell death, and when they were administered in rats, enhanced cognitive function and reduction in macromolecular free radical destruction were observed [154]. Zinc plays a crucial role in growth, development, and well-being for mammals [156]. Zinc oxide NPs were utilized for various purposes in TE applications. They have shown antibacterial and osteoconductive properties for guided tissue regeneration for periodontal TE [157]. Positive effect of zinc oxide NPs over cell adhesion and migration has made it a suitable candidate for blood vessel formation [158] and antibacterial wound healing [159
161] applications. Magnesium oxide NPs, which have excellent antimicrobial properties [162], have been used for TE. Enhancement in mechanical properties (tensile strength and elastic modulus) of nanocomposite scaffolds, and adhesion, proliferation, and differentiation of osteoblast-like MG-63 cells on magnesium oxide incorporated scaffolds were reported [163]. Low concentration of magnesium oxide NPs has been reported to induce antiapoptotic, antioxidative, and antidiabetic effects on rat pancreatic cells in vitro [164]. Iron oxides (Fe3O4 or Fe2O3) are well known and exclusively studied NPs in the biomedical field. Thanks to their magnetic properties, they have been used in versatile applications, including cancer therapy, diseases treatment, gene delivery, antimicrobial agent, TE, and imaging. Iron oxides can be modified using ligandmediated approach for stimuli-responsive applications [165]. These NPs were used for bone [166,167], cardiac [168], vascular [169], and cartilage [170], tissue/organ regeneration. It was reported that iron oxide NPs enhance osteogenic differentiation

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of DPSCs by increase in ALP activity, osteogenic gene expression, and bone matrix mineral synthesis [171]. Using iron oxide NPs within scaffolds not only can enhance the bone formation but also can provide platform for tracking the degradation of scaffold using noninvasive magnetic resonance imaging (MRI) [172]. Utilizing pulsed magnetic field on iron oxide labeled MSCs in rat model with cartilage defects has been reported to provide both in vivo monitoring of MSCs with MRI and upregulation of certain cartilage biomarkers; therefore such a system could be applied for chondrogenic differentiation [170]. Bioactive glass ceramic NPs (nBGs) were first introduced by Larry Hench and colleagues with a composition of Na2O
CaO
SiO2 [144]. Other elements such as magnesium, potassium, phosphorus, calcium, sodium, and boron can be incorporated in nBGs to modify them for various TE applications. Techniques such as laser spinning, microemulsion, gas phase synthesis, sol
gel technique, and precipitation have been used for synthesis of nBGs [173]. Application of nBGs is not limited to bone TE applications. nBGs were reported to stimulate rapid hemostasis and fibroblasts proliferation as well as angiogenesis enhancement [174,175]. Incorporation of nBGs (12 nm mean diameter) instead of conventional BGs (200 nm mean diameter) into gelatin hydrogels has been reported where a nanocomposite hydrogel with thixotropy (becoming less viscous) characteristics for wound healing applications was formed [174]. Such a composite is normally in the gel form but becomes injectable under shear stress. In vivo results have shown regeneration of cutaneous tissue within 7 days in rats. nBGs have been utilized with various materials and different morphology for bone TE to enhance mechanical properties of scaffolds or osteogenic differentiation of various cells (osteoblasts, bone marrow stromal cells, BM-MSCs, umbilical cord Wharton’s jelly-MSCs, and adipose derived-MSCs) [176]. nBGs incorporated in dextran hydrogel [177], poly(ethersulphone) nanofibres [178], coated over poly(L-lactic acid) [179,180] or polymer-coated bioactive glass [181], are among the vast number of the reported studied models for bone TE applications. Incorporation of other elements in the BGs can either enhance cellular proliferation and differentiation or add other characteristics such as antibacterial features to the scaffolds. Strontium (Sr21) and cobalt (Co21) ions substituted BGs can enhance the osteogenesis and angiogenesis, respectively, for bone TE [182,183]. Incorporation of copper (Cu21) was reported to enhance antibacterial characteristic of scaffolds [184]. Recently, boron-modified nBGs were reported to enhance odontogenic differentiation of DPSCs which were proven by immunocytochemical staining of dentin sialophosphoprotein, osteopontin, and collagen I [185]. Boron-doped nBGs regenerate the dentin when nBGs were incorporated within cellulose acetate/oxidized pullulan/gelatin-based dentin-like constructs through increase in cellular viability, ALP activity, and intracellular calcium deposition [186]. nBGs are not limited to act directly on the cellular activities, as they can be used as miRNA delivery carriers when they are synthesized with ultralarge pores with monodispersed characteristics [187]. Most widely used ceramic in TE is CaP
based bioresorbable family of materials that includes HAp, CaP, dicalcium phosphate dehydrate, tricalcium phosphate, calcium carbonate, calcium aluminate, and octacalcium phosphate [187]. HAp is the

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most widely used CaP ceramic due to its similar composition to the natural bone and suitable stability under acidic, neutral, and acidic conditions. HAp can be synthesized through precipitation, hydrolysis, hydrothermal methods, etc., or extracted from various natural sources [188]. Mono ion substitution of various metals into HAp lattice such as zinc, strontium, copper, magnesium, cobalt, and silver could change the scaffolds properties through modulating the degradation rate, mechanical properties, and cell
material interactions [189,190]. Enhancement of compression modulus and ultimate compressive strength were reported through incorporation HAp NPs in porous shape memory poly(urethane) [191] and PCL/ PLGA [192] scaffolds, respectively. However, enhancement in mechanical properties are not the only possible outcome of the incorporation of HAp NPs. For example, a decrease in the compressive modulus was observed when weight percentage of NPs incorporated in gelatin scaffolds was increased as NPs contribute in crosslinking efficiency of scaffolds negatively [193]. Similar to nBGs, HAp NPs have been incorporated within various carriers such as porous scaffolds [194], hydrogels [195], and electrospun fibers [196,197] to modulate the scaffold properties for bone and dental TE applications. Depending on the scaffold features and properties of HAp NPs, various cellular responses could be observed. For example, biomorphic PLGA/HAp NPs composite has enhanced proliferation, attachment, and differentiation of preosteoblastic cells [198]. Sometimes for appropriate enhancement of cellular responses, two or more ion substitution within the HAp lattice is required. Dual-doped HAp NPs with Ba21 and Ho31 ions were reported to act as contrast agents for computed tomography (CT) with improved CT contrast efficiency even under different CT operating voltages [199]. These kinds of systems can be used for clinical CT imaging or for tracking in vivo degradation of scaffolds in vivo (Fig. 5.2A and B). Dual doping of HAp with ferric (Fe31)/selenite (Se422) ions enhanced degradation and cytocompatibility of NPs which resulted in improved osteoblastic differentiation of stem cells (higher ALP activity and intracellular calcium) (Fig. 5.2C and D) [200]. CaP NPs were widely used by combining natural and synthetic polymers to form nanocomposite materials with superior mechanical properties mainly for bone TE [201]. Aerosol-derived flame spray pyrolysis method was reported to be used for synthesis of CaP NPs. Osteogenic differentiation of urine-derived stem cells treated with CaP NPs was reported [202]. CaP NPs were reported to be used successfully for guided bone regeneration in silk fibroin-PCL
PEG
PCL/PCL
based bilayer membranes [203]. CaP utilization is not limited to directly stimulate the bone-tissue regeneration. CaP can be used as nonviral vector for gene delivery [204] and for promoting wound healing [205] and angiogenesis [206]. BMP-2-encoding plasmid DNA-functionalized CaP NPs incorporated in nano-HAp
collagen scaffolds were implanted in rats [204]. Higher ALP activity was observed as gene transfer vectors in CaP NPs induced a larger yield of BMP-2 for a longer period than by scaffolds loaded with BMP-2 solution or naked plasmid. CaP NPs can be used for coating various surfaces. Laser melting produced titanium which was first modified by anodization and then followed by CaP NPs deposition on the surface of titanium [206]. Surface with anodization and CaP NPs deposition were shown to improve

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Figure 5.2 Efficiencies of CT contrast for undoped and doped HAp NPs with different concentrations determined at various voltages: (A) Hounsfield unit measurements (radiodensity measurement that is obtained through radiation attenuation in various tissues) and (B) CT images. (C) ALP activity and (D) intracellular calcium production of hFOB cells seeded on Fe-SeHAp disks after incubation in osteogenic differentiation medium at different incubation periods. ALP, Alkaline phosphatase; CT, computed tomography; HAp, hydroxyapatites; NP, nanoparticle. Source: (A and B) Reproduced from Zheng X, Wang S, Wu L, Hou X. Microwave-assisted facile synthesis of mono-dispersed Ba/Ho co-doped nanohydroxyapatite for potential application as binary CT imaging contrast agent. Microchemical J 2018;141:330
36 with permission from Elsevier. (C and D) Reprinted from Alshemary AZ, Engin Pazarceviren A, Tezcaner A, Evis Z. Fe(31) /SeO42(2) dual doped nano hydroxyapatite: a novel material for biomedical applications. J Biomed Mater Res, B: Appl Biomater 2018;106(1):340
52 with permission from John Wiley and Sons.

human umbilical vein endothelial cells adhesion, proliferation, and expression of angiogenesis-associated genes. The group suggested that these NPs could be used on implants to improve vascularization of hard and soft tissue interface. Metallic nanoparticles Metallic NPs have fascinated scientists for over a century and are now heavily utilized in biomedical sciences and engineering. Metallic NPs have been widely used in biomedical field for various purposes such as drug/gene delivery, imaging, and TE [207,208]. Synthesis and modification of metallic NPs with different chemical functional groups can provide the platform for the conjugation of various ligands for nanoscale-delivery systems for tissues and organs.

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Gold NPs (AuNPs) are most widely used metallic NPs thanks to their ease of manufacturing and different properties in nanoscale such as unique interaction with light compared to the bulk gold [209]. Various gold salt
reduction methods were developed for synthesizing AuNPs with a core diameter ranging from 1 to 150 nm, and various moieties such as phophines and thiols can be used to functionalize the surface of NPs [210]. Most of the applications of AuNPs were focused on the cancer diagnosis and therapy [211]; however, their applications in terms of tissue regeneration have recently increased. AuNPs are potential candidates as a safety measurement by targeting the remaining cancer cells after tumor removal through various approaches such as photothermal therapy [212]. This approach would provide a safer approach for the regeneration of desired tissue as replacement of tumor region. AuNPs have been reported to promote various cellular activities, including differentiation of stem cells [213]. AuNPs have been used in nerve [214,215], bone [216
222], cardiac [223
225], vascular TE [226] and also for wound healing purposes [227]. Dual functional AuNPs produced by conjugation of 2,2,6,6-tetramethylpiperidine-N-oxyl were reported to be internalized by MSCs and, consequently, induced a decrease in the overproduction of ROS, the suppression of adipogenic, and the enhancement of osteogenic differentiation of MSCs [228]. This system can provide both desired differentiation of stem cells and prevent dysregulated ROS secretion-dependent dysfunctions. Modifying AuNPs can enhance its action of targeting tissue. AuNPs modified with RGD have been reported to affect the differentiation of MSCs in a density-dependent manner [229]. Low density RGD decreased adipogenic gene expression and oil droplet formation, while high RGD density decreased ALP activity and enhanced adipogenic gene expression of MSCs. AuNPs-cells interaction can be influenced by various factors, and different downstream gene expression pathways could be up/downregulated. Understanding the mechanism behind these interactions can provide information to modulate the next generation of nanoscale-delivery platform to regenerate the required tissue and organ. AuNPs have been reported to modulate cardiomyocytes through calcium oscillation (during heating the NPs via 532 nm picosecond pulsed laser) [230], myotube activation by wireless stimulation using NIR technique for muscle TE [231], migration of M2 macrophages, and proliferation of neonatal cardiomyocytes using electrical stimulation [232]. Silver NPs (AgNPs) are the other widely used metallic NPs due to their significant antimicrobial characteristics along other distinctive physical and chemical properties, including electrical and thermal conductivity, chemical stability, and catalytic activity [233]. Addition of AgNPs into tissue-engineered constructs with high risk of infection can enhance the success of tissue/organ regeneration. Various methods based on chemical, physical, photochemical, and biological processes have been developed for the synthesis of AgNPs that were discussed in detail in literature [233]. Each method provides its own pros and cons associated with particle size, distribution, stability, and production costs. Antibacterial mechanism of AgNPs is reported to impact through precipitating the bacterial cellular proteins and blocking the microbial respiratory chain system [234]. Utilization of AgNPs in TE can be seen widely for skin-TE; especially for wound-healing applications. Chronic and

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burn-related wounds are highly susceptible to undergo infection and can be associated with multidrug resistance due to infection. Consequently, applying alternative approaches and materials such as utilization of AgNPs can provide an antimicrobial platform to be used for TE applications. AgNPs were incorporated into various scaffolds, including fibers [235
237], membranes/films [238,239], porous scaffolds [240
242], and hydrogels [243]. Impact of AgNPs in wound healing is not limited to the antimicrobial effect as accelerated diabetic wound healing and decrease in local and systemic inflammatory response through modulating cytokines were reported in mice treated with AgNPs [244,245]. Application of AgNPs within the hydrogel systems can promise better wound healing due to higher water uptake content of hydrogels. γ-PGA
based hydrogels coated with AgNPs were applied in mice and intact epidermis and collagen deposition during 14 days of impaired wound healing were observed in histological analysis (Fig. 5.3A and B) [246]. AgNPs have been reported to be useful for other TE and therapeutic applications. These include cytotoxic effect over cancer cells [247], podocyte differentiation of mouse kidney-derived stem cells [248], blood clotting capacity through denaturing the anticoagulant proteins [249], and structural stability and biocompatibility enhancement of decellularized tissue through cross-linking [250]. Antibacterial metal utilization is not limited to AuNPs and AgNPs. Copper NPs (CuNPs) have antimicrobial properties, and they were reported to facilitate wound healing through cell proliferation, angiogenesis, regulation of collagen, and synthesis of elastin fibers [251,252]. Yeast-immobilized and CuNP-dispersed carbon nanofibers were used to develop dual functional wound dressing material for diabetic wounds with an ability of controlling both glucose and bacterial infection simultaneously [251]. In another study, antibacterial, antioxidant, and scavenging properties of CuNPs were compared to AgNPs and TiO2 NPs which can provide comparative data for choosing the desired metallic NPs for a specific application [251]. Quantum dots Quantum dots (QDs) which are colloidal semiconductor nanocrystals with inorganic core were first introduced by Bell Labs’s Louis Brus in 1985. QDs have bright fluorescence, narrow emission, broad UV excitation, and high photo stability. QDs utilization includes enhancing biological imaging, medical diagnostics, cancer therapy, and biosensors [4]. Core inorganic part of QDs which can be from group II
VI elemental semiconductors such as CdSe and CdTe compounds, or group III
V semiconductors such as InP and InAs compounds, cannot interact with biological systems; hence, coating material with appropriate biocompatibility to encapsulate the core part can remedy this drawback. Various methods and materials were reported for the synthesis of QDs [253,254]. To minimize the cytotoxicity, Cd-free QDs and nonmetal QDs (silicon, carbon, and graphene) were synthesized [255]. Stability, nontoxic, and slow photobleaching characteristics of QDs have been reported, and they were used as a platform to follow lineage-tracing experiments for embryogenesis [256]. These methods could provide fundamental information to understand tissue and organ formation and application for optimum desired constructs for regeneration. QDs were utilized for microvasculature imaging

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Figure 5.3 (A) In vivo experiments for the treatment of wound infections in mice with gauze and different γ-PGA/Ag PECs hydrogels. (B) Histopathological evaluation of skin sections. Micrographs of H&E stained wounds treated with gauze and different γ-PGA/Ag dressings at different time intervals (4, 7, 10, and 14 days) (magnification 100 3 ). PGA, Poly(glycolic acid). Source: Reproduced from Wang Y, Dou C, He G, Ban L, Huang L, Li Z, et al. Biomedical potential of ultrafine Ag nanoparticles coated on poly (gamma-glutamic acid) hydrogel with special reference to wound healing. Nanomaterials 2018;8(5).

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to understand their interaction with stem cells [257]. Various methods were used for functionalizing QDs either for enhancing their stability or targeted delivery [258]. QDs were used to target various biological structures such as exosomes [259] microRNA [260] and stem cells [261,262]. For example, QDs conjugated with antibodies were used to track specific cells belonging to rare populations of hematopoietic stem and progenitor cells [261]. Similar systems could be useful in context of tissue/organ regeneration to understand and track cell
cell and cell
environment interactions. Surface of QDs was reported to be functionalized with folic acid for the delivery of therapeutic agents [263]. Similar systems can be developed for TE applications for the delivery of bioactive agents to regenerate desired tissues and organs. Other types of QDs have shown promising effect over cellular activities such as stem cell differentiation. Graphene QDs were reported to enhance osteogenic genes expression and specific biomarkers in MSCs through BMP and TGF-β relative signaling pathways [264].

5.3.1.2 Nanotubes NTs are self-assembled inorganic or organic sheets of atoms that form hollow tubular nanostructures with large internal volume. The tubular form of NTs can be formed from a single sheet (single-wall NTs) or multiple layers of sheets (multiwall NTs) [265]. They can be synthesized from various materials but are favored in using 2D layered compounds such as graphite, hexagonal boron nitride, and tungsten disulfide (WS2) [266]. Composition and geometry factors play a crucial role in determining the properties of NTs [265]. Various materials such as metals, clays, and carbon NTs (CNTs) utilized for TE and their recent applications will be discussed.

5.3.1.2.1 Carbon nanotubes CNTs are most widely used NTs which have been applied in many fields since their discovery in 1991 due to novel chemical, mechanical, and electrical properties. Based on their geometry, it is possible to categorize CNTs as single-walled CNTs (SWCNTs) and multiwall CNTs (MWCNTs) which can be synthesized through various methods such as chemical vapor deposition, arc discharge, and laser ablation [36]. CNTs cover a wide range of applications such as controlled drug delivery, targeted drug delivery, imaging, regenerative medicine, and biosensors. Various tissue regeneration applications were reported using CNTs. Myocardial tissue regeneration is widely targeted by CNTs due to their ability to be manufactured with aligned and conductive properties required for myocardial regeneration [36]. Recently, 3D elastomeric scaffolds
based on poly(dimethylsiloxane) (PDMS) integrated with MWCNTs have been developed. Neonatal rat ventricular myocytes seeded on scaffolds had shown higher viability, more defined and mature sarcomeric phenotype, enhanced connexin-43 gene expression, and increase in gap junction areas compared to 3D-PDMS control [37]. In another study, biocompatible CNT
silk hybrid fibers were manufactured for directed cardiomyocytes growth and electrophysiological detection [38]. Cardiomyocytes cultured on the scaffolds were grown in the

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direction along the fibers. Stable and regular current traces suggested that robust interfaces were formed between CNT
silk and cardiomyocytes. CNTs applications are not limited to the regeneration of cardiac muscle tissue. Regeneration of various tissues using CNTs was proposed such as meniscus [267], bone [39,268
270], cartilage [40], and nerve [41,271
273]. Phosphonates and poly(aminobenzene sulfonic acid) functionalized SWCNTs were produced as supporting scaffold for the growth of artificial bone. Negatively charged functional groups on SWNTs attract the calcium cations and result in the formation of well-aligned plate-shaped HAp crystals [274]. Recently, multilayered nanocomposites based on negatively charged MWCNTs and positively charged poly(dimethyldiallylammonium chloride) were fabricated through layer-by-layer assembly method to prevent aggregation of MWCNTs. Various cellular activities, including metabolic activity, adhesion, differentiation, neurite outgrowth, and electrophysiological maturation, of neural stem cell
derived neurons on nanocomposites have been achieved (Fig. 5.4A
E) [273].

5.3.1.2.2 Clay nanotubes Clay minerals, also known as sheet silicate or phyllosilicate, are inorganic layered nanoscale materials with a wide range of applications in drug/gene delivery, polymer composite formation, and regenerative medicine [275]. They can be classified into 1:1 and 2:1 types according to the layering of tetrahedral and octahedral sheets. Kaolinite and halloysite (1:1) and pyrophyllite and talc (2:1) are among the mineral clays which can be used in regenerative medicine [275]. Halloysite NTs (HNTs) can provide sustained release of bioactive molecules and enhance mechanical properties of scaffolds especially designed for wound-healing applications [276
280]. HNTs with ALP enzyme incorporated have shown higher biomineralization potency which could be used as a bioactive component of the scaffolds for bone repair [281]. Surface-modified HNTs with supermagnetic iron oxide, chitosan, and 2D calcium-phosphate nanoflakes were developed. Elements of modified surface were reported to contribute synergistically in enhancing the osteogenic differentiation of human adipose tissue
derived MSCs [282]. A sustained release of kartogenin from kartogenin-loaded HNTs incorporated in laponite hydrogels for at least 38 day was achieved for cartilage TE [283]. Hydrogel based on gellan gum, glycerol, and HNTs was developed for soft tissue applications [284]. Incorporation of HNTs results in improved biocompatibility of the carrier and provided a platform for tuning compressive modulus of the hydrogel. Higher metabolic activity of human fibroblast was observed in 25% HNTs bearing groups compared to control groups.

5.3.1.2.3 Metallic nanotubes Metallic NTs (MNTs) are another class of nanoscale-delivery systems that have been rarely used for tissue and organ regeneration. Electroless plating protocols with effective homogeneous deposits are one of the versatile routes toward the fabrication of parallel arrays of angular metal NTs using chemically and thermally robust mica templates [285]. MNTs have been used in a wide range of applications such as bimetallic Cu/AuNTs for reusable electrochemical sensing of glucose [286].

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Figure 5.4 (A) Neural stem cells (NSC) differentiation on CNT-multilayered substrates which were fabricated using the layer-by-layer assembly technique. E14.5 NSCs (passage 2) were cultured on the CNT-multilayered substrates in the differentiation medium. (B
E) NSC differentiation on various substrates in vitro. (B) Immunohistochemical staining of neurons (red, β III tubulin) and astrocytes (green, glial fibrillary acidic protein (GFAP)). Cell nuclei were stained using 40 ,6-diamidino-2-phenylindole (DAPI). The differentiation proportion of (C) neurons (D) and astrocytes (E) and viability of NSCs on different substrates were statistically analyzed. n 5 3 experiments,  P , .05. Scale bars: 50 μm. CNTs, Carbon nanotubes. Source: Reproduced from Shao H, Li T, Zhu R, Xu X, Yu J, Chen S, et al. Carbon nanotube multilayered nanocomposites as multifunctional substrates for actuating neuronal differentiation and functions of neural stem cells. Biomaterials 2018;175:93
109 with permission from Elsevier.

AuNTs were also used as safe in vivo imaging nanoprobes and photothermal conversion agents which can be cleared within 72 hours postintravenous injection through hepatobiliary pathway [287]. pH sensitive iron oxide NTs were also developed which can provide magnetically guided delivery system for insoluble

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anticancer drugs. Similar systems can be developed for TE applications where internalization of drug is a priority of the system as NTs could be quickly and extensively internalized [288].

5.3.1.3 Nanorods Nanorods (NRs) or nanowires (NWs) are one of the morphologies nanoscale objects can have. NRs and NWs are different in terms of aspect ratio (height-to-diameter ratio). They can be synthesized through electrochemical, template method, or seeded growth method [289]. In this section, the application of NRs/NWs with an emphasis on tissue/organ regeneration will be discussed.

5.3.1.3.1 Metallic nanorods Metallic NRs are generally manufactured from Au and Ag. AgNRs were generally used as antibacterial purposes and for biosensor applications. AgNRs were reported to be synthesized using cellulose nanocrystals both as reducing and stabilizing agent [290]. They have shown remarkable antibacterial activity toward Gram-positive and Gram-negative bacteria with no cytotoxic effect on human liver hepatocellular carcinoma (HepG2), human breast adenocarcinoma (Mcf7), and kidney normal (BHK) cells. AuNRs were reported to stimulate bullfrog sciatic nerves with minimal invasive approach utilizing 808 nm NIR laser that activates AuNRs and, consequently, promotes activity of nerves in deep layers of the tissue [42].

5.3.1.3.2 Composite nanorods Composite NRs are generally preferred to enhance biocompatibility and precise targeting of the delivery system to tissues. Composites are obtained either by coating NRs or through making the hybrid material which can diversify their applications through providing binding moieties for conjugation of bioactive agents. Poly(vinylpyrrolidone) as immunostimulator and PEG as immune response and degradation rate modulator were used as coating material for fabrication of AgNRs. These AgNRs were reported to elevate immunity of HIV vaccine through surface modification
dependent immunostimulation and overcome the toxicity of the adjuvant [291]. HAp NRs were used with graphene nanosheets as a composite for bone TE [292]. Composite material has shown good osseointegration characteristic with tissues in milieu, enhanced biocompatibility, and superior osteoblastic proliferation induction compared to pristine graphene oxide and HAp alone. Hybrid material composed of AuNRs with albumin electrospun fibers were developed as cardiac patch for suture-free engraftment [293]. Cardiac cells seeded scaffolds were positioned on the myocardium and irradiated with a NIR laser (808 nm) and AuNRs converted the energy into thermal energy. The thermal energy was reported to change the molecular structure of fibrous scaffold and attached it to the wall of heart, preventing any additional injury for patient caused by stitching procedure. AuNRs functionalized via PEG with a terminal amine can effect MSCs differentiation in 2D multifunctional nanocomposite system [294]. Neural differentiation was enhanced when the cells were seeded on the AuNRs compared to the tissue-culture surface. Electric potential difference
derived electrical fields are known to impact

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skin regeneration through modulating cellular behavior [295]. Unidirectionally aligned zinc oxide NRs
based piezoelectric patches were developed to generate piezoelectric potential through mechanical deformations generated by animal motion which could induce electrical field for wound healing. In vitro experiments have shown enhancement in proliferation, myofibroblastic differentiation and migration of dermal fibroblasts, and migration of keratinocytes for aligned piezoelectric patches compared to control groups. In vivo experiments using aligned piezoelectric patches on mice with wounds in their back showed that the wound size significantly decreased via enhanced skin remodeling and regeneration and modulation of inflammatory responses compared to control groups.

5.3.1.4 Nanocages Nanocages (NCs) are organosilicon compounds that can be in cubic metallosiloxanes, cage-like silsesquioxanes, and macromolecular NCs. Organosilicons have been utilized in many fields, including medical devices, protective coatings, sorbents, adhesives, lubricants, and catalysts, and their synthesis was reviewed in literature [296]. Applications of these NCs for tissue/organ regeneration will be discussed in this section.

5.3.1.4.1 Metallic nanocages Metallic nanocages are another candidate nanoscale-delivery system that have been applied in the context of tissue regeneration; however, their utilization is limited to gold nanocages (AuNCs) and gold-silver nanocages (AuAgNCs). AuNCs were reported to be synthesized by controlling the stoichiometry in a manner that localized surface plasmon resonance of AuNCs would be in the range of 600
1200 nm for optical imaging purposes [297]. AuNCs and AuAgNCs were developed as contrast agent for photoacoustic imaging and photothermal therapy [298,299], antibacterial [300], cell adhesive smart surfaces [301], and controlled delivery [302]. AuNCs were used for labeling MSCs for in vivo and in vitro applications and tracking these cells utilizing two-photon and photoacoustic microscopy techniques [303]. Use of AuNPs made possible to track MSCs for at least 28 days in culture with minimum cytotoxicity and without affecting the differentiation potential of MSCs into desired lineages. Such a system could be applied to track the fate of stem cells for tissue/organ regeneration.

5.3.1.4.2 Organosilicon nanocages Organosilicon-based NCs are composite material that can be formed from silsesquioxane family with chemical structure of RnSinO1.5n [304]. The structure consists of inorganic framework of silicon and oxygen atoms, surrounded by organic side chains with R group such as hydrogen, alkyl, alkene, aryl, and arylene functional groups. NCs-based on polyhedral oligomeric silsesquioxanes (POSS) have a regular 3D shape formed by a few units, each containing silsesquioxane. These composite materials have widely been used for tissue-regeneration purpose. Nanocomposite scaffolds
based on the poly(carbonate urethane) (PCU)
POSS were reported to enhance cell viability, total DNA amount, collagen, and sGAG protein of chondrocytes during 14 days of incubation compared to PCL scaffold (Fig. 5.5A
E) [43].

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Figure 5.5 (A) Clinical applications of POSS
PCU. (1) A schematic diagram of POSS nanocages. (2) The chemical structure of POSS
PCU with the poly(carbonate) back bone and POSS appendages. (3) POSS
PCU polymer fabricated as a coronary artery bypass graft implanted in the sheep animal model. (4) Photograph of the first human implantation of the POSS
PCU lachrymal duct system. (5) Photograph of the POSS
PCU tracheal replacement. (B
E) Graphical data demonstrating the cytocompatibility of POSS
PCU compared with PCU and PCL. (B) Cell viability, (C) Total DNA, (D) Cytotoxicity, (E) ECM protein production. POSS, Polyhedral oligomeric silsesquioxanes; PCL, Poly(ε-caprolactone); PCU, poly(carbonate urethane). Source: Reproduced from Oseni AO, Butler PE, Seifalian AM. The application of POSS nanostructures in cartilage tissue engineering: the chondrocyte response to nanoscale geometry. J Tissue Eng Regenerative Med 2015;9(11):27
38 with permission from John Wiley and Sons.

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In similar studies, PCU
POSS was reported to differentiate ADSCs to chondrogenic phenotype [44] while plasma treated PCU
POSS caused osteogenic differentiation of ADSCs [45]. Recently, nanocomposite membranes composed of chitosan and POSS were developed for guided bone tissue regeneration [305]. Incorporation of POSS was reported to enhance the ultimate tensile strength of the membrane. POSS has increased plasma protein adsorption on the surface of membrane and ALP activity of Saos-2 cells seeded on these membranes. POSS were also used for auricular cartilage reconstruction [306], laryngeal reconstruction [307], enhancement of blood compatibility of PU nanofiber membrane [308], and coating of QDs as photo-stability enhancer [309].

5.4

Emerging delivery technologies in nanoparticle area

Tissue/organ regeneration always requires precise biomimetic approaches while providing cost-effective and applicable designs. This approach includes manufacturing both nanoscale-delivery systems and scaffolds. Nowadays microfluidic systems and 3D-printing technologies are being used to enhance the regeneration field toward more precise delivery systems and native tissue
like constructs. In this section, applications of these technologies in the tissue/organ regeneration will be covered.

5.4.1 Microfluidic devices for production of nanoparticles Microfluidic technology takes the advantage of nano/microscale manufacturing method for developing the fluidic microenvironment in controlled and reproducible manner. Microfluidic systems can eliminate the conventional drug-delivery problems, including batch to batch variation, fabrication quality control, heterogeneous particles, and restriction in mass production [310]. Another demand for developing microfluidic technologies in TE field arises from urgent needs for in vitro biomimetic models of human organs which are called organ-on-a-chip. Such systems could minimize the number of animal experiments and provide better tissue and organ mimicking models. In addition, such systems will provide an easily applicable platform to test and develop better targeted delivery systems and apply the combination of vast bioactive agents to understand their effects in tissue/organ using minimum amount of sample and media volumes. In the context of this chapter, recent advances in the application of nanoscale-delivery platforms using microfluidic approach will be covered. While most of the works in the microfluidic systems are focused in cancer therapy and effect of drugs on tumor, these systems can inspire the developments for regeneration purposes. In a study, members of cationic poly(beta-amino ester) family were used with negatively charged plasmid DNA to form self-assembled polyplex NPs for transfecting cancer cells in microfluidic device for continuous and scalable production [311]. Enhancement in in vitro efficacy in multiple cancer cell lines compared to NPs produced by bulk mixing was

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observed. Similar approaches can be used for gene delivery for TE applications. Size of nanoscale-delivery systems plays a crucial role for targeting tissues or passing through the cells [312], or controlled release of the cargo [313,314] which all can be modulated using microfluidic systems to produce desired nanoscale-delivery system. Due to large pore size of alginate nanogels in conventional manufacturing procedure, burst release and low encapsulation are inevitable. It has been reported that microfluidic approach with on-chip hydrodynamic flow could adjust the pore size [313]. Entrapment of bovine serum albumin in the alginate nanogels was reported to yield high encapsulation efficiency and high protein release with slow release profile compared to bulk produced alginate nanogels. Such a system can be used for delivery of GFs and other soluble cytokines in the context of tissue/organ regeneration. Dynamic fluid characteristic of microfluidic systems not only enhance the transport of the NPs deep in model tissues [315] but also could mimic the physiological microenvironment for tissue regeneration. One of the most important reasons for the utilization of microfluidic systems is the necessity for standard platforms for the evaluation of cytotoxicity of the NPs used in TE applications. In one study, PLA-patterned electrospun fibers integrated with PDMS microfluidic system. Hepatocytes spheroids were formed in patterned chip under dynamic condition [316]. Nanotoxicity effect of AgNPs was investigated under dynamic and static conditions. Hepatocytes have shown higher lactate dehydrogenase leakage (cytotoxic marker) under dynamic condition. Recently, microfluidic devices were also used for understanding the effect of cell type, particle charge, and NPs internalization in mouse cortical neurons and SH-SY5Y cells and mechanism of uptake [317].

5.4.2 Recruiting 3D printing and nanoparticles for tissue engineering applications Additive manufacturing technologies has gained significant attraction for tissue/ organ regeneration. This technology not only promises the personalized tissue/ organ design and regeneration but also can be used for establishing disease models, educational, and medical researches [318]. Various 3D printing systems have been established since the 1980s by the introduction of stereolithography which ignite the engines for advances in additive manufacturing till nowadays, but further progress is still needed for optimizing 3D printing for regeneration of functional organs. Different materials such as metals, ceramics, polymers, and composites can be used for 3D printing through using suitable manufacturing procedure that have been reviewed in literature [318
320]. Incorporation of various nanoscale materials into 3D printed materials could modulate the scaffolds through enhancing their mechanical properties, delivery of bioactive agents, or directly modulating cellular activities. Utilization of incorporated nanoscale materials in 3D printed scaffolds and their impact in tissue/organ regeneration will be covered in this section. 3D printing of electroconductive materials such as SWCNTs [321], MWCNTs [322], poly(3,4-ethylenedioxythiophene): poly(styrene sulfonate) [323] and

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graphene oxide [324] for the transmission of electrical stimulation to cells and their neuronal differentiation have gained attentions in the context of neural TE. Scaffolds made of MWCNTs and functionalized with amine groups in a PEG diacrylate (PEGDA) matrix were developed using stereolithography 3D printing [322]. Scaffolds with adjustable pore size were reported to be manufactured. Incorporation of MWCNTs in scaffolds have enhanced neural stem cell proliferation and early neuronal differentiation while applying biphasic pulse stimulation with 500 μA current promoted neuronal maturity later. Applications of 3D printing for bone TE are focused in developing composite inks to enhance mechanical properties, MSCs proliferation, and their osteogenic differentiation. Incorporation of HAp NPs [325], aspirin-laden liposome composite [326], MWCNTs [327], nBGs [328,329], and AuNPs [330] into various carriers are among the studies focused on bone tissue regeneration. To enhance the poor conductivity and delayed efficient electrical coupling which can be seen between adjacent cardiac cells AuNRs incorporated gelatin methacryloyl
based bioinks were developed [331]. Cardiac cells showed higher cell adhesion and better organization when compared to the constructs without AuNRs. AuNRs were reported to bridge the electrically resistant pore walls of polymers and enhance cell
cell coupling and synchronized contraction of the scaffolds. Synthetic poly(urethane)
based 3D printing feeder has been used recently in the context of TE due to suitable mechanical properties, high biocompatibility, and tunable chemical features [332
334]. Poly(urethane)-based elastic NPs were utilized with hyaluronan as a cell aggregation enhancer, and Y27632 inhibitor (TGF-β3 alternative) to develop feeder materials for water-based 3D printing to enhance chondrogenic differentiation of MSCs for customized cartilage TE [334]. MSCs seeded in the scaffolds were self-assembled into MSC aggregates and underwent chondrogenesis effectively.

5.5

Conclusion and future perspectives

World of nanoscale delivery will help to shape our perspective for optimizing the tissue/organ regeneration. Cellular activities, including proliferation, migration, adhesion, and differentiation, can be influenced in nanoscale either through the delivery of bioactive agents or effecting cell fate through modulating scaffolds through providing various surfaces for controlling cell
material interaction or exerting various signaling factors such as electrical, thermal, and mechanical stimuli. Optimum tissue/organ regeneration in future requires integrating and monitoring the nanoscale-delivery systems with minimum cytotoxicity and highly controlled and targeted delivery of bioactive agents that can facilitate regeneration in cost-effective and fast manner. This approach is possible for utilizing various materials with multimodal properties or combining different types of materials through utilizing the advantages of each type to construct desired tissue and organ.

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68.

Surface functionalization of biomaterials for cell biology applications

6

E. Ada Cavalcanti-Adam and Wenqian Feng Max Planck Institute for Medical Research, Heidelberg, Germany

6.1

Introduction

In regenerative medicine, the integration of materials within tissue and the subsequent growth of tissue at the interface with the material are key processes, which determines the successful outcome. The development of new materials for tissue engineering applications is benefiting from a growing body of information on research regarding the molecular and physical cues coded within the extracellular environment. Information derived from the biological functions of the extracellular environment has been used in the design of bioinspired materials and scaffolds. The financial aspect in generating these materials that are often efficacious but require sophisticated preparation remains a challenge. Thus research has been fostered in recapitulating the extracellular environment in its simplified forms by using chemistry to direct cell responses. However, while materials able to optimize cell adhesion are of great advantage, there is still demand for the development of surface chemistry to achieve stability of surface properties and avoid adverse effects. The application of physicochemical concepts and modern surface nanoscience to address different aspects of cell responses has proven to be a successful approach for the design of bioinspired surfaces and materials. The first step is to determine the interaction of cells with such substrates with a focus on the impact of biochemical and physical properties of different materials on cell adhesion and functions. To further understand the process of tissue formation, it is necessary to develop novel strategies to immobilize adhesive and growth factors on materials to induce cell differentiation. A current challenge is to achieve a local presentation of these molecules in such a way that spatial and temporal controls are taking place at the interface to guide cell responses. In the future these promising advances will contribute to clinical and commercial aspects involved in tissue engineering.

6.1.1 From artificial to bioinspired materials: challenges at the cell
material interface The increase in aging population worldwide has led over the last decades to increasing treatments for trauma and regenerative therapies. As a result, there is a high Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00004-0 © 2020 Elsevier Ltd. All rights reserved.

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demand for materials, such as implants and scaffolds, which are able to guide the responses of tissues and cells. The materials used so far can be artificial, consisting, for example, of metal, ceramics, polymers, or can include natural components of tissues. To this end, bioinspired materials have been designed and fabricated with the scope of mimicking the natural properties of biological tissues to direct cell responses [1]. Up to date, tissue engineering products and materials are already in clinical use for skin replacement; there are also a growing number of clinical trials for other products that are applied to different tissues, for example, cartilage, bone, and blood vessels. While the benefits of such approaches are becoming increasingly evident, the costs related to the production of these materials often represent an obstacle to their commercial success. Aside from issues due to difficulties in scaling up and controlling the shelf life of products, a central dilemma is how to influence cell behavior while minimizing the complexity of information coded in the materials, thus avoiding the making of overengineered sophisticated devices. An emerging trend in tissue engineering is the design of a new generation of bioinspired materials which, rather than resembling the complexity of natural tissues, present key components that are able to foster and guide cell responses and unlock the body’s innate ability of growth and repair. Of particular interest is the presentation of chemical, physical, and biological cues at the interface between materials and cells. To achieve control over the different signals, scientists and engineers have been inspired by the multiscale structure of tissues that confers them multiple functionalities [2]. Bone is an example of a hierarchically structured tissue, presenting an organization, which ranges from macroscale down to nanoscale organization. While the structures at the micrometer scale in bone may vary depending on the type of bone, mineralized collagen fibrils are highly conserved and constitute the primary building blocks of bone [3]. Such microarchitecture of bone might provide load bearing properties and respond to load regulation [4]. Here, nanotopography is present because of the fibrillar nature of the extracellular matrix, having collagen fibers a characteristic 66 nm axial period and presenting different mechanical properties in overlay and gap regions. The latter may represent nucleation sites for hydroxyapatite crystals, leading to matrix mineralization. Thus understanding the full potential of hierarchical structures present in the body and achieving the optimization of cues at the cell
material interface might be the key for developing efficient and simple strategies for tissue regeneration.

6.2

Engineering the cell
material interface

The molecular and physical features of the extracellular environment are known to influence cell responses. Thus common design of biomaterials comprises chemical and/or physical cues that mimic the different properties of the extracellular matrix (ECM) of target tissues (Fig. 6.1). Molecular features include the presentation of matrix molecules in their diversity in composition and organization, as well as the delivery of soluble factors, which could be either released from materials or

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Figure 6.1 Properties of engineered materials to control cell responses. The materials can present different mechanical properties, ranging from few Pa to several MPa. The biochemical properties can be varied using strategies for adsorbing or grafting proteins, fragments of proteins, or peptides. Finally, the spatiotemporal properties can be adjusted to achieve targeted degradation of the material, or controlled presentation of molecules at the micro- and nanoscale.

immobilized on surfaces [5,6]. Regarding the physical features, particular attention has been posed on the importance of ECM topography and stiffness that can tightly influence mechanical signaling in cells [7,8].

6.2.1 Surface mimics of the extracellular matrix Bioinspired materials are designed to interact with cells by mimicking key molecular features of the ECM. The ECM present in solid tissues contains different types and amounts of macromolecules, such as proteoglycans, collagens, glycoproteins, as well as sequestered growth factors. The biological activity of the ECM strongly depends not only on the type of ECM proteins presented to cells, but also on the type of molecular information within the protein. A prominent example is represented by specific motifs within ECM proteins that are recognized by and bind to cell surface receptors such as integrins. The variety in integrin receptor types results in binding to a diverse range of ECM ligands and in triggering a multitude of signaling cascades in cells [9,10]. The signals due to integrin
ECM interactions are indispensable for tissue formation and development: for example, the absence of the integrin β1 subunit or of fibronectin is lethal at the early embryonic stage [11].

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Purified ECM proteins can be used as surface coatings or included in scaffolds to provide sites for integrin attachment. For example, modular ECM systems, based on derivatives of gelatin (a denaturation product of collagen) and hyaluronic acid, are applied for in vivo and in vitro growth of cells [12]. The different ECM components might be combined in varying proportions, cross-linked with synthetic polymers such as poly(ethylene glycol) (PEG) diacrylate and supplemented with heparin-bound growth factors. The use of composite ECM protein coatings and scaffolds still finds limitations due to the difficulty in (1) controlling the local amount and distribution of the different components, (2) performing purification and processing of the different molecules without affecting their biological activities. As an alternative, bioinert materials have been functionalized to achieve bioactivity in a more controlled fashion. Protein-repellent materials, such as PEG hydrogels, can effectively be used as a blank template onto which different biological cues can be built with precision and minimal modifications. For example, with this type of materials, small functional domains of ECM proteins, such as the integrin-binding arginine
glycine
aspartic acid (RGD) sequence, are used in place of the full protein [13]. When immobilized onto inert materials, RGD peptides are able to support the adhesion and growth of cells. The efficacy of peptides, ECMmimetics, and fragments has to be further demonstrated in clinically relevant models. In tissues, cells remodel the ECM during homeostasis, embryonic development, and healing. Matrix remodeling involves the concerted action of proteases, for example, metalloproteases, which digest the matrix, and deposition of new matrix proteins, then arranged in a network. These processes require the generation of active traction forces and mechanotransduction in cells, which in turn couples the assembly of matrix and cell locomotion within the ECM [14]. As such, cell movement must occur in balance with the physical properties of the extracellular environment. Emerging tools based on the use of synthetic and natural materials allow the mechanistic regulation and analysis of relevant parameters for cell adhesion and motility. For example, in hydrogels cross-linked by using enzyme degradable peptide sequences, degradation and adhesion processes are combined, resembling processes that take place during in vivo ECM remodeling [15].

6.2.2 Chemical and spatial control of cell adhesion to surface materials Most adherent cells express different integrin types to enable adhesion to ECM proteins. The cell begins to flatten on the interface only when the interaction of the integrins with specific sequences present in ECM proteins is initiated. To create a local environment suitable for the attachment of living cells on material surfaces, proteins from the ECM that have been locally deposited for promoting cell attachment represent a suitable starting. These immobilized proteins possess active, accessible, and specific recognition sites, thus exerting their biofunctionality toward cell binding. The deposition of proteins on planar surfaces can be achieved by

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incubating the surfaces in a solution of proteins prior to seeding cells, or by placing the surfaces into contact with cell culture media that is enriched with different proteins from the serum supplement during cell seeding onto the surfaces. Although protein films on surfaces create good platforms for cell adhesion experiments, some limitations affect their applications: the nonspecific adsorption of proteins based on physisorption is not mechanically stable, that is, the cell-adhesion-mediating film is either washed away or replaced by other molecules in a more thermodynamicfavorable way [16]. Moreover, the adsorption of proteins is lacking biostability, that is, proteins are subject to proteolytic degradation over time, which causes the loss in their biological function. The biological instability of these protein coatings remains a big issue even when the biomolecules are cross-linked to the matrix (chemisorption) [17]. An alternative way to guide the adhesion of living cells on artificial surfaces is to use distinct peptide moieties instead of whole proteins to selectively present the recognition sites on the surface. A number of cell-binding peptide sequences, such as RGD, have been grafted to surfaces via covalent binding of amine or carboxylic acid end groups. Self-assembling of thiol/disulfide-conjugated peptides on gold surfaces via stable Au
S bond has been used to guide cell adhesion [18]. Compared to systems that rely on protein physisorption, the covalent immobilization of small peptide moieties to the solid matrix supplies a mechanically and chemically stable platform for cell adhesion, allowing long-term studies on cell responses to the designed surfaces. Benefiting from the small molecules handling, peptide orientation as well as concentration can be easily and excellently controlled. This influences the accessibility and amount of recognition motif for integrin binding, the key factors to mediate cell adhesion. To achieve the control of integrin-mediated cell adhesion down to single-molecule level, a platform of highly regular gold nanoparticles deposited into a nearly perfect hexagonal nanopattern on substrates has been developed [19,20]. Since each gold nanoparticle has a size of approximately 8 nm and is functionalized with thiol-conjugated RGD peptides, this artificial surface provides a template for the binding of single integrin heterodimers. The role of the spatial distribution of single cyclic-RGD peptides on integrin clustering and its impact on cell adhesion and spreading has been then studied by varying the lateral spacing of the gold nanoparticles [20]. In a recent work, such patterns have been transferred to polyacrylamide gels whose rigidity and nanometer-scale distribution of RGD ligands can be precisely and independently controlled [21]. In this setup, it was probed specifically the mechanosensing function of integrins as a function of varying substrate rigidities, and the regulation of adhesion forces through integrin clustering. Building on the previous work that showed the importance of integrin lateral clustering in regulating focal adhesion dynamics [20], clustering is crucial in regulating molecular forces of single receptor
ligand bonds [22], as well as integrin mechanosensing and activation of mechanotransduction signaling. In many studies, fabrication of patterned surfaces in microscale for guiding cell adhesion and growth on the designed locations is highly needed to understand the correlation between control of cell spreading and shape and resulting activation of cell functions, such as gene expression or metabolic activity. To spatially introduce

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these “cell-adhesion” specific functionalities to a substrate, rending it as a cell-patterned surface, many modification techniques at the interface, including microcontact printing, inkjet printing, dip-pen nanolithography, electron-beam lithography, and photolithography, have been explored. For an in-depth discussion of this topic, the interested reader is referred to more comprehensive reviews [23
25]. Interestingly, control over the distribution of ligands on surfaces with micro
nano hierarchical structure has been achieved by self-assembly of diblock copolymer micelles prestructured by e-beam lithography [26], or by partial removal of gold nanoparticles with the aid of photoresist layer under ultrasonication [27]. For fabricating cell micropatterns, besides creating the promoted and specific cell-adhesion regions as discussed before, material surfaces have to possess a protein-resistant background or barrier to prevent unspecific protein adsorption that can trigger cell adhesion or migration (Fig. 6.2). Although bovine serum albumin (BSA) has been used to form a protein layer to resist cell adhesion in serum-free medium for short term, the passivation with BSA is not stable in the presence of other proteins [28]. Among the chemical materials that can be covalently immobilized on surfaces, biocompatible and hydrophilic polymers such as poly(dimethylacrylamide) [29] and poly(oligo ethylene glycol methyl ether methacrylate) [30] show efficient protein-repellent properties. In particular, PEG derivatives are widely used. Such a PEG polymer layer with only a few nanometers in thickness seems sufficient to prevent protein deposition onto glass. To ensure that cell adhesion is mediated entirely by the presentation of integrin ligands at the nanoscale, PEG with a molecular weight of 2000 g/mol has been used to passivate the areas between

Figure 6.2 Spatial control over presentation of adhesive and repelling regions on material surfaces. Cell micro- and nanopatterning is achieved and in this way integrin-mediated adhesion can be controlled. Several approaches can be taken to immobilize adhesive molecules and to create the repelling layer to prevent adhesion and control biding sites and cell shape.

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metal nanoparticles against cell adhesion [19]. Hydrogels with similar chemical surface functionalities have been successfully used as protein-resistant surfacecoating materials as well [31]. Superhydrophobic (SH) surfaces have also been applied to fabricate protein-resistant barriers [32]. Owing to the air trapped in the SH areas to prevent the protein adhesion at liquid
air interface, the SH surfaces exhibit the ability to prevent cell adhesion and migration. In recent years, slippery liquid
infused porous surfaces where a stable, water-immiscible, low-hysteresis lubricant overlay is locked on nanostructured surfaces show an outstanding antibiofouling performance [33]. Ueda and Levkin reported that the micropatterns of the lubricant overlay exhibited long-term stability and excellent cell-repellency with more efficiency than conventional PEG or SH surfaces in controlling eukaryotic cell adhesion [34].

6.3

Delivery strategies for growth factors at the cell
material interface

During tissue healing and embryonic development, growth factors such as members of the TGF-β superfamily are the key in stimulating cells to proliferate and differentiate. In addition, these signaling molecules influence stem cell fate [35]. In the preparation of biomaterials, growth factors are often embedded to direct cell differentiation and functions, often in a combination with cell adhesive proteins. Thus carrier matrices have been developed using either synthetic, such as degradable synthetic polymers based on poly(lactic-co-glycolic) acid, or natural polymers, such as collagen, hyaluronan, or silk fibroin [36,37]. Up to date, the greatest resonance, both financial and mediatic, has been achieved by a Food and Drug Administrationapproved product that is employed in orthopedic surgeries for the treatment of spine fusion. This product provides recombinant human bone morphogenetic protein 2 (BMP-2) added to a collagen sponge. Its use became controversial during the last years because of increasing reports on complications and conflicting results from different clinical studies. The protein is released from the carrier matrix over time, resulting in complications and side effects, such as ectopic bone formation. The rapid local depletion and diffusion of growth factors can be overcome with the increase in the doses provided to efficiently trigger tissue regeneration. This results in the increase of costs and might lead to local inflammation, formation of undesired mineral deposits, and increased incidence of cancer in distant tissues [38,39]. Aside from the presentation of growth factors and chemokines from the surface, materials have also been designed to promote the expression of these molecules and their receptors in cells, in order to achieve controlled differentiation and guide cell functions. To this end, surfaces functionalized with peptides to capture growth factors may represent a possible strategy to potentiate adhesion and autocrine signaling in cells [40]. An additional aspect is the alleviation of inflammatory immune responses to further facilitate the recruitment of cells for regeneration and revascularization.

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6.3.1 Biofunctionalization strategies for tailoring the spatiotemporal delivery of growth factors To overcome the short half-life and release of growth factors in solution, controlledrelease or immobilization strategies have been developed. In tissues, growth factors are bound to ECM components or are part of membrane complexes that are released upon matrix degradation or remodeling [41]. Thus the mode of presentation and release of growth factors from scaffolds and material surfaces might elicit different functions of the proteins and unravel biological functions (Fig. 6.3). The covalent immobilization of growth factors onto surfaces has been a matter of debate, mainly due to concerns over receptor accessibility and loss of bioactivity of surface-immobilized proteins. To overcome these problems, different tethering strategies have been developed to improve the presentation of molecules at the material interface [45]. Covalent tethering of TGFβ1 to PEG not only retains its biological activity for stimulating ECM production but also is more active than the soluble form at comparable concentrations [46]. Sitespecific modification can be engineered through genetic insertion of amino acids to the termini of the growth factor, for example, cysteine-containing fusion tag consisting of short amino acid sequences at the N-terminus of the protein, or enzyme substrate sequences that lead to localized proteolytic release [47]. The immobilization strategies might also involve the use of heterobifunctional cross-linkers that can be tailored in respect to the chemical character of the underlying support. Thus chemical molecules containing an amine-reactive N-hydroxysuccinimide group and a sulfhydryl-reactive maleimide moiety have been used to bind BMP-2 to polycaprolactone scaffolds or chitosan membranes [48]. The covalent immobilization of growth factors brings the additional benefit of preventing internalization of the growth factor
receptor complexes by cells, thereby prolonging growth factor availability and reducing the amount of the protein necessary to trigger biological responses.

Figure 6.3 Platforms created to engineer the dialog between adhesive and growth factors. (A) Growth factors embedded in a polysaccharide network show enhanced downstream signaling, as indicated in the western blot for phosphorylation of the Smad 1/5 complex [42]; (B) directed cell migration is promoted by the copresentation of matrix proteins and growth factors arranged in stripes, shown by fluorescence microscopy images [43]; (C) nanoscale presentation of growth factors covalently immobilized on nanopatterned surfaces: single growth factors are immobilized on the nanoparticles and their spatial arrangement on the surface is mapped by atomic force microscopy [44].

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There is still need for an appropriate design to achieve the spatial control of tethered growth factors. Recent advances in combining growth factor presentation with surface nanopatterning approaches result in molecule surface densities in the range of pg/cm2; here single growth factors are presented to cells at spacing of 30
100 nm [44]. Another important aspect to consider in the design of surface materials for the presentation of growth factors is the in vivo occurrence in spatial gradients that guide regenerative responses, such as during angiogenesis. To this end, the use of porous materials or layered approaches facilitate the formation of growth factor gradients that are presented at the interface with cells [49,50]. Many growth factors are present in vivo either in soluble or in matrix-bound form. For the latter, proteoglycans facilitate the local accumulation of proteins that might be released during matrix remodeling processes. Proteoglycans interact with matrix proteins, forming a structural network that supports and regulates adhesion and integrin binding. At the same time, the direct effect of proteoglycans on the oligomerization of growth factor receptors is controversial. Biomimetic surfaces that comprise the presentation of all these factors in a controlled manner are challenging, both in terms of precise control the spatial presentation and the molecular interactions that lead to enhancement of bioactivity. Only recently [42], heparan surface-BMP-2 binding has been characterized in terms of kinetic affinity, stoichiometry, and conformational changes. This is an important starting point to understand the effect of proteoglycans on growth factor activity. As for today, with these materials it is still difficult to entangle the mere effects arising from the spatial distribution and how these are related to temporal events. The temporal control for the in vivo situation cannot be easily achieved and would require additional efforts in the design of switchable linkers or molecules for a localized and targeted presentation and, when needed, release of the molecules. One promising approach is represented by combining material surface functionalization with growth factor encapsulation in particles [51]. Thermo-responsive PNIPAM-DMAEMA/cellulose sulfate complexes loaded with BMP-2 allow a slow release of the growth factor at physiological temperatures while maintaining its localized biological effect.

6.3.2 Guidance of cell responses by growth factors complexed with surface materials Growth factors regulate cell adhesion and motility. During adhesion to the extracellular matrix, macromolecular complexes comprising integrins and growth factor receptors are formed, resulting in activation of cytoskeleton dynamics [52]. Integrins colocalize and coprecipitate with several growth factor receptors, and there is accumulating evidence of an extensive crosstalk between integrin and TGFβ signaling, both in terms of physical association of receptors and modulation of downstream effectors. For example, TGF-β1 regulates the expression of integrinassociated proteins, such as paxillin, PINCH, and ILK, which increase integrin activation [53]. Several growth factors, including Epidermal growth factor, platelet derived growth factor-BB, and basic fiborblast growth factor, activate extracellular

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signal regulated kinase signaling pathways only upon integrin clustering and ligand occupancy [54]. BMPs enhance the formation of focal adhesions and stress fibers by increasing α5 and β1 integrin expression and trigger migration by enhancing the incorporation of β1 integrin into lipid rafts [55]. Material surfaces presenting combined growth factors and adhesive ligands have been designed to guide and further explore the interplay between growth factor presentation and cell adhesion. A major contribution in the design of materials for regenerative purposes has been the identification of FNIII12
14 as growth factor binding region [56]. This fragment, fused to fibrin matrix in combination with VEGF-A, induces endothelial cell reorganization, whereas in combination with BMP-2 it promotes mesenchymal stem osteogenic differentiation in vitro and in vivo. Importantly, low amounts of these growth factors are sufficient to trigger cellular responses. Such synergy of cytokines and growth factors in adhesion has been also explored in terms of stem cell function. SDF1alpha presentation together with RGD sequences could maintain hematopoietic stem cell self-renewal [57]. An intriguing aspect to consider in this concerted action of growth factors and adhesive ligand is that through the modulation of adhesion and spreading area, resulting selfrenewal and differentiation programs could be achieved in cells at the material interface. This could be translated into the combination of soluble and physical signals that mimic the native microenvironment of stem cells when utilizing biomaterials.

6.4

Conclusion and outlook

Several challenges at the cell
material interface have been pursued over the last years by employing bioinspired materials, designed according to different types of information from the ECM, namely, (1) specific chemical signals and modes of presentation of the molecules, (2) physical properties of the cell environment, and (3) topography of the binding sites of the microenvironment at the nanoscale. The spatial control over the presentation of molecules on material surfaces has been achieved by patterning approaches. These have provided further means to characterize cell receptor clustering and the relevant threshold for the formation of stable contact to the surface, in relation to adhesion force generation. The combination of both nanotopography and ligand presentation allows the regulation of phenotype and adhesion, indicating that nanoscale surface properties should not be neglected but rather optimized for the design of interfaces of scaffolds and implant materials. The immobilization of growth factors represents a novel strategy to direct cell responses while favoring cell adhesion to the material surface. The main focus of research has been on osteogenesis and angiogenesis, given the socioeconomic impact of orthopedic injuries and the need for improving healing. Surfaceimmobilization of growth factors not only prolongs but also elevates the signaling responses that lead to cell differentiation. The potential to achieve controlled

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presentation of these molecules and to reduce the amount of these factors has been demonstrated in combination with surface patterning approaches and offers future advantages for tissue engineering applications. Thus, physicochemical approaches prove to be the right tools not only to understand the influence of microenvironment on cell responses but also to optimize artificial materials making them become biomimetic at different length scales. In summary, developing new strategies for improving cell
material interfaces and for guiding-specific cell responses allow studies on fundamental biological events, such as cell adhesion, differentiation, and migration, which are on the basis of both physiological and pathological processes. Future work should be devoted to combining both adhesive and growth factor cues for directing cell responses through adhesion and investigating the local crosstalk between the distinct signaling pathways during cell differentiation. These types of surfaces would allow studying long-term effects of low surface concentrations of growth factors. In addition, such materials would mimic the naturally occurring coexistence of adhesion promoting molecules and growth factors in close proximity, which might affect the interplay of different types of receptors. This setup might alter the mode of signaling, resulting not only in prolonged cell stimulation but also addressing different receptor complexes. Since cell migration appears to be directed by surface topography and fibrillar organization of the ECM, future directions should also include the investigation of gradients of growth factors on fibrillar matrices.

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69.

Stem cells: sources, properties, and cell types

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Melis Asal1,2 and Sinan Gu¨ven1,2,3 1 Izmir Biomedicine and Genome Center, I˙zmir, Turkey, 2Izmir International Biomedicine and Genome Institute, Dokuz Eylul University, I˙zmir, Turkey, 3Department of Medical Biology, Faculty of Medicine, Dokuz Eylul University, I˙zmir, Turkey

7.1

Introduction

Tissue engineering consists of studies aiming developing functional tissues that can replace absent or diseased tissues. This concept requires utilizing both engineering and life sciences approaches. In order to construct a 3D tissue, cells need to be combined with scaffolds to functionally and mechanically recapitulate the natural system. Therefore choosing ideal cell source is crucial. The utmost improvement in the tissue engineering field occurred with the discovery of stem cells, which present an unlimited cell source. This chapter focuses on providing an insight into stem cells, their sources, properties, and types (Fig. 7.1). We focus on the types of stem cells proper for tissue engineering: induced pluripotent stem cells (iPSCs) and mesenchymal stem cells (MSCs), and lighten their applications in the field of tissue engineering.

7.2

Stem cell properties

Stem cells are the primal, undifferentiated cells of all multicellular organisms. These cells are characterized by being able to renew themselves through cell division and differentiating into a wide range of cells (Fig. 7.2) [2,3]. Stem cells are vital to the development, growth, maintenance, and repair of all tissues and organs. They are present in the human body from the early stages of development until death.

7.2.1 Self-renewal Self-renewal is the ability of stem cells to divide numerous times while retaining their undifferentiated state. This ability ensures expanding the number of stem cells during development and then maintaining these cells in the adult body throughout the life, also to restore the number after traumas [3,4]. Stem cells do this via dividing asymmetrically or symmetrically to give rise to one or two daughter stem cells Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00007-6 © 2020 Elsevier Ltd. All rights reserved.

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Figure 7.1 A summary of various types of stem cells that can be obtained from numerous sources from the human body [1].

Figure 7.2 Asymmetric division of stem cells, which lead to one stem cell and one specialized cell through utilization of the self-renewal and differentiation properties.

retaining potencies similar to the mother cell (Fig. 7.2). The self-renewal capacity of the cells is controlled by cell-intrinsic mechanisms, which are regulated by signals from the cell microenvironment. Called the stem cell niche, this microenvironment preserves stem cells and regulates their function. In response to changes in the niche, stem cells go through changes in the cell cycles that lead to changes in their

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potencies. Even though stem cells possess great self-renewal, they do not selfrenew that extensively under physiological conditions; they usually go through fewer divisions than their potentials. However, in the case of injuries that require higher regeneration, their potential changes in accordance with the changes in the physiological conditions [4].

7.2.2 Potency Stem cells can be isolated from a variety of tissues in the adult body, but their differentiation potentials differ [2]. Potency is the stem cells’ potential to differentiate into different cell types [3]. The cells are categorized mainly into four groups according to their differentiation potential. Totipotent stem cells are produced when an oocyte is fertilized by a sperm and form the zygote [3,5]. The zygote has the potential to differentiate into both embryonic and extraembryonic cell types to develop into a whole organism and to form the structures necessary for development [5]. The cells produced as a result of the first few divisions of the zygote, at least until the four-cell stage embryo, are also totipotent [3,5] (Fig. 7.3). Pluripotent stem cells are the cells that have the capacity to differentiate into all three germ layers: endoderm, mesoderm, and ectoderm, but they cannot develop into a whole organism [5]. The examples of pluripotent stem

Figure 7.3 Summary of stem cells in terms of their potencies.

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cell are embryonic stem cells (ESCs) and iPSCs [6] (Fig. 7.3). Multipotent stem cells can develop into many specialized cells, but that of a certain lineage [7] (Fig. 7.3). The stem cells retained in the adult body are either multipotent or unipotent. Unipotent cells can differentiate only into one cell type but have the selfrenewal capacity, separating them from the nonstem cells [3]. Progenitor cells also often differentiate into one cell type, but they lack the self-renewal capability [7]. From bone marrow transplants to burn treatments and to the treatment of corneal disorders, stem cells are a part of standard clinical care. Due to advances in stem cell research, new possible areas that can benefit from stem cells have emerged. Today’s prevalent idea is to replace aged, damaged, or diseased organs/tissues with their engineered substitutes by employing the tissue engineering approach.

7.3

Cell types

For tissue engineering applications, stem cells have the potential to offer an infinite source of cells. Although all stem cell types can possibly be of value, each type has its own potencies and associated limitations (Table 7.1). Thus, the initial step to use stem cells is to choose the most suitable stem cell type for the tissue to be engineered.

7.3.1 Embryonic stem cells After the eight-cell stage of the development, the blastomeres increasingly lose their totipotency as the cells differentiate into the first two lineages: the inner cell mass Table 7.1 Summary of the stem cell potencies, their sources, types, and features.

Differentiation potential Sources

Types

Totipotent

Pluripotent

Multipotent

A whole organism

The three germ layers ICM of the blastocyst

Cells of a certain lineage Adult tissues

Fertilized oocyte and the following few divisions Zygote Embryo up to fourcell stage G

G

G

Limitations

Ethical problems

G

G

G

ICM, Inner cell mass.

Embryonic stem cells Induced pluripotent stem cells Ethical problems Teratoma formation

Adult stem cells (hematopoietic, mesenchymal, etc.)

G

G

G

Low number Difficulty of isolation Restricted differentiation potential

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(ICM) and the trophectoderm (TE). ICM comprises the cells that develop into the fetus, and TE is destined to form the extraembryonic structures. ESCs are confined in the ICM at the blastocyst stage and are pluripotent, being able to differentiate into endoderm, mesoderm, and ectoderm (Fig. 7.4). In vivo, pluripotent cells within the ICM rapidly differentiate as the development of the fetus goes on. However, these cells can be isolated and cultured in vitro to keep them undifferentiated with the potential to differentiate into any cell type [5]. Their ability for dividing indefinitely allows expansion to high numbers required for development of tissues [8]. In 1981 Martin and Evans were able to derive ESCs from the ICM of the inbred mouse [9,10]. In 1998 human ESCs were isolated by Thomson et al. from In vitro fertilization (IVF)-produced human embryos [11]. Since their discovery, ESCs have received a great amount of interest due to their promising potential in tissue engineering. ESCs can be cultured theoretically indefinitely by maintaining the transcriptional activity and epigenetic regulators supporting pluripotency in vitro. ESCs possess enzymatic activities such as alkaline phosphatase and telomerase and express genes that are downregulated during differentiation, such as OCT4 and NANOG [5]. However in the case of changes in the culture conditions, cells begin to rapidly differentiate, meaning that the actual challenge for using stem cells in tissue engineering is to direct the differentiation to the desired cell types [8]. The ability of ESCs to form teratomas in vivo, tumors consisting of all three germ layers, underlines the importance of transplanting terminally differentiated cells without any hidden pluripotent stem cell properties during the tissue engineering applications. This can be achieved by differentiating the cells directly to the desired cell types in vitro [5,8]. Since diseases result from defects in different types

Figure 7.4 The types of cells that ESCs and iPSCs can give rise to. ESC, Embryonic stem cell; iPSC, induced pluripotent stem cell.

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of cells, ESCs may offer a possible cure due to their ability to generate any type of cells. The biggest problem regarding the use of ESCs in clinical applications is due to ethics. The fact that the embryo needs to be destroyed to isolate the ESCs has led to ethical debates within the authorities. As a consequence, isolating and working with ESCs is rarely permitted today. ESCs can potentially allow personalized treatments through production of autologous ESCs via therapeutic cloning. Therapeutic cloning refers to creating embryos de novo by transferring the differentiated cells’ nuclei to de-nucleated oocytes. In this process, the cytoplast reprograms the nucleus and initiates the embryonic development. If the development is stopped during the blastocyst stage, ESCs can be obtained. Although this technique possesses great potential, the ethical problems still emerge. Ethical debates about the use of human ESCs have been overcome by the discovery of iPSCs.

7.3.2 Induced pluripotent stem cells In 2006 Takahashi and Yamanaka showed that stem cells with properties similar to ESCs could be generated from mouse fibroblasts by simultaneously introducing four genes: OCT4, SOX2, C-MYC, and KLF4 [6]. They named these cells iPSCs. In 2007 they introduced the same four genes to human fibroblasts and generated the human iPSCs [12]. Simultaneously and independently, Thomson et al. generated human iPSC using different factors: OCT4, SOX2, NANOG, and LIN28 [13]. In latter studies, iPSCs were generated from lymphocytes and neurons [14,15]. One of the most important questions regarding iPSCs is whether they are different from ESCs and, if so, whether any differences that do exist are functionally relevant. Although there were prior reports reporting differences between iPSCs and ESCs, in terms of differentially expressed genes, DNA methylation, and epigenetic memories of donor cells, it is concluded that the variations are due to differences in induction and culture conditions, and that iPSCs and ESCs are hardly distinguishable [16] (Fig. 7.4). Preferably, iPSC studies should be conducted in parallel with ESCs.

7.3.2.1 Reprograming techniques IPSCs were generated using retroviruses and lentiviruses in the first studies [16]. In the retroviral approach; the reprograming factor genes carried by the vector are integrated into the host genome. Lentiviruses work similarly but are reported to show less variability and higher efficiency [17]. These methods were proven to give rise to iPSCs effectively. However, due to associated risks such as mutagenesis and immunogenicity, the ideal approach to generate iPSCs has become avoiding vector integration into the host genome. Several protocols have been reported avoiding vector integration such as the use of Sendai virus (SeV), plasmids, transposons, synthetic RNAs, and proteins [16].

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SeV is a negative sense single-stranded RNA virus. SeV is used for the transduction of the RNAs of the OCT4, SOX2, KLF4, and cMYC genes. This reprograming method is reported to be effective and also safe, due to loss of the virus genome at higher passage numbers. Still, this approach is not yet considered appropriate for clinical applications, which restricts its usage [17,18]. In methods that employ plasmids for transfection of the reprograming factors, daily introductions are generally required. However, in the episomal plasmid approach, the plasmid contains EBNA-1 and OriP sequences obtained from the Epstein
Barr virus [17]. These sequences allow extending the expression of reprograming factors by aiding in plasmid replication during cell division. This approach offers developing human iPSCs from fibroblasts as well as blood cells. Although there is a possibility of integration into cell genome, this method is advantageous due to rapid absence of plasmids after reprograming [18]. Transposons are DNA sequences that bear the ability of integration into host genome. In order to induce cells using the PiggyBac transposon, cassettes incorporating reprograming factors are introduced to the host cells. With the employment of transposase for the second time after a successful induction, the cassettes can be discarded from the host genome. This feature makes the PiggyBac approach a nonintegrating one [17]. Synthetic RNAs with modified nucleotides also enable iPSC induction with the introduction of reprograming factor RNAs into host cells. By employing RNAs, human iPSCs were generated from fibroblasts and blood. This approach requires daily introductions due to the short life of RNAs [17]. In addition to this, since foreign RNA introduction induces an immune response, measures need to be taken to diminish it [18]. The rapid removal of foreign RNA sequences makes this approach a nonintegrating one with a lower risk of mutagenesis than the methods that employ DNA [17]. Recruiting proteins to reprogram cells into iPSCs is another technique. This approach utilizes peptides with the cell penetration ability to introduce reprograming factors into host cells [17]. Inhibition of chromatin factors can also be employed to improve reprogramming efficiency of iPSCs. [19]. Since iPSCs can differentiate into the cells of all three germ layers and allow personalized treatment, they show remarkable potential. However, their clinical application still needs to overcome numerous limitations. Thus the use of adult stem cells seems to be a more practicable approach.

7.3.3 Adult stem cells The inevitable progresses occurring in the stem cell field have led to the discovery of the unpredictably great regeneration potentials of adult stem cells. This potential gave rise to the development of new methods for stem cell differentiation to engineer functional tissues. Adult stem cells are undifferentiated cells found in the adult body that possess both self-renewing and differentiation properties. These cells are multipotent and are capable of giving rise to the types of cells of the tissues that they reside in. They conserve the homeostasis by replacing cells lost due to injuries.

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It is still unclear whether stem cells are found in all tissues. The well-known adult stem cells are found in the bones, adipose tissue, blood, muscle, skin, and intestines. The best studied adult stem cells are mostly located in the bone marrow that hosts both hematopoietic and MSCs.

7.3.3.1 Hematopoietic stem cells Stem cells were first discovered by Till, McCullough, and Siminovitch, in 1963, with the discovery of the hematopoietic stem cells (HSCs) as cells that can give rise to fully differentiated cells [20]. HSCs reside in the bone marrow and develop the whole hematopoietic system by giving rise to blood progenitor cells, which then differentiate into mature cells [7,21]. The ability of HSCs to form a whole hematopoietic system was first utilized in the 1960s with successful bone marrow transplantations to treat diseases associated with the immune system. Today HSC transplantation is still employed widely in order to cure cancer, autoimmune diseases, and metabolism disorders among many others. HSC studies nowadays focus on utilizing tissue engineering techniques to generate HSCs from iPSCs in order to enable autologous transplantation [21]. Bone marrow also harbors another well-known stem cell type: MSCs.

7.3.3.2 Mesenchymal stem cells MSCs were discovered by Friedenstein in a series of studies, who described these cells as a type of fibroblastic cells that can produce bone and reticular tissue, found in the bone marrow of mouse and guinea pig [22,23]. After it was discovered that these cells can also give rise to cartilage and adipose, they were named MSCs. MSCs are also called mesenchymal stromal cells as they are part of the stromal tissues that adhere to plastic. However, these cells are still considered stem cells for possessing the two distinct features of the stem cells: self-renewal and differentiation [24] (Fig. 7.5). For cells to be considered as MSCs, they need to be adherent, have fibroblastlike appearance when seeded, and have the capacity to differentiate into three lineages: osteogenic, adipogenic, and chondrogenic. These cells should express CD73, CD90, and CD105 markers and be negative for CD34, CD45, CD14 or CD11b, and CD79α or CD19. It has been shown that MSCs can differentiate into tenocytes, smooth muscle cells, and stromal cells of the bone marrow in addition to osteoblasts, chondrocytes, and adipocytes [25] (Fig. 7.5). MSCs have been studied in numerous studies since their discovery and have been utilized in several fields of tissue engineering. In addition to this, MSCs are seen as a significant cell source with promising clinical results. These cells show promise due to their great differentiation capacity, high proliferation capability, and different sources in the adult body that allow easy isolation.

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Figure 7.5 Differentiation potentials of mesenchymal stem cells.

MSCs have been harvested from many fetal and adult tissues (Table 7.2). Among these sources, bone marrow and adipose tissue are the mostly studied and characterized sources [21,25].

7.3.3.2.1 Bone marrow derived mesenchymal stem cells Bone marrow
derived MSCs (BMSC) reside in the bone marrow and serve as osteogenic progenitors. BMSCs also secrete trophic factors and therefore maintain the HSC niche by regulating HSC function [24]. There is no defined marker to distinguish BMSCs, but they are generally characterized with the expression of CD29, CD44, CD105, and CD166. Then again, they are not positive for hematopoietic lineage markers, such as CD14, CD34, and CD45 [21]. In a study conducted in 2007, SSEA-4 was identified as a marker for human BMSCs [36]. MSCs are known to reside in many other tissues, suggesting that a common MSC niche is present all over the adult body [24].

7.3.3.2.2 Adipose-derived stem cells A population of cells named lipoaspirate cells isolated from human adipose tissue was reported by Zuk et al. in 2001. The cells were separated from the adipose tissue with the aid of collagenase digestion of the stromal vascular fraction (SVF) [28]. These multipotent cells were later named adipose-derived stem cells (ASCs). ASCs are MSCs that reside in the adipose tissue. ASCs were shown to have the potential to differentiate into adipocytes, osteoblasts, chondrocytes, myocytes, neurocytes, endothelial cells, smooth muscle cells, and other cell types ( for more information, please see the chapter 15: Adipose Tissue Engineering, Louis et al.) [37]. Adipose tissue contains numerous cell types. To isolate ASCs, the cells in the SVF are separated according to their expression profile of surface markers [38]. MSC markers expressed by ASCs are CD13, CD29, CD44, CD58, and CD166. Different than other MSCs, ASCs express CD13, CD29, CD49d, CD73, CD90,

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Table 7.2 Mesenchymal stem cell sources. Sources Fetal tissues

Adult tissues

References Amniotic fluid Amniotic membrane Chorion membrane Chorion villi Decidua Placenta Cord blood Wharton’s jelly Umbilical blood Umbilical cord tissue Bone marrow Adipose tissue Peripheral blood Blood vessels Brain Endometrium Menstrual blood Skin Dental pulp Spleen Liver Kidney Lung Heart Cartilage Tendon Skeletal muscle Thymus Pancreas

[25,26] [25,26] [26] [26] [26] [25,26] [26] [26] [26] [27] [26] [28] [26] [25] [29,30] [31] [31,32] [33] [34] [35] [35] [35] [35] [24] [24] [24] [35] [35] [35]

CD133, MHC I, and MHC II but they lack the expression of the BMSC marker, CD106. ASCs are also identified by the lack of CD45 and CD31 expression [37]. ASCs are harvested from human subcutaneous adipose tissue. In the human body, there are two adipose depots with different functions, named the brown adipose tissue (BAT) and white adipose tissue (WAT). BAT offers thermogenic properties. Although studies reporting the presence of BAT in adults have emerged, BAT is mainly found in infants and the amount of it declines with age. WAT is much more abundant than BAT and found throughout the adult body. WAT has energy storage and insulation functions. Although there are known differences between the functions of BAT and WAT, their distinct stem cell contents are still unknown. However, it is discussed that stem cells of the WAT have greater potential for differentiation. Even in terms of WAT, there are differences that arise such

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as that the SVF number of abdomen is higher than that of the flank and axilla [37,38]. Properties of ASCs are also affected by the patient age. It was shown that ASCs obtained from younger patients possessed higher angiogenic and osteogenic properties than older age groups. Decreased proliferation and differentiation was also shown in stem cells obtained from older patients [37]. ASCs appear to be highly promising for tissue engineering applications due to their abundance in the adult body and defined isolation techniques. Unlike ESC, ASCs are nonproblematic in terms of ethics since the cells are usually autologous.

7.3.3.2.3 Bone marrow
derived mesenchymal stem cells and adipose-derived stem cells Although MSCs are found in numerous tissues, the widely employed sources for tissue engineering applications are the BMSCs and ASCs [24]. Bone marrow is harvested from the posterior superior iliac spine or the iliac crest with the aid of a syringe. In general, this procedure is implemented under local sedation. Bone marrow aspiration allows attaining approximately 20 mL of bone marrow. Even though it is considered to be a low-risk procedure, it often causes pain. Bleeding and infection are also associated risks with the procedure. The low cell yield is another drawback [24,37]. Contrarily to BMSCs, ASCs reside throughout the body in subcutaneous adipose tissue. The preferred sites for adipose tissue harvesting are the abdomen and hip region. Existing techniques for ASC harvesting consist of Coleman’s technique, liposuction, and excision [37]. Although the viable cell proportion is significantly higher in liposuction than in excision, there is no difference in cells’ differentiation potential between the two methods [37]. These procedures are typically carried out under general anesthesia, which contributes to the morbidity and mortality rates. Albeit liposuction being minimally invasive, it bears the risks of nerve damage, bleeding, edema, infection, and necrosis. Excision bears even more complications such as problems in wound healing, infection, necrosis, and even death [24]. Even though ASC harvesting seems to be complicated, there are no further risks to isolate ASCs when this procedure is already performed for cosmetic purposes [37]. Moreover, the number of cells that can be harvested has a big impact on choosing one source over another. In bone marrow aspiration, cell numbers are reported to be lesser in contrast to ASCs. Thus BMSCs are usually insufficient for tissue engineering applications in terms of cell number, making ASCs a more practical source [24]. Another difference between ASCs and BMSCs is their tendency to differentiate into distinct cell types. Due to their natural niches, ASCs have an increased capacity to differentiate into adipose cells, and BMSCs are more likely to differentiate into osteogenic cells. It is still noted that these tendencies are affected by the sex, age, and health of the patients, as well as differentiation protocols [24]. Consequently, the choice of the appropriate MSC depends on various factors and it is not possible to say exactly which source is more suitable for a specific application.

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7.3.3.2.4 Risks of mesenchymal stem cells Currently, there are numerous ongoing trials recruiting MSCs, which will shed light on the risks associated with their use. As for the completed trials, no direct relation was found between MSC introduction and health complications such as malignancy, toxicity, and death. Some studies however reported fever, infection, and organ damage [24]. In order to appropriately assess the risks associated with MSCs, long-term follow-ups are required. Although as for now, MSCs are known not to preserve in the body for long term after being introduced systemically. The transient presence of MSCs following introduction causes requirement of high cell numbers that cannot be obtained from a single patient [24].

7.4

Applications of stem cells in tissue engineering

Tissue engineering and stem cell applications keep holding promise for treatment of incurable or challenging diseases. The sources of cells can be autologous, allogeneic, or xenogeneic. Autologous refers to using the host’s own cells, allogeneic means using cells from another individual, and in xenogeneic transplantations, the cells are obtained from an animal other than human. Xenogeneic sources bear immunogenic risks that restrict their use in tissue engineering. Autologous and allogeneic sources on the other hand are widely used in studies. Stem cell therapy can employ stem cells directly or differentiate the cells before transplantation. Stem cells can be used for their differentiation potential as well as regulatory effects such as that of MSCs [39]. Tissue engineering studies employing stem cells mainly involve disease modeling, drug screening applications and treatment of numerous diseases (Fig. 7.6) [40
43]. Umbilical cord blood (UCB) hosts hematopoietic cells that can be transplanted to cure hematological cancers. Numerous UCB banks are found in different countries, with some UCB derivatives being approved by Food and Drug Administration. These products allow human leukocyte antigen (HLA) compatible allogeneic transplantation of the HSCs to the patients following myeloablation [39]. UCB is not only used for blood disorders. Another product, which is used to treat cartilage defects, consists of allogeneic human UCB
derived MSCs and hyaluronic acid (HA). The results show that the use of UCB-derived MSC along with HA has promising value for cartilage regeneration. After 7 years of follow-up, the clinical results were stable, with minor adverse effects and no tumorigenesis, making the use of UCB-derived MSCs safe in addition to being effective [44]. There are also studies that differentiate the cells before transplanting, rather than using stem cells themselves. There are a vast variety of studies employing iPSCs to treat retinal diseases. In a 2017 study conducted by Takahashi et al., autologous iPSCs were differentiated into retinal pigment epithelial (RPE) cells and were transplanted to the patient suffering from macular degeneration. IPSCs were generated

Figure 7.6 Different tissues derived from stem cells a T-tubules differentiated from hiPSCs. (Green: Wheat germ agglutinin, Red: Cardiac troponin T (cTnT), Blue: Nuclei; scale bar:10 μm). b Vocal fold epithelial mucosa derived from hiPSCs. (Green: Cytokeratin 14, Red: P63, Blue: Nuclei). c Kidney organoid derived from hiPSCs. (Blue: WT1 transcription factor, White: Tetragonolobus lectin, Green: Jagged Canonical Notch Ligand 1, Red: E-cadherin; scale bar: 200 mm). d Hepatic organoids with mature functional hepatocytes derived from hiPSCs. (Green: Albumin Red: Zonula occludens 1, blue: Nuclei; scale bar: 20 μm).

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using episomal plasmids. The resulting RPE cells were transplanted without using any scaffolds. No immunogenic or tumorigenic complications were observed after transplantation. However, the study was halted at some point due to the discovery of genetic mutations in the iPSC-derived cells of another patient. Still, the results obtained from one patient are promising in terms of safety and tolerability, as well as long-term cell survival [45]. The use of iPSCs is not limited to retinal diseases. In another example, human iPSCs were differentiated and mixed with fibroblasts in fibrin hydrogel strained between two flexible pillars to generate cardiac tissue. The fibrin hydrogel was stretched to recapitulate the mechanical forces in the cardiac tissue. The system was then exposed to electrical stimulation. The authors showed that the engineered cardiac tissue resembled mature cardiomyocytes with improved mitochondrial function [46,47]. In a study employing an engineering-based approach, a bioacoustic levitational (BAL) assembly technique was developed in order to be used for constructing 3D tissues. BAL offers contactless and rapid assembly of cell populations up to four million cells/mL. The functionality of the device was tested by creating cerebral cortex constructs using human stem cell
derived neuro-progenitor cells (NPCs). The cells were mixed with fibrin prepolymer that was designed to have mechanical properties comparable to brain tissue and stabilized the construct as it polymerized. The NPCs were in situ differentiated into neural cells using a differentiation medium. After 30 days, the functionality of the device was shown as the developed tissues formed 3D organization like that of the brain cortex [48]. Development of personalized treatment options is on the rise, with microfluidic systems holding a unique place in the field. A microfluidic chip culture system with continuous flow was utilized for culturing functional embryoid bodies (EBs). The mouse steroidogenic stem cell
derived EBs were cultured and differentiated into cells expressing ovarian antigens in the system, and the measured estradiol and progesterone levels were found to be physiologically relevant. The system allowed long-term culture of the EBs, as well as suitability for storing by cryopreserving. The system holds promise for stem cells to be used in microfluidic systems for generating personalized, on-demand therapies, and drug screening platforms [49]. Adult stem cells often come to play in tissue engineering studies. In an outstanding study, Hirsch et al. were able to restore 80% of the skin of a patient suffering from a severe genetic disease, using fibrin-cultured grafts containing genetically modified autologous keratinocytes and epidermal stem cells [50]. As an example utilizing the paracrine effects of stem cells, Swartzlander et al. used BMSC to overcome foreign body response occurring against scaffolds used for tissue engineering. They encapsulated murine BMSCs with poly(ethylene glycol) hydrogels. In vitro, the BMSCs weakened the macrophages by suppressing gene expression and protein secretion. In vivo, BMSCs or osteogenically differentiating MSCs were implanted subcutaneously to mice and again reduction in macrophage activation was observed. The reduction correlated inversely with osteogenic differentiation, meaning that immunomodulatory properties of MSCs diminish

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along with differentiation. This study shows that employing MSCs can regulate the immune response and is of great importance for use with scaffolds to enhance the tissue engineering outcome [51]. The potential clinical use of SVF cells in bone fractures was investigated in a study conducted by Saxer et al. in 2016. In this study, human SVF cells were implanted in nude rats having femoral fractures, by seeding the cells within fibrin gel onto ceramic granules. SVF-treated rats exhibited ossicle formations arising from human cells. These cells were further implanted to patients with proximal humeral fractures. Bone ossicle formation was again observed, which appeared to have arisen from the implanted cells. It was shown that without in vitro differentiation and cell expansion, SVF can form bone tissue and vasculature when implanted in a fracture microenvironment [52]. Since the discovery of the effect of matrix elasticity on cells, this hallmark has been utilized to direct stem cells to a particular lineage [53]. In a 2015 study, micro- and nano-patterns were created in polyimide to compare in vitro differentiation of MSCs on groove/ridge structures. It is shown that 15 μm ridges support adipogenic differentiation and 2 μm ridges increase osteogenic differentiation, relevant to the cell morphology. The nano-patterns with 650 nm periodicity enhance both adipogenic and osteogenic differentiation. It was observed that the differentiation preferences were not merely due to the structures and were actually induced by the differentiation media. Still, the effect of surface topography on supporting differentiation to particular lineages is significant [54]. Tailoring materials to direct stem cells to a particular lineage is a rising trend [55]. Employed nanocomposite hydrogels that mimic the extracellular matrix to induce bone regeneration, without recruiting osteoinductive factors. The researchers used gelatin methacryloyl (GelMA) with nanosilicates (nSi) to encapsulate human BMSCs. They showed that GelMA supported cell migration and proliferation while nSi enhanced osteogenic differentiation. They also displayed the biocompatibility of the scaffold by in vivo studies, opening new doors for osteoinductive factor
free bone regeneration [55]. Using a method other than classic differentiation approaches employing growth factors or soluble signals, in a study conducted in 2016, MSCs were differentiated into cartilage directly by inhibiting angiogenesis. It is well known that the stem cell niche has significant importance for differentiation. Since the microenvironment of articular cartilage is avascular and hypoxic, it was hypothesized that inhibition of angiogenesis could mimic the specific microenvironment to support cartilage formation. A decoy-soluble vascular endothelial growth factor (VEGF) receptor-2 (sFlk1) was transduced to human MSCs. The transduced cells were seeded on collagen sponges to be implanted in nude mice ectopically. These cells lacked vascular ingrowth and differentiated into a stable hyaline cartilage. It was shown that the differentiation was in fact due to the hypoxic environment, and MSC differentiation was possible without additional factors or signals. This phenomenon can be further utilized to improve cartilage repair using MSCs by blocking VEGF [56].

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The lack of vascularity in engineered 3D tissues is a major cause of cell death due to reduced diffusion rates of oxygen and nutrients as the thickness of the construct increases. This is one of the main burdens of osteogenic constructs. In a study, endothelial and mesenchymal progenitors obtained from the SVF of human adipose tissue were investigated in terms of supporting vascularity in vitro. Hydroxyapatite scaffolds were seeded with SVF using a perfusion bioreactor. After 5 days, capillary networks were observed in vitro and the networks anastomosed with the host vascular system in ectopic nude rat implantation. The results suggested that SVF supported vasculature better than BMSCs and ASCs. Overall, it can be inferred that SVF cells enhance engraftment of osteogenic grafts and improve bone tissue generation [57]. Another approach to improve vasculature of osteogenic grafts depends on the VEGF. In the study conducted by Helmrich et al., BMSCs were genetically edited to express rat VEGF. BMSCs were seeded on osteoconductive silicate
substituted apatite granules. In the ectopic nude rat model, increased vascular network was observed 8 weeks after implantation. An unexpected outcome, decrease in bone quantity occurred, still VEGF did not negatively affect BMSC engraftment in vivo. These findings suggest VEGF overexpression can be beneficial for vascularization of osteogenic constructs, but impairment in bone homeostasis should be overcome for clinical applications by finding an equilibrium between angiogenesis and osteogenesis [58]. Insufficient supply of oxygen and nutrients is also a major issue hindering transplantation of engineered skin grafts. One approach to engineer dermo-epidermal skin substitutes with vasculature is the use of SVF-derived endothelial cells. To maintain the endothelial cells in 2D culture, the cells were encapsulated in 3D fibrin or collagen type I hydrogels. The skin substitutes transplanted to immune-deficient rats gave rise to vascular networks that anastomosed to host vasculature by day 4. The newly formed epidermis maintained homeostasis along with negligible contraction in the dermis and was sustainably covered with epidermis. The perfusion of these SVF-based constructs paves new ways for treatment of skin loss [59].

7.5

Conclusion

In tissue engineering field, the search and use of suitable multipotent or pluripotent stem cells is a rising concept. Research on engineering tissues has primarily focused on providing functionality comparable to natural tissues. Major tasks are adjusting differentiation protocols to obtain specific lineages, purifying the resulting cells and combining with suitable scaffolds. After risks regarding carcinogenicity and infection are discarded, the formed tissues can then be implanted in vivo, to replace and/ or assist the function of the diseased site. Tissue engineering approaches to culture and implant differentiated lineages of stem cells show great potential and will have a significant contribution to regenerative therapies. Although great improvements have occurred in tissue engineering, advanced studies are required in order to develop fully functional organs/tissues.

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Immune cells: sources, properties, and cell types

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S. Jung1,2 and Florent Meyer1,3 1 Oral Biology Department, Faculty of Dentistry, University of Strasbourg, Strasbourg, France, 2INSERM UMR_S 1109 (Molecular ImmunoRheumatology unit), Strasbourg, France, 3INSERM, UMR_S1121, Biomaterials and Bioengineering, Strasbourg, France

Abbreviations Ab ADCC Ag APC BCR BM cDC CD CLP CNS CMP DAMP DC ECM ECP EDN EPO FasL Fc FO FDC GC HSC IFN Ig IL ILC IEL iNKT LN LTi MBP

antibody antibody-dependent cell-mediated cytotoxicity antigen antigen presenting cell B cell receptor bone marrow conventional dendritic cell cluster of differentiation common lymphoid progenitor central nervous system common myeloid progenitor damage-associated molecular pattern dendritic cell extracellular matrix eosinophil cationic protein eosinophil-derived neurotoxin eosinophil peroxidase Fas ligand fragment crystallizable region follicular B cell follicular DC germinal center hematopoietic stem cell interferon immunoglobulin interleukin innate lymphoid cell intra-epithelial lymphocyte invariant NK T cell lymph nodes lymphoid tissue inducer cell major basic protein

Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00009-X © 2020 Elsevier Ltd. All rights reserved.

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MoDC MHC Mϕ MZ NADPH NET NK NOD PAMP PBMC PC pDC PDGF PMN ROS TCR TD TF TFH TFR TGF-β Th TI Treg TLR VEGF WBC

8.1

monocyte derived DC major histocompatibility complex macrophage marginal zone B cell nicotinamide adenine dinucleotide phosphate neutrophil extracellular trap natural killer nucleotid-binding oligomerization domain pathogen-associated molecular pattern peripheral blood mononuclear cell plasma cell plasmacytoı¨d DC platelet derived growth factor polymorphonuclear cell reactive oxygen species T cell receptor T-cell dependent transcription factor follicular helper T cell follicular regulatory T cell transforming growth factor beta helper T cell T-cell independent regulatory T cell Toll-like receptor vascular endothelial growth factor white blood cell

Introduction

The use of any biomaterial will trigger the immune response regardless of its composition (synthetic or natural). Indeed, it results in the break of the body’s physical and chemical barriers with the production of damage-associated molecular patterns (DAMPs). DAMPs are prone to activate the innate immune system by the mean of specific receptors located on innate immune cells, and more largely on several cell types such as endothelial or epithelial cells. Innate immune system activation tends to help the wound healing. It is a complex mechanism that implies the activation of several different cell types; locally and temporally regulated. The knowledge of such reactions and of the physicochemical characteristics that drive the immune response is of paramount importance to design and evaluate a biomaterial. In this chapter, we describe the immune response to biomaterials and during wound healing in order to decipher the elements that are crucial to avoid the worst case scenario, that is, foreign body reaction. A description of the different immune cell types will help the reader to strengthen his knowledge on their different characteristics and their implications in the immune response. Finally, a specific focus will be

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put on the use of such cells in vitro to characterize biomaterials and more specifically on how to prepare such cells.

8.2

Immune system consideration in the use of biomaterials and tissue regeneration

8.2.1 Overall description of the immune system: innate versus adaptive system (“know your basics”) The immune system is the “organ” in charge of the overall protection of the organism. We can find such system in any Metazoa. In humans, it is classically described as a dichotomic system that is divided into the innate and the adaptive immune arms. Both systems rely on each other to control 99.9% of the threats. If the adaptive immune system is found in only 5% of the species, the innate immune system is found in every specie from the plants to insects. Both systems are complementary with the innate immune arm giving a rapid (few hours) but less specific response, and the adaptive system that takes 2
5 days for a proper installation, but that is characterized by a much higher specificity and efficiency. In bioengineering, the immune system is of paramount importance because tissue healing and remodeling rely on the activation of immune cells. The activation of the adaptive immune system depends on the prior activation of the innate immune system. Indeed, it cannot recognize the pathogen in its whole but only after antigen (Ag) processing by cells of the innate immune system. Moreover, costimulatory signals delivered by innate immune cells are mandatory to trigger the development of the adaptative immune response.

8.2.1.1 Just a matter of recognition The activation of the innate immune system is done by recognition of molecular patterns that allow to discriminate normal self from damaged self and nonself. Pathogen-associated molecular patterns represent all the molecular trademarks found in pathogens (viruses, bacteria, and fungi), such as sugar polymers, single strand RNA, non-methylated CpG-rich DNA. DAMPs represent components of the host’s cells but that are either not in their good location or in a degraded/denatured form. For example, ATP or uric acid can trigger innate immune response when they are found in the extracellular compartment. Low molecular weight hyaluronic acid triggers immune response as well because it represents a degraded form of the native molecule [1]. Those molecular patterns are recognized by innate immune receptors called pattern recognition receptors that are located at the cell surface (in the cytoplasmic membrane), or inside the cell, in the cytoplasm or in vacuoles (endocytic vesicles or phagosomes). Those receptors include, among others, Tolllike receptors (TLRs) and nucleotide-binding oligomerization domain (NOD)
like receptors [2]. Some receptors are secreted in the extracellular environment, especially in mucosa, such as ficolin in the lungs [3]. Receptors of the innate immune

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system share two specificities: first, they are genetically programmed and are therefore invariable, and second, they are expressed by a broad range of cells (all the cells that can encounter a pathogen). The innate immune system is assisted in pathogen/foreign material recognition by a secretory system called the complement system. This system is composed by proenzymes that are activated by catalysis. These enzymes lead to the destruction of the pathogen either by the creation of protein pores inside the membrane of infected cells and bacteria, or by facilitating the interaction with the immune cells. Indeed, once activated, this system will produce C3b protein that can interact with the surface of an infected cell or bacteria and with immune complexes. C3b receptor (complement receptor 1) is expressed on several immune cells, such as neutrophil granulocytes and macrophages, and plays an important role in processing of immune complexes, phagocytosis of C3b-bearing microorganisms, and regulation of the immune response. The complement system is worth citing because it is activated by three different ways: interaction with an antibody (Ab) bound to its Ag (immune complex; classical pathway), direct interaction with sugar decorating extracellular part of the pathogen membrane (lectin pathway), and by autocatalysis (alternative pathway). It is this third way of activation that is important in bioengineering. C3 protein undergoes spontaneous hydrolysis. The enzyme becomes activated and can interact with the B protein to form a C3 convertase that will catalyze the degradation of C3 protein (amplification of the system) [4,5]. As the C3 enzyme undergoes constant activation, it is important to regulate the system to avoid any damage. For this purpose, C3b is active only when it binds covalently to proteins at the surface of the pathogen. This reaction has to be quick; otherwise, the reaction takes place inside the C3 protein and inactivates its enzymatic capabilities. Even though this system is devoted to the destruction of pathogens, the reaction can occur at the surface of a biomaterial if it is composed of proteins or integrates NH2 moieties. Hence, it could lead to the inadvertent activation of an immune response directed against the material [6]. Besides, the adaptive immune response relies on B and T lymphocytes (B and T cells) that are heterogenous cell types characterized by a variety of receptors and modes of action. B lymphocytes are primarily responsible for the humoral immunity, and T lymphocytes are involved in cell-mediated immunity. On a very simplified basis, T cells can be divided into CD4 1 T cells that are responsible for the activation and guidance of the adaptive responses and CD8 1 T cells that mediate cytotoxic responses to viruses and tumor cells. As its name implies, the adaptive immune response adapts the reaction to the characteristics of the pathogen/lesion. In a simple and schematic way, an extracellular threat will be more processed via the humoral way with the production of Abs. Intracellular threats will be processed via the cytotoxic way aimed at killing all the cells that present signs of abnormality (tumor cells) or infection (viruses, intracellular bacteria). Unlike B lymphocytes, T cells do not recognize directly the Ags but after their processing by cells called antigen-presenting cells (APCs). Indeed, during the first encounter with an extracellular pathogen, APCs can internalize the pathogen by phagocytosis and digest it. Peptides derived from the pathogen are then presented on the cytoplasmic

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membrane by specific receptors called major histocompatibility complex (MHC) class II molecules to CD4 1 T cells. The recognition of the Ag presented by APCs (in particular dendritic cells; DCs) in the secondary lymphoid organs leads to the activation of naı¨ve T cells. As the specificity of recognition is highly increased in the adaptive immune system and considering the huge variety of Ags, the repertoire of Ag-receptors is therefore characterized by an extremely high diversity (total diversity up to 1018). Each cell bears a single type of receptor with a unique specificity of recognition. This diversity is generated by somatic gene segment rearrangements that take place during lymphocyte’s development. In B cells, additional modifications of the rearranged regions allow an increase of their affinity for the Ag and are beneficial to ensure a faster and greater response in the case of a second exposure to the same pathogen.

8.2.1.2 Just a matter of amplification and increased efficiency Innate immune response is initiated locally. The first reaction of recognition and destruction of the pathogen is driven by tissue-resident cells such as macrophages, and/or DCs. The reaction is then rapidly amplified by the development of an inflammation that is driven by the production of proinflammatory cytokines such as interleukin (IL)-1, IL-6, or TNF-α. One of the consequences of inflammation is the local recruitment of other innate immune cells to amplify the reaction. Those innate immune cells are granulocytes [neutrophils, eosinophils, and basophils; also named polymorphonuclear cells (PMNs)], monocytes/macrophages, DCs, and natural killer cells (NK cells). Cell recruitment is driven by cell homing and by the modification of local parameters (e.g., serum exudate) [7,8]. All innate immune cells are capable of pathogen destruction. Different strategies are used such as phagocytosis, excretion of various enzymes (i.e., proteases, lysozyme) and antimicrobial peptides, or production of reactive oxygen species (ROS). This reaction primes the adaptive response. Indeed, APCs can migrate from the lesion site to secondary lymphoid organs (lymph nodes and spleen) where lymphocytes will initiate the adaptive immune response. Moreover, as almost every lymphocyte bears a unique Ag receptor, an amplification is mandatory. After the binding of its specific Ag, the cell is activated to divide and to produce clones in a process known as clonal expansion, explaining why the adaptive reaction takes days rather than hours to develop. CD4 1 T cells, also called T helper cells (Th), orchestrate the immune responses. Depending on the Ag and the costimulatory factors produced by the APCs, activated CD4 1 T cells can differentiate into different T helper cells subsets (see Table 8.3). For example, Th2 cells activate B cells that differentiate either into plasma cells, specialized in the production of Abs, or into memory B cells that provide a rapid Ab response upon secondary exposure to the same Ag. Th1 cells are crucial for cell-mediated responses driven by macrophages and CD8 1 cytotoxic T cells. The latter are activated by the recognition of intracellular Ags presented by MHC class I molecules. In the final activation of the immune response, cells from the innate immune system such as neutrophil

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granulocytes, macrophages, or NK cells continue to perform their action that is even further amplified by lymphocyte cytokines‘ production. An important characteristic of the adaptive immune response is the establishment of a memory. Indeed, thanks to the receptor specificity, the response after a second encounter with the same pathogen will be both faster and greater in magnitude. If the notion of memory was for a long time specific to the adaptive immune system, it is now discussed for the innate immune response too.

8.2.2 Tissue regeneration/wound healing (“why immune system is so important”) Despite a constant improvement in the knowledge of tissue repair mechanisms, it still remains not fully understood why mammals have a tendency for scar formation and imperfect healing. During evolution, the increase in immune competence appears negatively correlated to the regenerative capacities and the priority seems to be given to rapid wound closure, in order to prevent the entry of pathogens, rather than to regeneration [9
11]. Indeed, wound repair results almost systematically in the formation of a scar, which is composed of acellular extracellular matrix (ECM) and is very different from the original tissue, with a loss of specialized functions [9]. The goal of regenerative medicine is to restore tissues back to a normal physiological and functional state [9]. Regenerative medicine is therefore intimately linked to tissue healing and to its “Holy Grail”, that is, perfect regeneration [12]. Despite this apparent inverse relationship between immune capacities and regeneration during evolution [11], it has now become evident that the immune system plays a crucial role in the regulation of healing process [13]. The type of the immune response and the cells involved significantly influence the tissue-healing process that can range from pathological repair (i.e., fibrosis) to regeneration (i.e., complete restoration) of a functional native tissue [9]. In the field of bioengineering and regenerative medicine, immunomodulation is becoming a very attractive approach [14]. Modulation of the host inflammatory response may improve biomaterial integration and support regenerative therapies but could also be an alternative to the use of stem cells and growth factors [9]. A good knowledge of the immune-mediated mechanisms of tissue healing is, therefore, an indispensable prerequisite for the design of regenerative strategies. Tissue healing is a tightly regulated and highly dynamic process that relies on complex interactions between immune cells, tissue resident cells, soluble mediators, and molecules from the ECM. It can be divided into three main stages: inflammatory phase, tissue formation, and tissue remodeling (Fig. 8.1). A tissue injury will induce the immediate onset of an immune response. Even in the absence of any pathogen, a sterile inflammation will be triggered by various danger signals (DAMPs or alarmins) released from cells [15] that are injured following the implantation of a biomaterial, for example. Resolution of the inflammatory response is however an integral part of the process and an essential prerequisite to allow

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Figure 8.1 The main actors of the immune response following tissue injury. (A) Kinetic of immune cells‘ mobilization after tissue injury. Tissue resident cells such as tissue resident macrophages sense tissue damage (DAMPs and alarmins) and trigger the mobilization of other immune cells. Neutrophil granulocytes are followed by monocytes/macrophages and T cells. The relative amount of each cell type recruited is not represented. (B) Overview of the initial inflammatory phase following tissue injury. Chronological events are represented from left to right. Tissue damage is sensed by tissue-resident macrophages via DAMPs. Neutrophil granulocytes are the first circulating immune cells recruited to the site of injury, promoting inflammation and monocyte/macrophage recruitment. The inflammation is initially maintained by “pro-inflammatory” M1 macrophages (also called “classical activated” macrophages), before being eventually resolved with the help of “antiinflammatory” M2 macrophages (also called “wound healing” or “tissue repair” macrophages). (C) Overview of the proregenerative immune mechanisms. A critical number of macrophages displaying an “anti-inflammatory” phenotype contribute to regeneration through a crosstalk with Tregs, which in turn help sustain the antiinflammatory phenotype via secretion of antiinflammatory cytokines such as IL-10 and TGF-β. Tregs may also enhance the regenerative capacity of endogenous stem/progenitor cells through the secretion of growth factors. Th2 cells induce/maintain antiinflammatory macrophages. γδ T cells are able to recruit innate immune cells and to directly trigger tissue growth. Black arrows (Continued)

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L

complete tissue healing. An excessive or prolonged inflammatory response will therefore be associated with defective healing [16]. The initial phase of the inflammatory reaction involves the innate immune system, and the neutrophil granulocytes (or polymorphonuclear neutrophils; Table 8.1) are usually the first immune cells that are recruited (first line of defense) [17]. Neutrophil granulocytes are major pathogen-fighting cells that were initially considered as mainly proinflammatory agents. They phagocytose pathogens and kill them through various cytotoxic mechanisms, including the release of antimicrobial substances (i.e., ROS and antimicrobial peptides), the secretion of various cytokines, and the deployment of neutrophil extracellular traps (NETs) [18]. However, they also exhibit antiinflammatory functions and play an important role in tissue healing. Despite their short life span, they remain in the inflammatory site up to 3 days [17]. They orchestrate the recruitment of monocytes and macrophages that remove dying neutrophil granulocytes as well as cellular debris via phagocytosis. Neutrophil granulocytes promote, therefore, their own removal and contribute to the resolution of the inflammatory response [19,20]. Mast cells (Table 8.1) are also early effectors of the innate immune response. Although they are well-known for their proinflammatory properties (degranulation), there is increasing evidence that they can also have potent antiinflammatory effects and regulate healing process [20,21]. Besides their role as scavenger cells that phagocytose cellular debris and apoptotic cells, macrophages (Table 8.1) are also key players in the regulation of tissue healing. After tissue injury the population of resident macrophages is enriched by an important recruitment of migratory blood monocytes. The latter arrives in the inflammatory site about 1
3 days after neutrophil granulocytes and differentiate into macrophages or, albeit to a lesser extent, into monocytes derived DCs. Macrophages have been classically divided into two main categories, namely, “classically activated” M1 macrophages and “alternatively activated” M2 macrophages [22]. M1 macrophages have a proinflammatory phenotype and play a crucial role in the elimination of pathogens through the production of proinflammatory cytokines (e.g., IL-6, IL-1β, and TNF-α), a high phagocytic activity, and their Ag presentation capacity with a subsequent activation of Th1 responses. At the end of the acute phase, the presence of IL-4 and IL-13 leads to a phenotypic switch from “proinflammatory” M1 macrophages to “antiinflammatory” and “wound-healing” M2 macrophages. The latter participates in the successful outcome of the healing response through the secretion of antiinflammatory cytokines [e.g., IL-10, transforming growth factor (TGF)-β], indicate a differentiation path or secretion of immune modulators/morphogens. Black dashed arrows indicate a hypothetical differentiation path. Red arrows indicate induction. Blue arrows indicate inhibition. DAMP, damage-associated molecular pattern; IL, interleukin; TGF, transforming growth factor. Source: Adapted from Julier Z, Park AJ, Briquez PS, Martino MM. Promoting tissue regeneration by modulating the immune system. Acta Biomater 2017;53:13
28. Available from: https://doi.org/10.1016/j.actbio.2017.01.056 [9].

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the promotion of cell proliferation, ECM deposition (fibrosis), and tissue remodeling [23]. However, M1/M2 dichotomic classification is oversimplistic, and macrophage activation spectrum is now considered to be much broader [24,25]. Indeed, these cells are able to respond dynamically and to adapt their functional state to the different mediators present in the microenvironment [24]. Depending on the signals that they receive, macrophages can therefore exacerbate inflammation, promote tissue repair or even drive regeneration. It has been shown quite recently that this cell type is crucial for the regeneration of different tissues [26,27]. Thus, macrophages appear to be promising targets for the design of future regenerative therapies [9,20]. The role of DCs (Table 8.1) during tissue repair is not completely understood. Considered as professional phagocytes as well as professional APCs that link innate to adaptive immunity, DCs may also have immunoregulatory functions during wound healing through the control of macrophage homeostasis [9]. T lymphocytes (Table 8.3), in particular regulatory and γδ T cells subsets, play a crucial role in tissue healing and regeneration [9]. Regulatory T cells (Tregs) create an appropriate environment for tissue repair. They secrete different immunosuppressive cytokines such as IL-10 and TGF-β that dampen the immune responses [28], but they also promote M2 macrophage phenotypic switch [29]. γδ T cells are able to recruit innate immune cells and to directly trigger tissue growth [9]. The immune system directs the quality and the outcome of the tissue repair in both positive and negative fashions [9]. Therefore, the modulation of the inflammatory response appears to be a very promising approach in the field of regenerative medicine.

8.2.3 Immune response to biomaterials: when all goes wrong, that is, the foreign body reaction The foreign body reaction is described as the attempt of the immune system to isolate the materials from the body by the formation of a thick and dense surrounding fibrous capsule [30
32]. From a biomaterial point of view, it is a complete failure as it is the integration of the material that is expected. This reaction is seen preferentially with synthetic biomaterials rather than with biological-derived materials. Such difference is related to the degradability of the material and the ability of the immune system to process it. Classically, the foreign body reaction is divided into five different stages (Fig. 8.2). The first stage is the coating of the material by a protein layer. Such reaction is the consequence of the surgical lesion that is associated with the introduction of the biomaterial. The first body fluid encountered is the blood. Plasma proteins will therefore coat the biomaterial. Classically, two layers that are linked to two different mechanisms are described. The first thin layer (2
5 nm), directly in contact with the material, is composed of plasma proteins and its composition is directly influenced by plasma composition. Albumin is the first protein that

Platelet

Neutrophil granulocyte Mastcell

Histamine IL4– Il13

Macrophage

IL8 MCP-1 MIP-1β

DAMPS

Il4/IL13 M1/M2?

PDGF, VEGF TGF-β β

Lymphocyte

Giant multinucleated cell Fibroblast

Proteasis RGD Biomaterial

Figure 8.2 Foreign body reaction. Description of the successive phases of foreign body reaction once a material is implanted: - Formation of a protein coating (RGD sequence (arginine-glycine-aspartic acid) containing proteins) around the material with the development of fibrin clot containing platelets and helping cell interaction with the material. The release of chemotactic molecules (C5a or C3a), by direct activation by proteins that are coated on the material surface, along with the production of DAMPs following the surgical injury, helps the development of an acute inflammatory response orchestrated by neutrophil granulocytes and mast cells. - Arrival of first macrophages. - Development of a chronic inflammation. Macrophage fate (M1/M2) is driven by locally infiltrated lymphocytes. Macrophages fuse to form giant multinucleated cells because of frustrated phagocytosis. - Fibrosis and capsule formation following fibroblasts invasion. Progressive thickening of the capsule with reduced number of cells.

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adheres at the surface of the biomaterial, followed by fibrinogen, high molecular weight kininogen, fibronectin, and vitronectin. Mechanism and kinetic of plasma proteins adhesion and desorption from the surface was nicely studied and described by Vroman [33]. The second, thick, layer (hundreds of microns) is formed principally by fibrin, and its composition is close to a blood clot. It has been proven that factor XIII can be activated when it interacts with cationic surfaces [34]. It subsequently activates thrombin, leading to fibrin clot formation. Even if fibrinogen is not transformed, the adhesion of fibrinogen on a surface changes its conformation and induces the unmasking of some platelet’s adhesion sites. Such coating is important for the immune response, first because it provides adhesion sites for cells, and second because it traps soluble factors such as complement and immunoglobulins (Igs). In this protein coating, we can find C3 that is activated through the alternative pathway (autohydrolysis). C3b is formed at the surface of the material [6]. In the meantime, C3a that is a strong chemoattractive signal for immune cells is released in the surrounding area. Aside, some Igs are trapped in the protein coating and are able to trigger the classical complement pathway. C3b represents a ligand for some integrins that are present at the surface of leukocytes and can therefore help their adhesion to the surface. This step triggers acute inflammation. The second step is characterized by local infiltration of neutrophil granulocytes and mast cells. Three different messages act concomitantly: the complement activation, the DAMPs associated with the surgical lesion, and somehow the direct interaction with the material itself. Some other factors such as von Willebrand factor, P-selectin, platelet factor 4, or factors VII and XI trigger neutrophil granulocytes infiltration. Neutrophil granulocytes and mast cells degranulate and release the content of their intracytoplasmic granules (e.g., myeloperoxidase, ROS), leading to the degradation of the surrounding environment and helping in the communication with immune cells (cytokines). For example, neutrophil granulocytes release IL-8, monocyte chemoattractant protein 1 (MCP-1) and macrophage inflammatory protein 1 beta (MIP-1β) and mast cells release IL-4, IL-13 and histamine. Moreover, various growth factors present in the material protein coating and/or ECM are released under the action of proteases excreted by immune cells [e.g., platelet derived growth factor (PDGF), TGF-β]. Such reaction is intended to degrade the protein coating and the material in order to trigger healing. Cytokines such as IL-8 are strong chemoattractants for macrophages. Depending on the material composition and the degradation rate, the inflammatory reaction can become chronic with the arrival of macrophages that represent the third step [35,36]. In most cases, this step lasts for 2
5 weeks and is characterized by an infiltration of monocytes/macrophages as well as lymphocytes. Macrophages bear on their surface scavenger receptors and TLRs that can recognize directly the material surface. Adhesion leads to macrophage activation and differentiation. Several populations of macrophages have been described, and their phenotype depends, among others, on the type of cytokines that they produce. As described above, M1 macrophages produce proinflammatory cytokines (e.g., IL-1β, IL-6, IL-8, and TNF-α) that trigger a strong

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inflammation with degradation of the surrounding tissue and a potential degradation of the foreign material. M2 macrophages produce antiinflammatory cytokines (e.g., IL-10 and TGF-β) that rather induce tissue remodeling via the action of metalloproteases. It is noteworthy that macrophages can switch from one phenotype to the other depending on the extracellular messages. For example, IL-4 and IL-13 favor M2 phenotype. Lymphocytes, in particular CD4 1 T cells, can be found in the vicinity of the foreign material. They secrete IL-4 and IL-13 and drive the reaction toward inflammation resolution with an increase of M2 macrophages. The switch M1/M2 leads in that case to the fusion of several macrophages that form multinucleated giant cells and sign the fourth step [32]. It could increase the phagocytosis and inhibit their anoı¨kis or apoptosis. Multinucleated giant cells are characteristic of a foreign body reaction. The macroscopic signal for this transformation is the persistence of the material and a frustrated phagocytosis. Such cell type is meant to degrade the material. The components that they are secreting vary from proinflammatory to tissue remodeling mediators. This diversity signs that they arise from a heterogenous population of macrophages that fused together. In the long term, the presence of these cells is deleterious because they produce ROS, enzymes and acids. It can help the degradation of the material, but in the meantime, it also damages the surrounding tissue. This local tissue remodeling is associated with the release of profibrotic and angiogenic factors such as PDGF, vascular endothelial growth factor, and TGF-β that are produced by M2 macrophages, but also by endothelial cells, keratinocytes, fibroblasts, thrombocytes, or adipocytes [37]. This final step sees the arrival of fibroblasts at the surface of the material. Those cells lay down ECM components. Over time, there is a maturation of the tissue from a highly cellular to a more fibrous (denser and less cells) tissue with a maturation of the collagen type from type III to I. TGF-β can trigger the differentiation of fibroblasts into contractile myofibroblasts. This final step should normally end with the apoptosis of myofibroblasts and fibroblasts and with the reduction of the collagen scaffold under the action of macrophages (special population). However, in the case of material persistence, this phenomenon can continue and lead to the formation of a thick fibrous capsule without tissue remodeling.

8.3

Immune cell description

8.3.1 Myeloid cells Table 8.1 Myeloid cells derive from the common myeloid progenitor. Cell types

Granulocytes or polymorphonuclear cells (PMN)

Neutrophils or neutrophil granulocytes

Localization

Main features

Key functions

Reviewed in

Blood (patrolling) Migration toward the sites of inflammation

Short lifespan (,24 h) Most abundant WBC type in human blood (50%
70% WBC) Production of B1011 neutrophil granulocytes per day in the BM Hallmark of acute inflammation: first cell type that is recruited to inflammatory sites

Effector cells of innate immunity Pathogen destruction: immunity to extracellular pathogens (bacteria and fungi) Phagocytosis: “professional” phagocytes ROS production via NADPH oxidase Degranulation: antimicrobial peptides, and proinflammatory cytokines, proteases, etc. NETs generation (NETosis) Immune regulation Polarization toward distinct phenotypes (pro/antiinflammatory) in response to environmental signals Cellular cross talk: maturation of DCs, etc. Inflammation resolution: wound healing, revascularization, etc. Immunity to parasites (helminths) Role in allergic diseases and anaphylaxis Induction of Th2 immune responses via the secretion of pro-Th2 cytokines (IL411) Low phagocytic activity

[38]

G

G

G

G

G

Basophils

Blood

Short life span (2
3 days) Least common granulocytes type in human blood (,1% WBC) Production of proinflammatory and vasoactive mediators Degranulation: preformed mediators (histamine, heparin, proteases, etc.) Synthesis of cytokines, lipid mediators, etc. G

[39]

G

(Continued)

Table 8.1 (Continued) Cell types

Eosinophils

Mast cells (mastocytes)

Localization

Main features

Key functions

Reviewed in

Blood (brief time B18 h) Migration to thymus and gastrointestinal tract where they reside under homeostatic condition (2
18 days) Accumulation in tissue inflammatory foci (prolonged survival)

1%
3% WBC IL5: key cytokine for eosinophils differentiation and survival. Production of several mediators contained in specific granules: cationic proteins (MBP, EPO, ECP, and EDN), cytokines, enzymes, etc. lipid bodies: lipid mediators

Low phagocytic activity Immunity to parasites (helminths), some viruses and bacteria Role in allergic diseases in particular asthma Modulation of the functions of other leukocytes: T cell activation (antigen presentation) Regulation of T cell polarization Mast cells activation through cationic proteins released from the granules DC maturation

[40,41]

Tissues: mucosal type (immature), connective tissue type (more mature)

Life span: weeks (mucosa) to months (connective tissue) Production of proinflammatory and vasoactive mediators Degranulation: preformed mediators (histamine, serotonin, and proteases) Synthesis of cytokines, lipid mediators, etc.

Share many functions with basophils Immunity to parasites (helminths and protozoa) and bacteria Role in allergic diseases and anaphylaxis Regulation of T cell responses according to the released cytokines Immune surveillance in tissues Role in wound healing.

[39,42]

G

G

G

G

G

G

G

Monocytes and macrophages (Mϕ)

“Professional” phagocytes Monocytes

Patrolling monocytes: CD16 1 , found in the circulating blood, 2%
6% WBC Migratory monocytes: CD14 1 Recruited in tissues and secondary lymphoid organs in inflammatory conditions!transendocytosis, migration, and differentiation into Mϕ In homeostatic conditions, resident Mϕ: Differentiation from yolk sac precursors and maintenance during adult life independently of monocytes (local proliferation)!long-lived cells Tissular localization: red pulp Mϕ (spleen), alveolar Mϕ (lungs), Langerhans cellsa (epithelia), microglia (CNS), Kupffer cells (liver), peritoneal Mϕ, etc. In early stages of inflammation, mainly “classical activated” M1 Mϕ: Differentiate from migratory monocytes and resident Mϕ Proinflammatory phenotype: metabolization arginine to nitric oxide (cytotoxic activity, inhibition of cell proliferation) At the end of the inflammatory response, phenotype shift from M1 to M2 Mϕ Antiinflammatory phenotype, metabolization of arginine to the repair molecule ornithine

Immune surveillance and maintenance of endothelium integrity Phagocytosis and antigen presentation

[43,44]

Maintenance of tissue homeostasis (scavenger cells): elimination of dead cells and debris without activation of an immune response (“noninflammatory” phagocytosis) Secretion of proinflammatory cytokines (IL-6, IL-1β, TNF-α, etc.), phagocytosis, antigen presentation, activation of Th1 response (against intracellular pathogens) Secretion of antiinflammatory cytokines (IL-10, TGF-β), promotion of cell proliferation, ECM deposition, tissue repair, and remodeling

[23,45,46]

G

Mϕ

G

G

G G

G

(Continued)

Table 8.1 (Continued) Cell types

DCs

Localization

Main features

Key functions

Heterogeneous group of “professional” APCs, phagocytic cells Note: Although they have a similar morphology, FDCs are not DCs. They are not derived from the BM HSC but are of mesenchymal origin. Located in the follicle of secondary lymphoid organs, they capture and present Ag and immune complexes to B cells cDCs

Basal expression of MHC class II molecules Immature cDCs In the blood, at body surfaces (skin and mucosa) Upon Ag capture, DCs undergo a process of “maturation” Mature cDC Migration to lymphoid organs G

G

pDCs

Circulating cells

Monocyte derived DCs (MoDCs)

Low numbers in homeostatic conditions: expand in inflammatory conditions Inflammatory DCs: derive from CD14 1 inflammatory monocytes

Reviewed in [47]

Immune surveillance: sensing and sampling the environment for self and nonself Ags, low T cell activation potential, high endocytic capacity, production of “innate” cytokine (i.e., IFNα) Enhanced Ag processing, induction of MHC molecules, costimulatory molecules and cytokine production Presentation of processed Ags to naı¨ve T cells: T cell activation Polarization of the response: immune responses (Th1-2-17 responses) or tolerance (Tregs) Production of cytokines: activation of B and NK cells Robust production of type I IFN: innate immune response against viruses Ag presentation to naı¨ve T cells: poor capacity compared to cDCs Initiation of adaptive responses: Ag presentation to effector T cells, pathogen clearance, cytokine production, etc.

Ag, Antigen; APCs, antigen-presenting cells; BM, bone marrow; cDCs, conventional DCs; CMP, common myeloid progenitor; CNS, central nervous system; DCs, dendritic cells; ECM, extracellular matrix; ECP, eosinophil cationic protein; EDN, eosinophil-derived neurotoxin; EPO, eosinophil peroxidase; FDCs, follicular DCs; HSC, hematopoietic stem cell; IFN, interferon; MBP, major basic protein; MHC, major histocompatibility complex; Mϕ, macrophage; MoDCs, monocytes-derived dendritic cells; NADPH, nicotinamide adenine dinucleotide phosphate; NET, neutrophil extracellular trap; NK, natural killer; pDCs, plasmacytoı¨d DCs; Treg, regulatory T cell; WBC, white blood cell. a Originally categorized as a DCs, Langerhans cells have been re-classified with macrophages on the basis of ontogeny.

8.3.2 Innate-like lymphocytes Table 8.2 Innate-like lymphocytes. Cell type ILCs

Localization

Main features

Key functions

[48
50]

Early effector immune cells: rapid response Mirror the phenotypes and functions of T cells: But do not express specific Ag receptors or undergo clonal selection and expansion once activated Type 1 ILCs

NK cells

4%
29% of peripheral blood lymphocytes (median 11.4%)a Also found in peritoneal cavity, liver, lungs, lymphoid organs, and uterus during gestation

Innate lymphoid counterpart of cytotoxic CD8 1 T cells Master TFs: T-bet and Eomes Activity mediated by the balance between signals from activating and inhibitory cell surface receptors Cells expressing normal levels of MHC class I molecules!strong inhibitory signal Downregulation of MHC class I molecules by viral infected or transformed cells (“missing self”)!reduced inhibitory signal Recognition of molecules expressed by viral infected and transformed cells!strong activation signal Two main subsets: CD56dim CD161 (90% of blood NK cells, high cytotoxic activity) and CD56bright (10% of blood NK cells, cytokine production) G

G

Reviewed in

Immunity to viruses, tumor surveillance Important role in pregnancy and graft rejection Direct cytolysis of infected or transformed cells: release of cytoxic granules containing soluble mediators (i.e., perforin and granzymes) Production of Th1 cytokines (IFN γ, TNFα): antitumor and antiviral effects Control of the switch from innate to adaptive immunity: interaction with Mϕ, DCs, iNKT, etc.

G

(Continued)

Table 8.2 (Continued) Cell type

Localization ILC1

Type 2 ILCs: ILC2

Type 3 ILCs: ILC3

Close to the epithelial barriers: interface between the environment and the associated lymphoid tissues ILC1, ILC2, and ILC3: 3 subgroups reflecting the cytokine expression profiles of CD41 Th cell subsets Activation by tissue signals (i.e., stress, and danger signals) and local environment modifications rather than by Ag

Main features

Key functions

Innate lymphoid counterpart of CD4 1 Th1 cells Master TF: T-bet

Immunity to intracellular pathogens, viruses, and tumors Role in chronic inflammation Production of Th1 cytokines (IFNγ and TNFα): Mϕ activation, ROS

Innate lymphoid counterpart of CD4 1 Th2 cells Master TFs: GATA3/RORα

Immunity to extracellular parasites (helminths) Involved in metabolic homeostasis Role in the pathogenesis of allergy and asthma Production of Th2 cytokines (IL4, IL5, IL13): alternative Mϕ activation, mucus production, eosinophilia, ECM deposition and tissue repair, vasodilation, and thermoregulation

Innate lymphoid counterpart of CD4 1 Th17 cells Master TF: RORγt Note: debate whether lymphoid tissue inducer cells (LTi) that are essential for the development of lymphoid organs and tissues belong to ILC3 subset

Immunity to extracellular pathogens (bacteria and fungi) Role in the pathogenesis of chronic inflammation Production of Th17 cytokines (IL-17 and IL-22): recruitment of neutrophil granulocytes, phagocytosis, antimicrobial peptides, and epithelium homeostasis

Reviewed in

γδ T cells

Rare subset in the blood and secondary lymphoid organs: 1%
16% of peripheral blood lymphocytes Tissue-specific localization of oligoclonal subpopulations sharing the same TCR chains Enriched in epithelial tissues (epidermis, gastrointestinal and reproductive tracks): .10% Part of a larger group of epithelial residing lymphocytes called IELs

“Unconventional” T cell population: much less frequent than αβ T cells First T cells to develop in the thymus Expression of heterodimeric TCRs composed of γ and δ chains Not restricted by MHC: activation by direct recognition of stress Ags in an MHC independent manner

First line of defense in tissues: broad functional plasticity Direct cytolysis of infected or transformed cells: release of soluble mediators (i.e., perforin and granzymes) Ag presentation: can serve as professional APCs Production of various cytokines Activation of other immune cells

[51,52]

Rare subset in the blood: 0.01%
1% of PBMCs Very long-term resident cells in lymphoid (thymus, spleen, and BM) and nonlymphoid tissues (liver11, fat tissue, etc.)

Thymic origin Share characteristics of both NK and T cells Expression of a limited repertoire of TCRs (invariant α chain) Recognition of a restricted number of Ags (glycolipids in particular α-galactosylceramide) that are presented by the nonpolymorphic MHC class I
like molecule CD1d expressed by APCs Expression of T cells and NK cells surface markers (CD16, CD56, etc.) Rapid activation by danger signals and proinflammatory cytokines

Immunoregulatory cells Different iNKT subsets that mirror CD41 Th cell subsets (i.e., cytokine production) Rapid modulation and activation of other immune cells (e.g., DCs, Mϕ, T, B, NK cells, etc.)

[53,54]

G

G

iNKT

Ag, Antigen; APCs, antigen-presenting cells; BM, bone marrow; CLP, common lymphoid progenitor; DCs, dendritic cells; ECM, extracellular matrix; IEL, intra-epithelial lymphocyte; IFN, interferon; IL, interleukin; ILCs, innate lymphoid cells; iNKT, invariant NK T cells; MHC, major histocompatibility complex; NK, natural killer; PBMC, peripheral blood mononuclear cell; ROS, reactive oxygen species; TCR, T cell receptor; TF, transcription factor. a Reference ranges and median values from [55]. Innate-like lymphocytes derive from the common lymphoid progenitor (CLP).

8.3.3 Lymphocytes Table 8.3 Cell type

T lymphocytes or T cells

Main features

Key functions

Suggested reviews

51%
84% of peripheral blood lymphocytes (median 72.5%)a Maturation in the thymus Express a TCR on the cell surface: recognition of Ag as peptides bound to MHC molecules presented by APCs Cellular immunity component of the adaptive immunity CD4 1 T cells: helper T cells (Th cells)

24%
54% of peripheral blood lymphocytes (median 39.5%)a Expression of CD4 coreceptor Central orchestrators of immune responses Recognition by the TCR of peptides presented by MHC class II molecules expressed by APCs!differentiation of activated CD4 1 T cells into effector Th cells guided by specific costimulatory signals and cytokines Multiple Th cells subsets described to date, including Th1, Th2, Th17, Th9, Th22, regulatory T cells (Tregs), follicular Th (TFH), and follicular Tregs (TFR) Due to space considerations, Th9 and Th22 subsets will not be presented here Th1 cells

Differentiation driven by IL12 Key TF: T-bet

Cell-mediated immune responses Immunity to intracellular pathogens (bacteria, viruses) Production of cytokines: IFNγ (signature cytokine), TNFα, etc. Activation of Mϕ (phagocytosis11), support of CD8 1 T cell functions Role in the pathogenesis of autoimmune diseases

Th2 cells

Differentiation driven by IL4 Key TF: GATA3

Humoral immune responses Immunity to extracellular pathogens (helminths) Production of cytokines: IL4, IL5, IL13, etc. Provide help to B cells (isotype switching, production of IgG, IgE, etc.), regulation of eosinophils (IL5), basophils, and mast cells functions, promotion of goblet cells hyperplasia Role in the pathogenesis of allergy and asthma

Th17 cells

Described more recently (2005)

Immunity to extracellular pathogens (bacteria and fungi)

[56
58]

[59]

Differentiation driven by IL6, IL21, and IL23, TGFβ Key TF: RORγt Proinflammatory Th cell subset Enriched in epithelial barrier sites (gastro-intestinal track, skin, lungs, etc.) Exhibit a certain instability and functional plasticity: can produce cytokines specific from other Th cell subsets

Production of cytokines: IL17A (signature cytokine), IL17F, IL22, IL21, etc. Detected at sites of inflammation early during the immune response: orchestration of the innate immune responses, additional neutrophil granulocytes recruitment and activation Maintenance of mucosal homeostasis Role in pathogenesis of autoimmune and auto-inflammatory diseases

Follicular helper T cells (TFH)

Reside in the periphery within Bcell follicles of secondary lymphoid organs (LN, spleen, and Peyer’s Patches) Key TF: BCL6 Expression of CXCR5 and PD1

Direct humoral immune responses: provide costimulation and stimulatory cytokines to B cells!trigger formation and maintenance of GCs, affinity maturation and differentiation of GC B cells to Ab-producing PCs and memory B cells Production of cytokines IL21, IL10

[60]

Regulatory T cells (Tregs)

Key TF: FoxP3

Negative regulation of immune responses and maintenance of immune homeostasis Suppressive functions: downregulation of the induction and proliferation of effector T cells Maintenance of the tolerance to food and self-Ags

[61]

G

G

Production of antiinflammatory/inhibitory cytokines: IL10, TGFβ, and IL35 Induction of effector cells apoptosis (granzyme) Prevention of T cell costimulation: constitutive expression of the inhibitory receptor CTLA4 that outcompetes CD28 for its ligands CD80 and CD86 Different subsets: Natural (nTregs) or thymusderived (tTregs) Tregs: generated in the thymus

G

Consumption of IL2 that is important for Th cells differentiation

(Continued)

Table 8.3 (Continued) Cell type

Main features

G

Follicular regulatory T cells (TFR)

CD8 1 T cells: cytotoxic T cells

Key functions

Suggested reviews

Control of the magnitude of normal GC responses to avoid the production of high-affinity auto-Abs

[62]

Immunity to intracellular pathogens (viruses and bacteria), tumor surveillance Direct cytolysis of infected or transformed cells: release of cytoxic granules containing soluble mediators (i.e., perforin and granzymes) Production of Th1 cytokines (IFN γ, TNFα): antitumor and antiviral effects Fas/FasL interactions

[63]

Peripheral Tregs (pTregs) and induced T regs (iTregs): TGFβ has the capacity to induce the differentiation of Tregs from mature CD4 1 T cells in vivo outside of the thymus at peripheral sites (pTregs) or in vitro (iTregs)

Reside in the GCs Expression of CXCR5 and PD1 such as TFH cells but function such as Tregs Express the TF FoxP3

13%
40% of peripheral blood lymphocytes (median 26.5%)a Expression of CD8 coreceptor Recognition by the TCR of peptides presented by MHC class I molecules expressed by all nucleated cells

G

G

G

Expression of FasL by activated CD8 1 T cells!Fas binding on the surface of target cells!activation of caspases cascade!apoptosis of target cells A CD8 1 T cell can express both Fas and FasL and kill another CD8 1 T cell (fratricide)!elimination of effector immune cells during the contraction phase (end of the immune response)

B lymphocytes or B cells

6%
21% of peripheral blood lymphocytes (median 11.5%)a Maturation in the BM Express BCR on the cell surface: recognition of Ag in its native form Humoral immunity component of the adaptive immune response!secretion of Abs (also called Igs) but also Ag presentation (“professional” APC) and cytokine secretion Migration of immature B cells from the BM into the spleen as transitional B cells (T1/T2 B cells)

Immediate protection against pathogens IgM: Ag agglutination, complement activation, etc.

Differentiation into FO B cells or MZ B cells depending on signals received through the BCR 5 mature naı¨ve B cells Activation in the secondary lymphoid organs after Ag recognition
Extrafollicular response (TI activation and early phase of the response after TD activation):

Persistent protection against pathogens IgG: neutralization, complement activation, opsonization (triggers phagocytosis), ADCC, etc. IgA: neutralization, mucosal immunity, etc. IgE: binding to Fc receptors on the surface of basophils and mast cells!induces degranulation

outside lymphoid follicles

Rapid Ab responses upon secondary exposure to the same Ag

G

G

G G

!generation of short-lived PCs that produce unmutated, low-affinity Abs from IgM isotype GC reaction (late phase of the response after TD activation): specialized structures within the lymphoid follicles where activated B cells undergo extensive proliferation, Ig class switching, and affinity maturation (somatic hypermutation)

G

!generation of long-lived PCs that produce somatically mutated, high-affinity Abs from switched isotypes (IgG, IgA, and IgE) Some of the Ag activated B cells develop into memory B cells Regulatory B cells (Bregs)

Low proportion of B cells (,10% of circulating B cells in healthy individuals)

Negative regulation of immune responses and maintenance of immune homeostasis Production of the antiinflammatory/inhibitory cytokines: IL1011, (TGFβ) Promotion of Treg differentiation Suppression of the differentiation of proinflammatory lymphocytes

[64
66]

Ag, Antigen; APCs, antigen-presenting cells; BCR, B cell receptor; CLP, common lymphoid progenitor; GC, germinal center; IFN, interferon; IL, interleukin; iTreg, induced regulatory T cell; LN, lymph nodes; MHC, major histocompatibility complex; nTreg, natural regulatory T cell; pTreg, peripheral regulatory T cell; TCR, T cell receptor; TF, transcription factor; TD, T-cell dependent ; TFH, follicular helper T cell; TFR, follicular regulatory T cell; Th, helper T cell; TI, T-cell independent; Treg, regulatory T cell; tTreg, thymus-derived regulatory T cell. a Reference ranges and median values of the different lymphocytes subsets from [55]. Lymphocytes derive from the CLP.

220

8.4

Biomaterials for Organ and Tissue Regeneration

Immune cell sourcing

In order to test biomaterials used in bioengineering strategies, it is helpful to perform a first screening of the potential immune response that is induced by those biomaterials. It is more interesting to test the immune response in vivo because it represents the overall mechanism with all the players. However, such experiments are time-consuming and a first in vitro assessment can be performed. To follow the 3R rules in animal testing (refine, reduce, and replace), such experiments are mandatory. Human immune cells used for research purpose can have two different origins: cell lines or primary cells. On one hand, cell lines are commercially available and present a phenotypic stability. Considering their widespread use in laboratories, a lot of studies are performed with cell lines, allowing to compare more easily results between different research groups. However, most of the time, such cell lines are immortalized cell lines derived from cancer patients’ tumors and can therefore not be classified as physiological. On the other hand, primary cells are directly cultured from their human source (patients’ organ tissues or blood cells) and are isolated in the laboratory. If the procedure to isolate those cells is available in most laboratories, unlike human cell lines, the phenotype of primary cells is not constant. Indeed, some variations can be observed from one patient to another and even between different batches from the same donor. The characterization of the cells prior their use is therefore mandatory. In addition, it requires to have access to blood samples. Primary immune cells can be isolated directly from blood or from buffy coat that is a fraction of an anticoagulated blood sample containing only white blood cells and platelets following density gradient centrifugation of the blood. Both sources are available from blood banks that check for viral infection to ensure the safety of the lab workers. Peripheral blood mononuclear cells (PBMCs) can be harvested by density gradient (Ficoll-Paque, Lymphoprep) from blood samples. Classically, the different mononuclear cells will be found at various proportions: lymphocytes B70%
90% (T cells B70% and B cells B15%), monocytes B5%, NK cells B10%, and DCs B1%
2% [67]. Characterization of the different subsets can be done by flow cytometry based on the expression of surface markers/Ags [cluster of differentiation (CD) classification system; Table 8.4]. The life span of the different subsets is variable due to their specific role in the immune homeostasis. In blood, monocytes survive for 3
7 days [68], NK cells for 2 weeks [69], DC cells for 1 week [70], whereas B cells survive up to 5
6 weeks, and T cells can be maintained in culture for weeks [71,72]. The different cell subsets can be isolated by magneticactivation cell sorting (MACS) or non-magnetic isolation (TACS technology) [73
75], by flow cytometry cell sorting, or by elutriation centrifugation [76,77] (for a complete review see Plouffe et al. [78]). In vitro, only the first steps of the immune response, that is, the detection of foreign material and the initiation of the immune response can be screened properly. The most widely used cell types are neutrophil granulocytes and macrophages.

Immune cells: sources, properties, and cell types

221

Table 8.4 Key surface markers that can be used for cell sorting. Cell type

Cell surface marker

T lymphocyte

CD3 CD4 (helper T cells) CD8 (cytotoxic T cells) CD19 CD20 CD11c CD123 HLA-DR CD56 CD34 CD14 CD33 CD66b CD16 CD11b

B lymphocyte Dendritic cell

NK cell Hematopoietic stem cell Macrophage/monocyte Neutrophil granulocytes

Source: From BD.

Neutrophil granulocytes that are of prime importance for the first interaction of the material with the tissue can be isolated from total peripheral blood. They are most of the time isolated with the red blood cells during the isolation of PBMCs with a density gradient [79]. Indeed, the layer located immediately above the erythrocytes layer contains mostly granulocytes. However, neutrophil granulocytes present a very short life span (12
20 hours) and their regular use is therefore difficult. Noteworthy, they can be isolated from the oral cavity with a simple mouth-rinsing protocol. Such neutrophil granulocytes subpopulation termed oral polymorphonuclear neutrophils present the same reactivity as the peripheral neutrophil granulocytes isolated from blood [80]. As they represent a critical element of oral homeostasis, their isolation can be of interest for dental material testing. Another source of patient’s immune cells is the bone marrow (BM), which is a primary lymphoid organ. It contains the hematopoietic stem cells (HSCs) that can differentiate into all the (white) blood cells types. The isolation of HSCs from BM aspirates is possible because they express the specific cell surface marker CD34. Aside from BM aspirates, HSCs can be isolated from umbilical cord blood [81,82] and even from the peripheral blood. Under steady-state conditions, HSCs circulate in the peripheral blood at a very low level but can be mobilized from their niches in the BM to the periphery by the administration of cytokines such as granulocyte colony stimulating factor [83,84]. The in vitro differentiation capability of CD34 1 cells was extensively studied. So far, differentiation in monocytes/macrophages [85], DCs [86], NK cells [87,88], B lymphocytes [89], T lymphocytes [90], and innate lymphoid cells [91] has been successfully implemented.

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Biomaterials for Organ and Tissue Regeneration

The cells line that are the most frequently used in bioengineering screening are those in front line with tissue repair/regeneration and foreign body reaction, that is, monocytes/macrophages. Monocytic or promonocytic cell lines such as THP-1, HL60 [92], U937 [93], ML2 [94], and MonoMac6 [95] are commercially available. THP-1 is probably the most popular cell line. It is a human monocytic cell line derived from the blood of a patient suffering from an acute monocytic leukemia [96]. It is considered as fully mature, but it expresses the vast majority of TLRs. It can be easily differentiated into macrophages by the use of phorbol-12-myristate13-acetate. What is of particular interest is that several modified THP-1 cell lines have been generated and allow to test different aspects of the immune response. THP1 Xblue cells incorporate the secreted alkaline phosphatase gene under the control of NF-kB and AP1, two transcription factors that are activated by TLRs. It, therefore, helps in the screening of biomaterials [97].

8.5

In vivo testing

In order to have a full vision of the interactions between a biomaterial and the immune system, it is mandatory to switch from in vitro to in vivo testing. The animal model of choice is probably the mouse. The work on wild type mouse models will of course give an overview, but the response of the murine immune system is not exactly the same as the response of the human immune system. It has been proven, for example, for cytokine expression and response to various stimuli in rodents [98]. Even for nonhuman hominids, some differences are described [99]. The answer to this problem could come from humanized mouse models. In such models the murine immune system is replaced by human cells. For example, a humanized mouse model can be developed by the implantation of human fetal thymus tissue and the injection of human CD34 1 fetal liver cells into an immunocompromised NSG mouse (NOD-scid IL2Rgamma null mouse; Jackson Laboratory). After engraftment with human immune cells, this murine model presents functional human T-, B cells, and DCs. The comparison of such model with the wild type model is a promising tool for the evaluation of biomaterials. For example, two different matrices derived from decellularized swine or human myocardium were tested. In such models the authors were able to show that the hydrogels generate a proinflammatory response in the early time after the implantation followed by a switch to a remodeling phase characterized by an infiltration of M2 macrophages and Th2 cells. However, it is during this switch that a difference was observed between the two murine models. Indeed, in the case of an allogenic transplantation (human decellularized pericardium), Th2 and M2 cells infiltration was reduced in the humanized mice compared to BALB/c wild type mice. However, the authors discussed the fact that only few human macrophages were found in the humanized mouse model with an infiltration of mouse macrophages instead. However, it shows that humanized mouse models mimic the potential human immune responses [100]. They could therefore be

Immune cells: sources, properties, and cell types

223

useful tools to assess the biocompatibility and proremodeling qualities of biomaterials prior to clinical translation.

8.6

Conclusion

The immune system is complex, dynamic, and highly adaptable. In this overview, our aim was not to give an exhaustive description but to point out the importance of each partner in healing process that could either lead to complete regeneration or defective cicatrization. It is, therefore, a key issue to consider in order to succeed in any biomaterial development. Keeping this in mind, in vitro and in vivo assessment of biomaterials should be carefully performed.

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Cell signaling and strategies to modulate cell behavior

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Claire Ehlinger1,2,3, Dominique Vautier1,2,3 and Leyla Kocgozlu1,2,3,4 1 UMR-S 1121 Inserm, Biomaterials and Bioengineering, Strasbourg, France, 2Faculty of Dental Medicine, Strasbourg University, Strasbourg, France, 3Federation of Translational Medicine, Strasbourg, France, 4Mechanobiology Institute, National University of Singapore, Singapore, Singapore

9.1

Introduction

In concert with responses to chemical signals, to hormones, to growth factors, and to cytokines, the cells respond to their mechanical environment according to the stress/strain conditions, the pressures, the viscosities, the confinements, and the topographies that they encounter. The cells develop a real “sense of touch” named “mechanosensing” to move, associate, dissociate, contract, or dilate. These mechanics of the living are present in each of the physiological processes of the cells when they adhere, migrate, proliferate, divide, repair lesions or activate a process of cellular extrusion, respond to inflammatory signals, reshape an embryo, differentiate, and degenerate into metastasis or die. By means of “mechanosensitive” proteins, cells are able to transform mechanical signals into biochemical signals. This process is termed “mechanotransduction” and makes the cell a mechanotransduction entity. The development of a functional tissue depends on an appropriate regulation of the cell behavior constituting it. At this end the reconstitution of the biochemical and physical signals of the microenvironment is a key element. Synthetic and natural polymers are materials of choice to integrate these different cues. This chapter focuses on recent progress discussing how cells respond to mechanical signals presented by biopolymer substrates mimicking the local extracellular matrix (ECM). Stem cells, very promising cells for tissue engineering, are highly plastic and can transition efficiently from one type to another. They have a specific molecular and physiological signature that drives their fate. The opening of chromatin appears as the priming early event in their reprogramming. Mechanisms by which mechanical signals achieved this phenomenon are just beginning to be elucidated. In particular, how mechanical signals modify the physics of the nucleus and the resulting consequences for the cells are emerging issues. Another exciting issue is how the cells memorize the mechanical signals, termed “mechanical memory,” interplays with a plethora of players to carefully orchestrate a phenotypic response [1,2].

Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00006-4 © 2020 Elsevier Ltd. All rights reserved.

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9.1.1 How to modulate cell adhesion, cell migration, and cell extrusion? Cell adhesion is one of the most important mechanisms studied to understand cell behavior. To develop strategies in tissue regeneration, new substrates for clinical applications modulating cell adhesion appear to be a key element. To modulate substrate rigidity to influence cell adhesion is one of the popular strategies [3]. During few decades, many substrates are developed to modulate cell adhesion as substrate rigidity, coating, cell confinement strategy in 2D, and for a few years in threedimensional (3D) to mimic cell environment [4
8]. Biophysics characteristics such as elastic and viscoelastic properties are involved in the mechanosensing response as cell adhesion. The most common substrates used to mimic microenvironment of cell are synthetic substrates such as polydimethylsiloxane (PDMS) [9], polyacrylamide (PAA) [10], polyethylene glycol [11], polyvinyl alcohol [12], widely used materials due to the ease of manufacture and biopolymers components such as ECM proteins such as collagen [13,14]. The polyelectrolyte multilayers (PEMs), corresponding to an alternate deposition of positively and negatively charged polyelectrolytes onto solid substrates [15], procure a wide range of rigidity and are useful to investigate the influence of stiffness to cell adhesion. The biophysical characteristics are mostly common for synthetic substrates and procure a range of substrate stiffness between few kPa to MPa. Depending on the cross-linked degree, PDMS materials provide a viscous state at weak cross-linked level versus an elastic behavior at high cross-linked state. The materials based on PAA procure only elastic behavior and across a stiffness range until 200 kPa [4,5,13,14]. The biopolymer substrates such as collagen provide a tunable stiffness matrix based on density and cross-linking degree of protein [16]. The behavior of the cells depends mostly on the natural microenvironment of mature cells. Engler et al. demonstrated for the first time the differentiation of mesenchymal stem cells (MSCs) depending on specific stiffness to neuronal, muscle, and osteogenic cells corresponding, respectively, to 0.1
1, 8
17, and 25
40 kPa [14]. The mechanotransduction of the cells studied to understand the cell response to the microenvironment is basically in three steps: the initial mechanosensing which occurs at the timescale of seconds and even less in subseconds (mechanical force propagation from the ns to μs); the early response of the cells is taken between second and minutes and involves the actin cytoskeleton and the cell motility actors; and the last step which occurs at the timescale to hours and days where the physical signals are converted to biochemicals signals and observed through cell morphology, cell proliferation, cell migration, cell death, etc. [17]. Cell adhesion due to mechanosensitivity occurs early with the integrin binding to the ECM and underlines that integrin mediates the cell response adapting cell shape, cell spreading, and motility to their microenvironment and its mechanical properties involving the recognition of substrate rigidity and ECM composition [18]. The other main parameter in the mechanotransduction pathway is the deformation of the substrate generated by the application of the force by the cell itself,

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the translation of the mechanical force will be transformed in biochemicals signals at the late response of the mechanosensitivity. To modulate mature cell with tunable stiffness, mimicking the stiffness of the biological microenvironment of the cell, different substrate conditions were tested from soft (50 kPa abbreviated E50) to stiff substrates (500 kPa abbreviated E500). In the study of the cellular response to the rigidity, the authors showed that substrate stiffness corresponding to 50 kPa inhibits focal adhesion assembly and actin fiber formation (Fig. 9.1) leading to the inhibition of replication, nevertheless, transcription activity is maintained. In response to a substrate with a stiffness of 200 kPa (E200) cell enables integrin activation, focal adhesion assembly (Fig. 9.1) accompanied by activation of FAK-Y397 kinase and replication becomes efficient and not interfere with activation of transcription [5]. The authors underline the strong relationship between cell adhesion and replication and propose a selective and uncoupled process of the substrate stiffness in the regulation of replication and transcription in epithelial cells. In tissue regeneration and wound healing process, increasing cell proliferation and promoting cell migration are required. Motility of the cells depends on mechanical conditions and stimuli. To understand the mechanisms of cell migration, researchers developed many strategies in 2D, 1D, and for a few years in 3D. Cell migration for isolated cell is mostly investigated for immune cells, neuronal cells, cancer cells specially during metastasis [13,19
21]. The actin cytoskeleton with their regulatory proteins constitutes a machinery driving force in the cell. This machinery generates protrusive forces (pushing) by polymerization of several actin filaments or contractile forces (pulling) via sliding actin filaments along the bipolar myosin II filaments. The migration process is linked to protrusive and contractile

Figure 9.1 αv-Integrin, actin fibers, and vinculin take place for different substrate elasticities. 4 hours of culture on E500 (500 kPa), E200 (200 kPa), and E50 (50 kPa), cells with anti-αv-integrin and anti-vinculin, labeled with phalloidin. Bar: 20 μm.

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forces exerted by the cell itself and related to the deformation of the microenvironment [22]. To understand specifically the role of actin fibers during cell protrusion, the deconstruction of various constrains is a strategy brought by the microfabrication techniques [23]. To study the rigidity sensing during polarization process which occurs during cell migration, microfabrication tools such as the combination of micropillars and rigidity could be required [9]. The actin cytoskeleton remodeling is the main driver during the polarization of the cell. To emphasize the large-scale mechanosensing probing from the cell, Trichet et al. [24] used micropillars substrate with different stiffness [25]. The authors demonstrate that the applied force on the substrate by the fibroblast cell increases with the focal adhesion area which is positively linked to the stiffness until to reach saturation at high rigidity corresponding to 85 nN/μm. The study reveals the strong coupling between the alignments of stress fibers with respect to stiffness increase. These results provide a better understanding between the cell cytoskeleton polarization observed by the large-scale mechanosensing given by the stress fibers organization and the directionality of mature cell migration which explain the durotaxis behavior. This process corresponds to a mode of cell migration in which cells are directed by rigidity gradients from physical properties of the ECM. Using a similar tool, another work led by the same team emphasizing the matrix rigidity sensing is guided by cell cytoskeleton remodeling and furthermore by an adaptive rheological response proven by a viscous behavior of actin cytoskeleton on soft substrates (9 nN/μm), whereas on stiff substrates (85 nN/μm) an elastic behavior is observed [9]. In other strategies to highlight the previous results, authors combined micropillars tools to study different stiffness and confined cell in an anisotropic shape to mimic polarized cell and isotropic shape for unpolarized cell [26]. Micropillars (Fig. 9.2A) are pillar-shaped microstructures that can be grouped together into artificial structure exhibiting many useful properties. As an example, micropillars are an approach commonly used to study forces in cellular systems. Each pillar has a defined spring constant that can be measured independently. The forces acting on a pillar can be linked with the deformation of the pillar imposed by cellular components. This approach underlines the actin organization under confined shape. In Fig. 9.2 a schematic of printing of confined shape on micropillars substrate with the contact printing of fibronectin with respect to unpolarized and polarized shape. On stiff substrate for all confined shape is given, Rat Embryonic Fibroblast cell [27] adopts an alignment of stress fibers (Fig. 9.2D). The combination of cell confinement and substrate rigidity demonstrates that both play a critical role in actin cytoskeleton remodeling, on high stiffness (85 nN/μm) the actin stress fibers reach an important length and aligned organization whereas at 38 nN/μm, we do observe a very well-organized fibers enriched with restitution of the bulk property of actin [26]. Collective cell migration is widely studied for tissue regeneration and to better understand cell extrusion which occurs to balance cell proliferation and maintain homeostasis [4,27]. To couple cell density evolution and cell extrusion rate, the authors confined epithelial cells to a circular pattern coated with fibronectin (Fig. 9.3A). The authors demonstrate that the increase of cell extrusion rate is linked

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Figure 9.2 Microforce sensing substrate and action organization: (A) images of micropillars PDMS substrate obtained with SEM, scale bar 1 μm; (B) functionalization process with fibronectin on the top of the micropillars; (C) fluorescent fibronectin micropatterned on the top of the pillars with the appropriate shape, scale bar 10 μm; and (D) actin cytoskeleton organization on 85 nN/μm stiffness. Representation of actin filaments labeled with phalloidin markers, scale bar 10 μm. PDMS, Polydimethylsiloxane; SEM, scanning electron microscope.

to the increase of cell density coupled to the velocity decrease (Fig. 9.3B
D). The contraction of the tissue observed for phase 1 is coupled to the local contraction of the tissue around the extruded cell [4], (Fig. 9.3E), when the density of the tissue is higher (phase 2), the entire contraction is lost and the local contraction during the extrusion process is less pronounced [4], (Fig. 9.3E, phase 2). This work demonstrates the link between the tissue packing and the local event as cell extrusion and emphasizes a distinct extrusion process guided by the tissue density, at low density cell crawling with lamellipodia extension play a key role during the extrusion process whereas for tightly packed tissue, a local actomyosin contraction takes in charge the extrusion event [4].

9.1.2 Synthetic matrices to control cell programming and reprogramming Totality of the genetic information of an organism, named genome, is sustained by its complete DNA sequence. The most essential role of DNA is to carry genes. The complex of DNA and protein is termed chromatin. Finally, specific information of an organism is content in its proteins and RNA. Nuclear organization relates to the

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Figure 9.3 Characterization of mesoscale behavior during tissue growth: (A) micropatterning process; (B) merged image with phase contrast and nuclei of Mardin-Darby canine kidney (MDCK) monolayer cells grown on adhesive patches at three different time evolution corresponding to low (left image), mean (middle), and high density (right image); (C) average cell density over the time; (D) quantification of extruding cells in function of time; and (E) time series of phase contrast image and corresponding velocity field around the extrusion point (t 5 0 hour) in phase 1 and phase 2.

location that particular regions of the genome occupy. The open, active euchromatin, which is efficient for gene activation, fills practically the nucleus whereas the condensed, inactive heterochromatin is restricted to an irregular ring situated at the nuclear periphery. The heterochromatin surrounds also the nucleolus and disseminates as aggregates in the nucleoplasm [28]. Epigenetic reprogramming operates exclusively in a single direction from pluripotency to differentiation [29]. Chromatin of pluripotent cells exhibits a typical open chromatin structuration, while differentiation results into significant remodeling and emergence of large dense chromatin areas [30]. Recent works evidence that an open chromatin organization is essential for the preservation or targeting of pluripotency [31]. To recover pluripotency a cell has to overcome various obstacles, such as the rewrite of epigenetic modifications and activation of key pluripotent genes [32]. Among the essential actors for somatic cell reprogramming to pluripotency [31], inhibitors of histone deacetylase (HDAC) and DNA methyltransferases have been identified. They are epigenetic modulators that potently facilitate reprogramming [33].

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Recent findings argue that physical properties of the cellular environment assume a crucial impact in cell fate [14,34] particularly in nuclear activity [5]. Recent reports showed that the Young modulus of the synthetic matrix acts on the chromatin organization and more precisely a soft matrix favors chromatin condensation [35,36]. Mechanotransduction phenomena by which cells feel their environment were widely studied on mechanosensitive proteins at focal adhesions and within the cytoskeleton [17,37
39]. Extra- and intracellular forces propagate across the cytoskeleton to the nucleus. These forces engage integrins at focal adhesions coupled to actin filaments, themselves connected to microtubules and to intermediate filaments. A protein complex, termed the linker of nucleoskeleton and cytoskeleton (LINC complex), establish a physical connection between the cytoskeleton and the nucleoskeleton. This complex permits force transfer across the nuclear envelope from the associated cytoskeletal filaments to the lamins [40
45]. These forces finally impact chromatin that constitutes a target of signal for activation or gene silencing [46,47]. Rabineau et al. [35] prepared hydrogels by alternate deposition of positively and negatively charged polyelectrolytes onto substrates leading to the formation of structured films, named PEMs. These hydrogels consisted of poly(L-lysine) and hyaluronic acid, coated with a poly(sodium styrene sulfonate)/poly(allylamine hydrochloride) (PSS/PAH) stratum, as a substrate simulating the ECM stiffness of physiological tissues [5,48,49], on which epithelial PtK2 cells were cultured. The stiffness of the substrate growth by adding PSS/PAH layer pairs [50]. Rabineau et al. demonstrated that on stiff substrates, chromatin was in its euchromatin configuration, while a soft substrate causes remodeling to some extent in its heterochromatin conformation. On a very soft substrate, cell goes through necrosis (Fig. 9.4). The initial hypothesis of these authors was to direct the cell behavior by altering the structure of the chromatin. The counterbalancing activities of histone acetyltransferases and HDACs dynamically regulate the acetylation of the chromatin and therefore the chromatin compaction, by, respectively, opening (euchromatin) or condensing (heterochromatin) chromatin [51]. Rabineau et al. [35] utilized a drug

Figure 9.4 Chromatin remodeling by substrate elasticity made of multilayer polyelectrolyte films. Ultrastructural images of PtK2 cells after 5 hours on (A) stiff substrate (200 kPa), (B) soft substrate (50 kPa), and (C) very soft substrate (0 kPa). Percentage of heterochromatin notifies the area heterochromatin at the total nuclear surface. Scale bar: 7 μm, arrow: heterochromatin, star: euchromatin.

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that blocks HDACs, trichostatin A (TSA), to preserve chromatin in euchromatin. On the very soft substrates, they observed that PtK2 cells incubated with TSA sustained acetylation of histones H3, preserved their chromatin in euchromatin and a diffuse nuclear localization of HP1β. These cells preserved also their nuclear envelopes unbroken and had a remaining intermediate filament network surrounding their nuclei. This permitted cells to live in a nonadherent condition without proliferation revealing that cells were in a quiescence state. Interestingly, when relocated on a rigid substrate, these cells maintained their ability to adhere, to spread, and to enter a novel mitotic cycle depending on their transcriptional activity. These data might be crucial to preserving cells in the best conditions within synthetic templates and in tissue-derived substrates utilized in tissue regeneration. In another study, Rabineau et al. used a hydrogel of elastic modulus of 20 kPa (E20) as a selective soft substrate for colorectal tumor cell survival. They showed that consecutive hard (culture glass)/soft substrates relocation of resistant cells elicited a significant opening of chromatin (Fig. 9.5) accompanied with an increase of cell survival, cell spreading, and cell motility. This survival seemed reversible, suggesting an adaptive event rather than irreversible gene mutation(s), as example mutation of the APC protein (adenomatous polyposis coli tumor suppressor gene), frequently expected during the malignant sequence of colorectal. This adaptation process was correlated with modifications in the gene expression profile. Importantly, expression of the homeobox gene CdX2 and the nuclear receptor family gene Hnf4α, two essential mediators of intestinal homeostasis with tumor suppressor function, was downregulated, whereas the Macc

Figure 9.5 Opening of chromatin by switching the substrate elasticity. Ultrastructural images of SW480 cells after 24 hours of culture on (A) first E20, (B) third E20. Percentage of chromatin in the nucleus. Scale bar: 8 μm, arrow: heterochromatin, star: euchromatin.

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1 gene involved in metastasis was stimulated. At the ultrastructural level the open chromatin induced by mechanical cues appeared to be very similar to those obtained chemically by TSA or by the molecule CYT296 [52]. A completely new way for chromatin opening exclusively founded on physical properties of the microenvironment without any drug mediation, preventing any side effects or antagonistic regulation, was thus evidenced [53]. In accordance, Anseth’s team demonstrates that prolonged exhibition of human MSCs (hMSCs) to stiff microenvironments directs chromatin decondensation in a persistent way [54].

9.1.3 Nuclear mechanics and mechanical memory The cells adapt to changes in the physical properties of tissues by modifying their gene expression profiles. Numerous biochemical signaling pathways involved in these responses have been elucidated. However, there is still little knowledge of the spatiotemporal relationships between the mechanical forces perceived at the level of the plasma membrane, the deformation of the nucleus, and the functional organization of chromatin. This part gives a brief overview of this new aspect of mechanobiology, centered on the physics of the nucleus. Iyer et al. applied mechanical forces to living cells using magnetic particles inserted into the plasma membrane. Using high-resolution fluorescence anisotropy imaging, the team revealed that mechanical forces induce actin polymerization and a change in actin F/actin G ratio. This variation results from the nuclear translocation of actin G in correlation with the import of the cofactor megakaryoblastic acute leukemia factor-1, linked to the transcription of RNA. Physical propagation of mechanical forces leads to nuclear deformation and reorganization of chromatin [42]. Versaevel et al. performed micromanipulations on isolated endothelial cells by varying the adhesion surface (making “micropatterns” of different shapes) and subjecting the cells to different compressive forces by hydrostatic pressure. They showed that nucleus deformation is due to lateral compression forces induced by the tension of actomyosin fibers. This tension modifies the spatial organization of the actin stress fibers and the geometry of the focal adhesion contacts allowing the elongation and spreading of the cell. Their results indicate that these changes in the shape of the cell and its nucleus are accompanied by a strong condensation of chromatin and a significant decrease in cell proliferation [55]. Using substrates consisting of micropillars mimicking the 3D confinement of the tissue environment, Booth-Gauthier et al. investigated the movement and nucleus deformation of healthy cells and cells with “Hutchinson
Gilford progeria syndrome.” The cells carrying this syndrome are engaged in an aging process. Their results show that the nuclei of these aging cells have increased dysmorphism and increased stiffness correlated with reduced mobility. The authors suggest that the nucleoskeleton shape defects (invagination, wrinkling, and bubble shape) of aging cells significantly alter the transmission of forces to the nucleus. These cells are then trapped in this 3D microenvironment [56]. Previous studies have shown that the transmission of cytoplasmic forces to the nucleus is mediated by a protein complex called the LINC complex, which physically binds the macromolecules of the cytoskeleton (actin filament, microtubule,

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and intermediate filament) to the nuclear envelope [40,41]. Anno et al. focused on the mechanical interaction between actin F and the nucleus via the nesprin-1, component of the LINC complex proteins. They developed a uniaxial cyclic stretching system to follow the nuclear deformations of endothelial cells, whose expression of the nesprin-1 gene was blocked by interference via siRNA. The results reveal that the nuclei of “nesprin-1-blocked” cells are smaller than the nuclei of wild-type cells. Under stretching the nesprin-1 deficient cells undergo a stronger compression exerted by the cortical actin layer. The absence of nesprin-1 induces the physical disconnection of actin-F from the nucleus. The nucleus is thus released from the tension forces that the actin-F is unable to longer transmit to it. Under these conditions the deformability of the nucleus is increased before mechanical stretching. This work specifies the essential role of the binding of actin fibers (actin-F structuration) to the nucleus via nesperin-1 allowing stable transmission of forces to the nucleus [57]. After imposing high amplitude deformations on HeLa cells, Haase and Pelling studied the return of the plasma membrane and nucleus to their initial shape. To perform these experiments, they used an Atomic Force Microscopy (AFM) as a nanoindenter coupled to a confocal microscope. The AFM imposes on living cell indentations of at most 5 μm with a maximum force of 20 nN over a period of 10 minutes. After retraction of the tip the plasma membrane and the actin cortex recover quickly their initial shape. They have shown that the cytoplasmic regions surrounding the nucleus are more resistant to long-term compression than the nuclear regions. This form of memory is supported by the intact actin cytoskeleton and the actomyosin contractile activity. In response to local deformations, these results suggest that the nucleus is weakly resistant to mechanical stress and does not play a determining role in restoring the shape of the cells. In contrast, the plasma membrane and actin cortex clearly provide remarkable mechanical stability to the cell [58]. Using a press system the Piel team is able to modify cell confinement. They thus determined a threshold strain for which the nuclear lamina breaks up and rebuilds, allowing significant changes in nuclear volume. These nuclear deformations are correlated with variations in the expression of genes involved in different intracellular mechanotransduction pathways [59]. Qin and Buehler are particularly interested in the physical response of the nuclear lamina to resist extreme mechanical deformation of 100%. In this work the nuclear lamina is considered as a structural protein meshwork of the nuclear membrane. The authors demonstrate that the nuclear lamina withstands the extreme mechanical deformation at the nanometric scale of unfolding, sliding, and transition of α and β chain proteins of the meshwork. These phenomena of extension of the protein meshwork prevented catastrophic crack propagation in the nuclear lamina. At the microscopic scale the propagation of mechanical failure to individual protein filament preserved the whole mesh [60]. Important work by the Discher group has revealed that the level of expression of nuclear lamina A, as a function of tissue elasticity, regulates the differentiation of MSCs [61]. The only way to produce a new cell is to duplicate an existing cell. This new cell inherits the exact copy of its mother’s genome. Daughter cells can also inherit a diversity of other “memories” from its mother cell such as proteins, RNA, and

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other “memory carriers” (such as growth factors or cyclins). In parallel, cells in a multicellular tissue are generally bound together by extracellular matrices that they synthesize. The matrix is critical for the organization and the function of the tissue. In particular, the mechanical properties of the matrix can instill to the cells heritable modifications in gene expression and/or protein activity. Memorization by the cells of these mechanical properties is termed “mechanical memory” [62]. Anseth’s team is one of the first to highlight the phenomenon of mechanical memory in hMSCs. They exposed hMSCs to supraphysiologically stiff cell culture grade polystyrene (Young’s modulus E B3 GPa). These cells were then transferred and cultured on soft poly(ethylene glycol) hydrogels (E B2 kPa). In this scheme of experiments the authors investigated the intracellular location of the yes-associated protein (YAP), identified as a molecular relay of mechanical signals exerted by the rigidity of the substrate. Nuclear import of the protein YAP requires Rho GTPase activity and tension of the actomyosin cytoskeleton [63,64]. Yang et al. showed YAP deactivation in hMSCs after 3 days on soft hydrogel. However, they evidenced a persistent activation of YAP after a prolonged duration of culture on rigid substrates followed by 3 days on the soft hydrogel [62]. Interestingly, they found that YAP was reversibly activated when the culture period on rigid substrate was short. Conversely, an extended culture of hMSCs on rigid substrates led to YAP activation eliciting irreversible lineage program [62]. In accordance with this work a mechanical memory has also been observed for MSCs by switching their biophysical microenvironments [65], or through changes in chromatin condensation for MSCs cultured on nanofibrous scaffold and subjected to dynamic tensile loading [36]. Pathak et al. developed a modular PAA substrate with regions of dissimilar ECM stiffness. They explored how the stiffness of the substrate instilled a “mechanical memory to migratory cells.” This process was shown to depend on YAP activity [66]. The mechanical properties of the microenvironment can potentially modify the cell fate. Cells have the ability to memorize these mechanical signals and continue their development under their influence, even when they are in contact with a different mechanical environment. While genetics refers to the written form of genes, epigenetics refers to their reading: the same gene can be read differently depending on the tissue or certain physiological responses. Indeed, epigenetic processes that permit heritable modifications in gene expression without changing the genetic sequence are one efficient means to turn an ephemeral environmental signal into a long-lived phenotypic change. A fascinating new question is open: how mechanical and epigenetic memories cooperate to orchestrate cell lineage commitment to ultimately produce a functioning organism or tissue.

9.2

Conclusion

In conclusion, to modulate cell behavior and cell signaling, studying cell in a various microenvironment specially with different stiffness/rigidity has been widely

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investigated. Cell type plays a critical role in the response to the substrate stiffness. With the progress of microfabrication and a large substrate mimicking microenvironment of the cell, great advances have been made in understanding cellular behavior. Cell adhesion is a key process in various biological processes such as cell migration, and widely mechanotransduction is directly linked to the stiffness of the substrate and opens the way to modulate and improves tissue regeneration. Furthermore, the coupling of stiffness and confinement techniques has allowed understanding the cell adhesion mechanosensivity and more precisely the largescale mechanosensing of the cell which involves the actin cytoskeleton. At the level of the tissue, for example, epithelial tissue, epithelial cohesion and tissue fluidity are a key element to maintain the robustness and the barrier of the epithelium. Cell extrusion events have been demonstrated to play a critical role to release excess of cell to keep homeostasis. With the help of micropatterning techniques and laser ablation systems coupled to a confocal microscopy, it has been shown two distinct mechanisms depend on cell packing which guide the release of the cell. One of these mechanisms involves collective cell crawling and the other one the actomyosin contraction of the neighboring cells around the extruded cell itself. For a few years, new techniques have been developed and allow to understand cell behavior in 3D microenvironment.

Acknowledgments We thank Chun Xi Wong for help in the design figures and Murat Shagirov for PIV analysis from Mechanobiology Institute from Singapore (MBI). Financial supports were received from National University of Singapore MBI, Alsace Contre le Cancer 2018. C. E. and L. K. are grateful to “Faculte´ de Chirurgie Dentaire” of Strasbourg for financial support.

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55.

Section 2 Biomaterials use in organ specific applications

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Cardiovascular tissue engineering

10

Richard A. O’Connor1,2, Paul A. Cahill2 and Garrett B. McGuinness1 1 Centre for Medical Engineering Research, School of Mechanical and Manufacturing Engineering, Dublin City University, Dublin, Ireland, 2Vascular Biology & Therapeutics Group, School of Biotechnology, Dublin City University, Dublin, Ireland

10.1

Introduction

Cardiovascular disease (CVD) accounts for an estimated 17.5 million global deaths each year and is currently the leading cause of mortality and morbidity in the world [1]. It is estimated that approximately 92.1 million US adults suffer from at least 1 type of CVD, with the mortality rate standing at 219.9 persons per 100,000 translating to 30.8% of all US deaths annually [2]. Increased public health awareness paired with early patient access and continuing advancements in medical device technology have resulted in declining CVD-related deaths across Europe and the United States in recent years [2,3]. Despite this, it remains a major source of economic concern with treatment estimated to have cost the US economy $316.1 billion in the year 2013 alone. Of this, $189.7 billion can be attributed to direct costs including hospital care and medication, and a further $126.4 billion attributed to the indirect loss of productivity due to patient absence from the workplace [2]. By 2030 an estimated $1 trillion will be spent on the direct medical costs of CVDs in the United States alone [2]. Coronary artery disease (CAD) is one of the most prevalent forms of CVD and can be attributed to approximately one of every seven deaths in the United States [1]. CAD is associated with the buildup of atheroma, also referred to as atherosclerotic plaques, within the coronary arteries of the heart. The coronary arteries supply the myocardium of the heart with oxygen and nutrient rich blood, which is essential for its efficacious functionality. Plaque formation results in both the stiffening of the arteries and the stenosis (narrowing) of the lumen preventing adequate blood flow reaching tissues downstream of blockage sites, ultimately resulting in medical conditions such as myocardial infarctions, known commonly as a heart attack. This chapter will address the potential for tissue engineering strategies to offer improved treatments for CVD, with particular focus on the development of tissueengineered blood vessels (TEBVs) for coronary artery bypass grafting (CABG), a pressing clinical need. An overview of the cardiovascular system and atherosclerotic disease is given, followed by a discussion of the CABG clinical procedure. Insights into the limitations of current approaches are considered, before setting out key requirements for successful scaffolds for TEBVs. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00011-8 © 2020 Elsevier Ltd. All rights reserved.

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The cardiovascular system

The cardiovascular system of the body consists of three essential components: the heart, the blood vessels and the fluid they contain, blood. Within this, blood vessels can be designated to be part of the systemic circulatory or pulmonary circulatory system [4]. The systemic system consists of the blood vessels that carry oxygenated blood away from the heart to the body and the vessels that return the deoxygenated blood back to the heart. The pulmonary circulatory system consists of the blood vessel loop that transports blood between the heart and the lungs. The right and left main coronary arteries, members of the systemic circulatory system, are small diameter blood vessels that branch from the ascending aorta at the right and left aortic sinuses, respectively. The right coronary artery encircles the right atrium and ventricle supplying them with oxygenated blood and has an internal diameter (ID) of 1.5
5.5 mm [5]. The left coronary artery supplies the left atrium and ventricle with oxygenated blood and is the larger of the two coronaries. It bifurcates into the left anterior descending (LAD) coronary artery with an ID range of 2.0
5.5 mm and the left circumflex coronary artery with ID 1.5
5.5 mm [4,5].

10.2.1 Arterial tissue Arteries are complex vessels whose architecture depends highly on their location within the vasculature system. Arteries can be categorized into two main groups, namely, elastic and muscular arteries. Elastic arteries are the largest arteries of the body, including the aorta and the common carotid, and are often located close to the heart. Muscular arteries are generally small to medium diameter vessels including the coronary arteries and the radial artery. All arteries contain three distinct layers referred to as the tunics: the tunica intima, tunica media, and the tunica adventitia [6,7] (Fig. 10.1). The tunica intima is the inner layer of the artery and comprises the endothelium and the basal lamina. It contributes little to the mechanical resistance of a vessel but is important for stable blood flow. The endothelium is a monolayer of endothelial cells (ECs) that are orientated in the direction of blood flow whose function is to provide a smooth nonthrombogenic lining allowing for smooth blood flow and selective diffusion [5]. A cushioning layer known as the basal lamina (approximately 80 nm in thickness) surrounds the endothelium allowing for the vessel to bend and change diameter under pulsatile flow conditions and provides a surface on which the ECs attach [8]. The basal lamina comprises collagen type IV, proteoglycans, and the adhesion molecules laminin and fibronectin. The tunica media is the mid layer of an artery and offers the main structural resistance of a vessel. The categorization of arteries as elastic or muscular depends on the composition of this layer. In elastic arteries the media layer is composed of 5
15 μm concentric rings of vascular smooth muscle cells (vSMCs) separated by sheets of elastic fibers. Thick elastic arteries may be composed of 40
70 of these

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Figure 10.1 Schematic cross section of an artery showing the three primary layers of arteries —the intima, media, and adventitia and their constituents [7]. Source: Credit: Reprinted by permission from Springer: Kluwer Academic Publishers, Journal of Elasticity and the Physical Science of Solids, A New Constitutive Framework for Arterial Wall Mechanics and a Comparative Study of Material Models, Holzapfel, G.A., Gasser, T.C. & Ogden, R.W., ©2000.

layers. Muscular arteries contain a single thick layer of helically orientated vSMCs embedded in a loose extracellular matrix of elastin and collagen fibers. The adventitial layer or the outer most layer of an artery is composed of fibroblasts, large-diameter type I collagen fibers, proteoglycans, nerves, and, in large to medium diameter vessels, the vasa vasorum [6]. Again, this layer does not contribute extensively to the mechanical resistance and response of the vessel but rather acts as a binding and anchorage layer to surrounding tissues.

10.2.2 Cardiac tissue Cardiac muscle tissue is found in the wall of the heart and is responsible for the pumping action of the heart, and the consequent pulsatile flow of blood in the arterial network. In common with skeletal muscle, cardiac muscle features striations are associated with myofibrils and sarcomeres. The contractile cells responsible for the actuation of the muscle are known as cardiomyocytes, with fibroblasts also present to perform their function in terms of production and maintenance of the extracellular matrix [collagen, elastin, and glycosaminoglycans (GAGs)].

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Heart valves also comprise an extracellular matrix with collagen, elastin, and GAG in different arrangements, and tissue-engineered heart valves (TEHVs) are also a significant focus of tissue engineering research. TEHV development has recently been the subject of a comprehensive review series [9,10].

10.3

Cardiovascular disease

Arteriosclerosis is the general classification of diseases that results in the thickening of artery walls and a loss of elasticity. Atherosclerosis is one form of arteriosclerosis and is a progressive disease associated with a chronic inflammatory response within arteries. A cascade of immune cell signaling and growth factor/cytokine secretion induce vSMC proliferation from the medial layer of the artery into the atherosclerotic legion. The vSMCs form a fibrous cap over the plaque protecting it from blood flow. A schematic of the lesion formation process can be seen in Fig. 10.2. Atherosclerotic plaques typically do not fully occlude the vessel with myocardial infractions predominantly occurring when the fibrous cap breaks open in response to chemical factors secreted by foam cells within the legion. These factors in turn lead to thrombosis formation that rapidly decreases the flow of blood to

Figure 10.2 Progression of atherosclerotic disease in a coronary artery. Source: Credit: Reprinted/adapted by permission from Springer Nature: Libby P. Insight: inflammation in atherosclerosis. Nature 2002;420:868
74, ©2002 [11].

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the myocardium [4,12,13]. Specific regions within the vasculature are highly prone to endothelium dysfunction and in turn atherosclerosis such as branches, bifurcations, and curvatures in arteries. This is hypothesized to occur due to increased turbulence and reduced shear stress at these sites [12].

10.4

Coronary artery bypass grafting

The development of noninvasive surgical techniques such as angioplasty and stenting along with pharmaceutical drug treatments has allowed for many CVDs, including stenosis of the coronary arteries to be relieved without the need for invasive surgeries [14,15]. In severe cases, coronary artery bypass grafting (CABG) may be required to fully revascularize the cardiac tissue. Bypass surgery is a highly invasive vessel-reconstruction procedure whereby a blocked coronary artery segment is bypassed with an alternative vessel referred to as a vascular graft. Fig. 10.3 shows a diagrammatic representation of a vascular graft used to bypass a blockage in the left coronary artery. Here a graft is used to connect the aortic root to the LAD coronary below the point of blockage, restoring an adequate supply of blood to the heart and the surrounding tissues. A number of noninvasive or minimally invasive techniques have been developed for the treatment of atherosclerotic lesion formation and the resulting medical complications. Drugs such as antihypertensives, beta blockers, and clot-dissolving agents have been designed to increase the blood supply to the heart and prevent thrombosis formation [14,17]. Percutaneous transluminal coronary angioplasty (PTCA) is a minimally invasive technique used to reopen narrowed coronary arteries, whereby an uninflated balloon is threaded through the vasculature system to the obstructed area and subsequently inflated stretching the arterial wall and

Figure 10.3 (A) Coronary artery bypass graft surgery and (B) compliance characteristics of current bypass grafts correlated with long-term patency characteristics [16]. Source: Credit: (A) with permission from Medmovie. (B) Salacinski HJ, Goldner S, Giudiceandrea A, Hamilton G, Seifalian AM, Edwards A, Carson RJ. The mechanical behavior of vascular grafts: a review. J Biomater Appl, 2001;15(3)241
78, ©2001. Reprinted by permission of SAGE Publications Ltd.

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compacting the atherosclerotic plaque [4]. PTCA is, however, a poor long-term management solution for atherosclerosis with approximately 42% of arteries experiencing restenosis within a 6 month period [18]. A variation of PTCA utilizes a fine wire tube known as a stent to aid in the maintenance of vessel patency. Despite improved patency rates, in-stent restenosis rates remain high with approximately 22.2% of patients’ arteries becoming reoccluded within 6 months of the procedure [19]. Coronary artery bypass grafting (CABG) is a highly invasive surgical procedure for the treatment of coronary stenosis with approximately 600,000 procedures performed each year in the United States [20]. In CABG, blockages within the coronary arteries are bypassed using small diameter vessels (,6 mm) referred to as bypass or vascular grafts. Oxygen and nutrient rich blood is channeled directly from the aorta through these vessels and into the coronaries below the point of blockage as demonstrated in Fig. 10.3. Continuous improvements in surgical techniques have led to an expected 30-day mortality rate of less than 1% and survival rates of 92% and 81% at 5- and 10-year postsurgery, respectively [21,22]. CABG has also demonstrated improved survival rates compared to percutaneous coronary intervention techniques particularly noticeable in multiple vessel disease treatment suggesting it is a superior long-term treatment strategy for coronary stenosis [23].

10.4.1 Vascular grafts Autologous vascular grafts, or grafts taken from an alternative site within the patient, are currently considered the gold standard vessels for use in arterial bypassing with the great saphenous vein (GSV) in the leg traditionally being the vessel of choice for small diameter applications [23]. Increased patient risk due to infection at the site of harvest and the declining supply of suitable host vessels in aging populations have, however, driven a search for an alternative source of suitable vascular grafts [24]. A number of synthetic grafts have been developed, including polyethylene terephthalate (PET, Dacron) and expanded polytetrafluoroethylene (ePTFE)
based vessels. Although proven successful in the bypassing of medium-to-large diameter blood vessels ( . 6 mm ID), they have shown poor patency rates when used in small diameter (,6 mm) applications, due to thrombus formation and anastomotic or intimal hyperplasia [8,25]. Due to the small diameter (,6 mm) nature of the coronary arteries, autologous vessels, which are vessels taken from an alternative site within the patient are considered the grafts of choice for use in CABG [26]. Their resemblance in size and mechanical properties to native tissues along with the removed risk of immune rejection makes them the optimum graft choice. The current gold standard vessel for CABG is the GSV [21]. Its long, bilateral nature and easy of accessibility has made it the favorable choice for surgeons [27]. Despite this, it does not fully match the biomechanical properties, namely, compliance, of the coronary arteries and has shown only a 66% patency rate at 10 years postsurgery [21]. Grafts derived from the internal mammary artery have demonstrated increased patency rates of up to 90% at 10 years; however, the difficulty in harvesting them has prevented their

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widespread use by surgeons [21,28]. With ever aging demographics the use of autologous vessels as a conceivable treatment strategy for coronary stenosis diminishes. Complications such as the degradation of host tissues along with previous myocardial or peripheral arterial reconstructions greatly limit the source of suitable vessels in ageing populations [25]. Allografts (grafts taken from a donor) have been studied as a possible solution to overcome this shortage. Fresh allograft tissues have been seen to experience rapid immune rejection and subsequent dissolution on implantation resulting in high short-term failure rates. Preserved and slightly degraded grafts offer increased clinical life due to decreased cellular content; however, they too experience immune rejection after some time. In order to successfully introduce allografts, immune suppression must be utilized that is not ideal for patients undergoing severely invasive surgeries [29]. A number of synthetic polymers have been employed in the development of vascular grafts for use in CABG. Synthetic approaches allow for the production of shelf ready grafts that offer convenience to surgeons along with reduced risk of secondary infection and immune rejection [27]. PET (Dacron) and ePTFE (GORE-TEX) have been used extensively in large ID (12
38 mm) and medium ID (6
10 mm) bypassing applications with significant clinical success, respectively [8]. These synthetic grafts, however, suffer from a mismatch in their compliance properties compared to the adjoining native vasculature. ePTFE demonstrates high occlusion rates with 64% patency at 1 year and 32% at 2 years postimplantation [10]. A lack of an EC lining also contributes to their failure due to turbulent hemodynamics resulting in thrombosis formation [30,31]. ePTFE grafts seeded with an endothelial lining have shown improved patency of 85% at 3 years and 65% at 9 years [32]. Due to the nonadherent nature of PTFE, the EC lining typically delaminates resulting in scaffold failure, so it has not be successfully adopted as a treatment strategy for small diameter bypassing [24,30].

10.4.2 Role of biomechanical compliance Failure of these synthetic grafts is often linked with the term “compliance mismatch” wherein the elastic properties of the bypass graft do not match those of the native arteries. This mismatch is suggested to generate flow instabilities that damage the endothelial lining of arteries leading to thrombus formation and intimal hyperplasia, ultimately resulting in long-term graft failure [33]. Fig. 10.3B shows the relationship between vessel compliance and patency rates for a range of vascular grafts used in femoral
popliteal bypassing. It is observed that with decreasing vessel compliance comes a reduction in vessel patency. The development of a small-diameter graft with biomechanical properties matching those of native coronary arteries is a pressing clinical need. In order to develop a suitable vascular graft for CABG, it is important to understand the biomechanical properties of the native tissues to be replicated. The unique architecture of soft tissues including vascular tissue gives rise to an equally unique set of biomechanical properties including nonlinearity, anisotropy, and viscoelasticity.

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Uniaxial tensile testing of vascular tissue has been performed by a number of researchers with nonlinear responses observed [34
36]. This nonlinear behavior arises due to the multicomponent composition of the arterial wall and is often referred to as the J-shaped profile of soft tissues. As reviewed previously, the layers of the arteries, most notably the tunica media, contain collagen and elastin fibers. Collagen is the primary load-bearing element of arteries and offers resistance during the systolic phase of the cardiac cycle. Collagen has a natural undulated morphology under normal physiological conditions. The elastin fibers of an artery provide a recoil effect that smoothens the flow of blood during the transition from the systole to diastole phase of the cardiac cycle. During initial loading of vascular tissue, elastin fibers provide the principal mechanical resistance. These fibers strain significantly under small applied loads giving a large initial strain region known as the “toe.” As loading increases, the undulated collagen fibers begin to straighten and align in the direction of loading. Collagen, the dominant load-bearing fiber does not strain significantly, so a stress-stiffening effect is witnessed. Fibrous and cellular content within the arterial walls are highly orientated aiding in resisting specific stresses experienced during pulsatile blood flow. As a result, arterial tissue is highly anisotropic with different stress
strain responses seen when loading occurs in the axial, circumferential, and radial directions. Circumferentially orientated intima and adventitia samples have been shown to stiffen at higher strains compared to longitudinal samples. In contrast the medial layers have been shown to stiffen at lower strains in the circumferential orientation compared to the longitudinal [7,36]. Pure elastic solids strain instantaneously when loaded and regain their original dimensions when the stress is removed. Newtonian fluids on the other hand exhibit a constant rate of strain when a constant viscous stress is applied and do not retract when the stress is removed. Arterial tissue contains both solid material and fluid constituents, so it exhibits a combination of these properties, a behavior known as viscoelasticity. Viscoelastic materials demonstrate creep, stress relaxation, and hysteresis characteristics. Creep is the time-dependent change in strain due to a constant stress, while stress relaxation is characterized by a decrease in stress under a constant strain [37]. In vivo arteries are subjected to both longitudinal stretching and internal pressurization. When the material is excised from the body, the corresponding stresses are released [38]. This stress release results in the samples visibly shrinking in the longitudinal and circumferential directions, indicating the sample has entered a no-load state [37]. If a sample is then loaded, unloaded, and reloaded, the stress
strain response can be seen to change between the first and second loading cycles. This noncoincidental loading path occurs due to the reorientation of fibers and cells within the tissue. If cyclic loading and unloading is performed a sufficient number of times, a stable stress versus strain loop is achieved, the sample is now said to be in a preconditioned state [37,39]. It can also be observed that the loading and unloading cycles do not follow an identical path, a characteristic known as hysteresis. This hysteresis occurs due to the internal changes of the tissue structure coupled with energy dissipation caused by friction forces generated between fibers during loading [38].

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Another important property of vascular tissue is compliance. Compliance is a measure of vessel distensibility and is defined as the ratio between the change in volume (ΔV) or diameter (Δф) of a vessel for a given internal pressure change (ΔP) [40]. Compliance is a significant characteristic to be considered when designing a vascular graft as numerous studies have shown the significance of reducing “compliance mismatch” for maintaining long-term graft patency [41
43]. A number of formulas exist for calculating the compliance of a vessel with each formula dependent on the experimental methodology used. The general formula for the compliance of a vessel is given by [43] Compliance 5


 Ds 2 Dd 3 104 %per mmHg 3 1022 Dd 3 ðPs 2 Pd Þ

(10.1)

where D and P are vessel diameter and blood pressure, respectively. Subscript d and s refer to diastole and systole measurements.

10.5

Tissue-engineered blood vessels

Continuous advancements in engineering and the life sciences have driven researchers toward the interdisciplinary field of tissue engineering as a possible solution to the shortage of suitable vessels for use in coronary artery bypass grafting (CABG) [25,44
46]. Tissue engineering combines the principles of engineering and the life sciences in order to develop biological substitutes that can be used to restore or replace tissue and organ function [47]. By combining cells with highly porous 3D biodegradable structures referred to as scaffolds, replacement tissue structures can be created either in vitro or in vivo [48]. These scaffolds act as templates for newly populating cells, offering stability and guidance during the early stages of tissue formation. The seeded cells are anticipated to secrete extracellular matrix (ECM) components allowing for the formation of the relevant ECM for the target tissue as the polymer degrades away [26]. The final envisaged product is a stand-alone biological construct that aims to match the properties of the native tissue to be replaced. The scaffolds used in this approach should provide a number of functions and features. Their architecture should provide sufficient void volume to allow for revascularization where applicable and to facilitate host tissue regeneration [33]. In addition, the biomaterials used within the scaffold should be biocompatible allowing for cells to attach, grow, and differentiate during in vitro and in vivo culturing [8]. Bioactivity of the scaffold is also desirable in order to facilitate and regulate cellular activity. This can include the addition of adhesion molecules to the scaffold surface to aid in cell attachment and appropriate surface topography to induce correct cell morphology and alignment as would be seen in the host tissue [49]. The incorporation of growth factors within the matrix to stimulate tissue regeneration is also of great interest [50]. The mechanical properties of the scaffold are also key

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attributes and highly depend on the shape and size of the defect or tissue to be regenerated [51]. The intrinsic mechanical attributes of the material should match that of the native host tissue to be replicated. The long-term success rates of CABG procedures are limited due to atherosclerosis and occlusion of the bypass vessel. The failures of these grafts are attributed to compliance mismatch between the bypass grafts and the native vessels [33,41,43]. Compliance mismatch arises due to a change in the rigidity or elasticity of a blood vessel along its length such as a native artery meeting a synthetic graft [41]. Physical variation of the artery diameter along with the pulsatile nature of blood flow result in altered wall shear rates [40]. Increased shearing may damage the EC lining giving rise to the onset of atherosclerosis and intimal hyperplasia, while decreased shearing leads to blood stagnation and in turn thrombosis formation [16,52]. Compliance mismatch between stenotic and distal segments of arteries has also been linked with the development of unstable atherosclerotic lesions leading to the onset of myocardial infarctions [53] or cerebrovascular attacks [54]. Intimal hyperplasia is the thickening of the tunica intima of an artery and is a biological response to the injury of a blood vessel. An increased growth of the intima due to vSMC proliferation and differentiation can lead to the graft becoming stenosed or occluded similar to atherosclerosis. vSMC remodeling occurs in response to the damage of the endothelial lining and is a healing response [42]. Intimal hyperplasia prominently occurs at the anastomosis (junctions) of two vessels of different compliances and is hypothesized to occur due to increased flow instabilities at these sites [55]. A schematic of intimal hyperplasia development at the distal anastomosis of a bypass vessel can be seen in Fig. 10.4 [40,52]. The compliance properties of a number of synthetic and biological grafts used in CABG for a range of mean pressures can be observed in Fig. 10.5. Synthetic vessels such as Dacron, ePTFE, and compliant poly(carbonate)polyurethane do not experience significant volume changes with increased pressurization. In comparison, biological tissues exhibit large volume changes at low-pressure levels and small volume rate changes at high pressures. It is noted that veins such as the GSV rapidly stiffen with increasing pressure compared to arteries, which show a

Figure 10.4 Intimal hyperplasia formation at the distal end of the vessel at anastomosis site [42]. Source: Credit: Reprinted/adapted by permission from Elsevier: Ballyk PD, Walsh C, Butany J, Ojha M. Compliance mismatch may promote graft
artery intimal hyperplasia by altering suture-line stresses. J Biomech ©1997.

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Figure 10.5 Compliance-mean pressure curves for vessels and grafts used in CABG [43]. Source: Credit: Reprinted/adapted by permission from Wiley: Compliance properties of conduits used in vascular reconstruction, Br J Surg, 2000;87(11):1516
24, ©2000. Table 10.1 Compliance and patency rates for arterial and venal grafts [33].

Host artery Saphenous vein Umbilical vein Bovine heterograft Dacron ePTFE

Compliancea

1-year % patency

2-year % patency

5.9 4.4 3.7 2.6 1.9 1.6


88 83 65 65 60


84 80 59 42 42

ePTFE, Expanded polytetrafluoroethylene. a % Radial change per mmHg 3 1022.

smoother transitional stiffening effect as pressurization increases indicating compliance mismatch issues amongst autologous vessels. The mean compliance values for arterial, venous and synthetic bypass grafts along with their patency at 1 and 2 years postimplantation are shown in Table 10.1. Increased compliance directly correlates with long-term patency and ultimately the success of the grafts. Limitations of current synthetic vascular grafts and the declining supplies of autologous vessels with ageing demographics highlights an imperative clinical need for a readily available bypass graft with improved biomechanical and compatibility

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Table 10.2 Properties of the ideal tissue-engineered blood vessel (TEBV) for use in vascular bypassing as set out by Walpoth and Bowlin. Properties of the ideal vascular graft G

G

G

G

G

G

G

G

Biocompatible (nontoxic and nonthrombogenic) Leak resistant, but adequate porosity for healing/tissue regeneration Compliance matching that of native artery Resistant to aneurysm formation Postimplantation durability after tissue ingrowth Easily manufactured, stored, and sterilized Suture retention Flexibility with kink resistance

Source: Adapted from Walpoth BH, Bowlin, GL. The daunting quest for a small diameter vascular graft. Expert Rev Med Dev, 2005;2(6):647
51.

properties. Tissue engineering approaches may provide a solution to this problem through the development of biological replacement vessels. These vessels should better mimic the structure and biomechanics of native coronary arteries and in turn out perform current vascular grafts. Considerable research has focused on the development of small diameter TEBVs for use in CABG. This approach provides a potentially sustainable, inexpensive, and effective solution to the current shortage of suitable bypass vessels. By designing these TEBVs to have equivalent structural, mechanical, and chemical properties as the native arterial structures they are replacing, vessels with appropriate longterm patency rates may be achieved [24,25,44,45,55]. Walpoth and Bowlin outlined the key properties they believed an ideal TEBV should possess, as summarized in Table 10.2. These criteria included ease of surgical handling, leak resistance to prevent critical blood loss, sufficient porosity to allow for tissue regeneration, and compliance characteristics matching native vasculature [25]. Tissue engineers should strive to develop scaffolds that achieve as many of these characteristics as possible, while maintaining a feasible design strategy. In 1986, Weinberg and Bell produced the first completely biological TEBV [46]. In this effort a collagen gel combined with bovine aortic cells was studied. The mechanical properties of the resulting scaffold were insufficient for in vivo implantation and despite the addition of a Dacron mesh for reinforcement, the vessel lacked the required burst strength characteristics. Konig et al. utilized a complete biological approach that removed the need for an exogenous scaffold. In this approach, human fibroblast cells were cultured until cell sheets were formed [56]. Subsequently these cell sheets were wrapped around a mandrel to produce a tubular scaffold and further cultured until layer fusion was achieved. The tissue was later dehydrated in air for several hours in order to form an acellular construct. This approach yielded a scaffold with good mechanical properties but required 3 months in culture before a graft that was sufficiently stable for implantation was produced. L’Heureux et al. similarly produced a completely biological tissue-engineered

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construct by culturing human vSMCs with ascorbic acid until cellular sheets formed. These cellular sheets were wrapped around a tubular mandrel to form the medial layer of the vessel. Sheets of human fibroblasts were subsequently cultured and wrapped around the vSMC layer to form an adventitia-like layer that comprised dense collagenous material. Following maturation, the tubes were removed and the lumen seeded with ECs. The TEBVs formed possessed burst pressures of 2000 mmHg but lacked the dense SMC stricture typically witnessed in native arterial tissue. When implanted, the vessels were prone to thrombosis formation resulting in overall patency rates of only 50% [57]. Hoerstrup combined nonwoven meshes of poly(glycolic acid) (PGA) with thin layers of poly-4-hydroxybutyrate, ECs, and myofibroblasts to produce their vessel [44]. Their research demonstrated advanced tissue formation through the use of “biomimetic” pulsatile culturing compared to commonly used static culture techniques. Burst strengths of approximately 3000 mmHg were achieved after 28 days of culture. These burst strengths are well above the hypertensive blood pressure threshold of 140 mmHg [20], and comparable with the burst strength values recorded for coronary arteries and the saphenous vein at 3000 6 500 [58], and 1599 6 877 mmHg, respectively [59]. Similar studies utilizing PGA meshes have again demonstrated that the addition of cyclic strain during culturing periods increased vessel burst strengths to approximately 2000 mmHg after 8 weeks in culture [45]. While these studies have taken important steps toward the development of suitable TEBVs, they also highlight deficiencies in the selection of suitable biomaterials. In order to reduce the need for prolonged culturing times, biomaterials with mechanical properties suitable for direct implantation should be considered. This will allow for in vivo stability, while vessel remodeling occurs within the patient.

10.5.1 Biomaterials for tissue-engineered blood vessels The identification and selection of appropriate biomaterials and fabrication techniques for the creation of 3D scaffolds is an ongoing source of interest for tissue engineers. Biomaterials used in these applications should offer mechanical stability while maintaining a suitable microenvironment for the growth and proliferation of the seeded cells [26]. In addition, these materials should be biocompatible, blood compatible, fatigue resistant, and ideally biodegradable [8]. Biodegradable materials should degrade in vitro and in vivo into either products that are normal metabolites of the body or products that can be eliminated with or without further metabolic transformations [26]. Two main classes of materials have been studied extensively for the development of TEBVs, namely, natural biomaterials and synthetic polymers. Natural biomaterials used in TEBV applications generally consist of protein components found within the natural ECM of arteries. The ECM is the structure left within a tissue once all cells are removed [38]. Scaffolds based on vascular ECM proteins include those derived from collagen [46,59,60], elastin [61],

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fibronectin [62], and combinations of these proteins [63,64]. Other natural biomaterials commonly explored include fibrin, a protein involved in the clotting of blood [65,66], and polymers derived from biological sources such as those of the polyhydroxyalkanoate family [67]. Biological scaffolds derived from natural biomaterials offer increased biocompatibility, nontoxicity, and increased cell adherence [68]. Despite these favorable characteristics, they suffer from poor mechanical properties, a lack of reproducibility when processed, rapid degradation, and can induce immune reaction and pose disease transmission risks if taken from a nonautologous source [68]. Synthetic polymers have been investigated due to their ability to create reproducible, inexpensive, and easily manufactured scaffolds. A number of nonbiodegradable synthetic polymers including ePTFE and Dacron have been studied as potential candidate materials. These materials have shown poor cellular interaction and mechanical characteristics ultimately leading to their poor success in small diameter applications [25,30]. Numerous biodegradable synthetic polymers have also been utilized, including PGA [44,45,69], poly(lactide-co-glycolide) [70,71], and poly(ε-caprolactone) [55,72
76]. The disadvantages of using synthetic polymers are the potential to induce an inflammatory response upon implantation, a problem commonly seen with such materials [77]. This inflammatory response includes platelet, macrophage and leukocyte activation which are known triggers for the initiation of thrombogenesis and intimal hyperplasia formation [52]. Despite this, the synthetic flexibility these polymers offer make them highly desirable for use as tissue scaffolds as they can be manipulated to offer the desired properties. By altering their chemical structure, the physical and biomechanical properties of various tissues can be replicated [26]. Studies have also shown the ability to blend these polymers with additional biomolecules that aid in increased cell adherence and proliferation [49,78,79]. Furthermore, surface and chemical modification of the synthetic polymers can be used to help reduce the inflammatory response they typically initiate upon implantation [80].

10.5.2 Stem cells in tissue-engineered blood vessel applications In order to develop an in vitro
based tissue-engineered construct a source of cells is often required in order to initiate population of the scaffold body prior to implantation. These cells are expected to proliferate and produce ECM as the scaffold degrades, allowing for stand-alone tissue structures to be achieved. The main successes in this field have come from the use of primary cells, taken from the patient, and used in conjunction with scaffolds to produce tissue constructs for reimplantation [81]. This method, however, suffers from a number of limitations including the invasive nature of cell collection and the potential for cells to be in a diseased state upon harvest [82]. An appreciation of the inherent diversities of organ systems paired with the difficulty of harvesting primary cells has therefore led researchers to focus upon the use of stem cells, including embryonic stem cells, mesenchymal stem cells, and induced pluripotent stem cells for tissue-engineering applications [83,84]. Stem cells provide a readily available source of cells that have the potential

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to be differentiated into a number of cell lineages [85], provide better proliferation rates compared to primary cell lines [58], while also retaining the ability to induce a reduced immune response upon implantation [86]. One of the critical steps of stem cell usage for regenerative medicine is the ability to control the differentiation of the cells into the desired tissue lineages. This is achieved through the control of their micro- and nano-environment [87], along with appropriately activating differentiation pathways through biochemical stimuli [88]. It is, therefore, desirable to use a stem cell niche that naturally aids in the creation and self-repair of a particular tissue of interest in order to better mimic the natural tissue formation process. As discussed previously the unique biomechanical and architecture properties of blood vessels require the careful consideration and selection of a cell-scaffold construct to ensure successful functionality upon implantation. A stem cell lineage that aids in vascular tissue repair has the potential to increase the long-term success of such scaffolds through the appropriate natural differentiation of cells and subsequent remodeling of the vessel. Recently, there has been evidence demonstrating the presence of resident vascular stem cells in the tunica media of arties, which have been coined as multipotent vascular stem cells (MVSCs) [89]. These cells have been shown to proliferate and differentiate into vascular SMCs whilst also be shown to possess the potential to be differentiated into neural cells, and mesenchymal stem such as cells giving rise to their multipotent classification [90]. Lineagetracing experiments have also importantly shown that these MVSCs are not derived from mature SMCs but rather a distinct cell group residing within the vascular walls. The unique nature of this stem cell group makes them an attractive potential for future TEBV studies.

10.6

Electrospinning of tissue-engineered blood vessels

Electrospinning is a scaffold fabrication technique capable of producing nano- to micron-scale diameter fibers that has been studied extensively in the last decade for the creation of tissue-engineered vascular grafts (TEVGs) [26,91]. Its ability to create fibrous structures that resemble the body’s natural ECM in both scale and architecture, along with high surface area
to-volume ratios and tuneable mechanical characteristics has made it a versatile scaffold production technology to study [92]. A significant challenge in the development of electrospun TEVGs is a lack of cell penetration and the insufficient transport of nutrients throughout the scaffold body, due to the small size, complex distribution and lack of connectivity of pores within electrospun structures [92,93]. Inadequate penetration of cells into the scaffold depths prevents complete remodeling of the tissue occurring throughout, often leading to graft failure as the fibers begin to degrade. A number of techniques have been proposed within the literature for the development of electrospun materials that demonstrate enhanced porosity characteristics such as porogen leaching [94,95] and sacrificial fiber spinning [96]. These techniques do, however, typically suffer from a number of limitations including a loss of mechanical integrity that has prevented them from being successfully employed [97].

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Two variations of the electrospinning process that may provide a potential route for the creation of electrospun scaffolds that demonstrate increased cell penetration properties are multimodal fiber spinning and dynamic liquid electrospinning. Multimodal electrospinning allows for the simultaneous creation of nano and micron-scale diameter fibers from a single spinning solution. This method offers increased pore sizes due to the crossing of large micron scale fibers within deposition planes, while intermixed nanoscale diameter fibers offer anchorage sites for cell attachment. The second approach considered is the production of linear bundles of electrospun fibers through dynamic liquid electrospinning, typically by depositing fibers into a circulating water vortex and continuously drawing fibrous bundles off to a rotating mandrel. Membranes formed from these fibrous bundles have been shown to exhibit increased porosity and pore volumes compared to traditional electrospun structures while also offering increased postprocessing potential due to their ability to be incorporated in to techniques such as knitting and weaving [98
100].

10.6.1 Fundamentals of electrospinning Electrospinning is a polymer-process technique capable of producing continuous fibers with diameters ranging from microns to nanometers in size [101,102]. Despite being patented over a century ago, the potential of electrospinning for biomedical applications was not realized until the 1990s [103]. At the laboratory scale the basic electrospinning setup consists of a high voltage DC power supply (1
30 kV), a capillary with attached spinneret (syringe with a flat tip needle), a syringe pump and a grounded metallic collector system. A schematic of the electrospinning setup can be seen in 1.6 [92,104]. The electrospinning process produces fine diameter fibers through the simple principle of applying a uniaxial stretching force to a viscoelastic solution [92]. The process begins by drawing a polymer solution usually derived from a raw polymer dissolved within a mono- or poly-solvent system into a syringe. A flat tip metallic needle is connected to the syringe and the solution is extruded at a constant and controlled rate using a syringe pump (Fig. 10.6).

10.6.2 Electrospinning parameters The electrospinning process and the morphology of fibers produced through it are highly dependent on a number of parameters. These parameters include solution properties, processing parameters, and ambient conditions. A summarized list of these parameters and their typical working ranges can be seen in Table 10.3 [105,106].

10.6.3 Collector systems for creating electrospun vessels Another important aspect of electrospinning is the type of collector system employed. Standard flat meshes with randomly orientated fibers are created by spinning onto aluminum foil or flat metallic plates [107]. In order to achieve

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Figure 10.6 The electrospinning process (A) schematic of the basic electrospinning setup with a flat metallic plate collector. Electrostatic forces result in Taylor’s cone formation and subsequent jet initiation. An scanning electron microscope (SEM) image of a typical 2D random fiber mesh is also shown [105] and (B) trajectory of fluid jet with instabilities regions shown [104]. Source: Credit: (A) Reprinted/adapted by permission from Elsevier: Tan SH, Inai R, Kotaki M, Ramakrishna S. Systematic parameter study for ultra-fine fiber. Polymer (Guildf). 2005;46(16):6128
34. (B) Reprinted/adapted by permission from Wiley. Li D, Xia Y. Electrospinning of nanofibers: reinventing the wheel? Adv Mater 2004;16(14):1151
70, ©2004. Table 10.3 Parameters known to affect fiber formation during electrospinning and typical working ranges. Solution parameters

Equipment parameters

Ambient conditions

Molecular weight (104
107 g/mol) Polymer conc. (1
40 wt.%) Surface tension (20
75 mN/m) Conductivity (0.05
30 mS/m) Solution volatility

Applied voltage (5
100 kV) Flow rate (0.01
1 mL/min) Collector characteristics Needle size Distance from tip to collector (1
100 cm)

Humidity Temperature Pressure

Source: Adapted from Tan SH, Inai R, Kotaki M, Ramakrishna S. Systematic parameter study for ultra-fine fiber fabrication via electrospinning process, Polym (Guildf.) 2005;46(16):6128
34; Wendorff JH, Agarwal S, Greiner A. Electrospinning. Weinheim, Germany: Wiley-VCH Verlag GmbH & Co. KGaA, 2012.

constructs with increased fiber alignment and complex geometries, a number of researchers have fashioned novel collector systems. The simplest method to achieve aligned fibers is through the use of rotating metallic drums. Matthews and collaborators utilized a grounded rotating drum spinning between 500 and 4500 RPM to

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collect collagen nanofibers [59]. Samples made at 500 RPM demonstrated random fiber orientation, while those created at 4500 RPM showed highly aligned substructures with fibers orientating themselves along the circumferential axis of the drum. Xu et al. again formed highly aligned structures but this time in the form of thin poly(L-lactic acid-co-caprolactone) filaments [108]. In this system a large diameter disk with a sharpened edge was used to collect thin linear bundles of fibers.

10.6.4 Limitations of electrospun scaffolds While the selection of appropriate biomaterials for use in the development of electrospun TEBVs is paramount to their success, perhaps, the most limiting factor facing this fabrication technique is the morphological attributes of the fiber constructs themselves. Researchers have demonstrated the increased adherence and proliferation properties that electrospun nanofiber-based materials offer compared to those of the micron scale [109
111]; however, a clear limiting factor in their success to date is the inadequate cell infiltration rates currently observed [93,112,113]. This lack of cell penetration has been attributed to the small size, complex distribution, and lack of connectivity of pores within nanofiber-based electrospun scaffolds [114]. The porous nature of electrospun structures arise due to the crossing of fibers within 2D deposition planes [106]. As fiber diameters decrease toward the nanoscale, their packing density greatly increases resulting in a severe reduction in the pore volume between fibers [115]. Without sufficient pore volumes and interconnectivity, cells cannot fully penetrate electrospun scaffolds and become trapped within the scaffold peripheries, often termed the “fishnet effect” [116]. This lack of complete infiltration prevents appropriate tissue remodeling occurring throughout the full thickness of the scaffold [117], along with cell necrosis due to inadequate nutrient and waste diffusion [114]. Deficiencies in new tissue generation can ultimately lead to the long-term failure of the grafts as the scaffold material degrades [113].

10.7

Future outlook for cardiovascular tissue engineering

The development of tissue engineering research toward successful clinical translation and commercialization is continuing apace [9,10]. It is a demand-led industry, and current trends toward personalized medical products have paved the way for the development of more medical devices with adjunct cell therapies. By mid-2007, according to Lysaght et al., approximately 50 firms with more than 3000 employees had generated sales in excess of $1.3 billion [9]. A recent review encountered 49 public tissue engineering companies employing nearly 146,000 people and generating an estimated h9 billion in sales of tissue-engineered products in 2017 [10]. Despite early positive predictions for tissue engineering and the market pull, the

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actual status is that a large innovation gap exists. Small diameter vascular grafts remain as a fundamental challenge to the field, alongside demands for more biologically compatible products tailored specifically to individuals using autologous cells, as outlined in this chapter.

Acknowledgments The authors gratefully acknowledge supported received from the Irish Research Council Postgraduate Scholarship Scheme under the Embark Initiative (Grant RS/2012/52) and Horizon 2020 (H2020-NMBP-2017), PANBioRA: Personalized, and/or Generalized Integrated Biomaterial Risk Assessment.

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54.

Bioartificial gut—current state of small intestinal tissue engineering

11

1 Thomas Daullary ¨ , Christina Fey1,2, Constantin Berger1, Marco Metzger1,2 and Daniela Zdzieblo1,2 1 Tissue Engineering and Regenerative Medicine, University Hospital Wu¨rzburg, Wu¨rzburg, Germany, 2Translational Center Regenerative Therapies, Fraunhofer Institute for Silicate Research ISC, Wu¨rzburg, Germany

11.1

Introduction

The small intestine is a multifunctional organ important for food digestion, absorption, and transportation of nutrients. The intestinal epithelium further plays a role as a supportive component of the body’s immune system and functions as a protective barrier to prevent infiltration of harmful substances or microbes into the body. In the past, animal models and human primary or cancer-derived cell lines represented the gold standard to study organ-specific functions, tissue homeostasis, and physiology of the small intestine. While fundamental breakthroughs were achieved, all of these models fail to closely mimic human in vivo
like features. Improved tissue models that reflect the in vivo situation more accurately are inspired by state-of-the-art and next-generation tissue engineering (TE) concepts. TE-based culture systems combine biological, material, biochemical, and/or biomechanical cues to support cell survival, behavior, and function of cells in vitro with the aim to stimulate the self-organization into native-like tissues. Tissueengineered in vitro models of the small intestine hold great promise to understand organ function and physiology under healthy and pathologic conditions in more detail. In this context, artificial organ substitutes are of great interest to model intestinal diseases such as celiac disease or inflammatory bowel diseases (IBDs) as well as for the study of host
pathogen interactions. In the following sections, we will review the main cell types used for small intestinal tissue engineering and provide an overview of the current innovative in vitro models.

11.2

The small intestine—structural organization and function

The small intestine is located in the abdomen connecting the stomach with the large intestine on a distance of 3
6 m in adult humans (Fig. 11.1A). It fulfills various tasks, including food digestion, absorption, and transportation of nutrients. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00025-8 © 2020 Elsevier Ltd. All rights reserved.

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Figure 11.1 Anatomy and cell types of the small intestine. (A) The small intestine lies in the abdomen connecting the stomach (gray) with the colon (brown) and can be subdivided into duodenum (green), jejunum (blue), and ileum (red). (B) The small intestinal wall is built of different layers (from the outside to the inside: tunica serosa, linear and circular muscle layers, tela submucosa, mucosa). Three layers form the mucosa: the lamina muscularis, lamina propria, and the intestinal epithelium. (C) Submucosal protrusions (plicae circularis) covered by the mucosa lead to an enlargement of the small intestinal resorption surface. Vasculature enables nutrient supply, while the nervous system transfers signals and stimuli. (D) Villi intestinalis are projections of the lamina propria lined by a monolayer of epithelial cells. At their base, villi structures segue into cavities called crypts. At the bottom of each crypt, SCs reside within the SC niche. SC-derived TACs leave the SC niche and migrate along the villus axis while differentiating into specialized epithelial cell types. (E) Regulation of cellular identity and function is mainly driven by signaling gradients whereby stem cell
specific growth factors comprising Noggin (NOG), EGF, signaling factors of the Wnt (Continued)

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L

Furthermore, the small intestine represents a protective barrier to prevent infiltration of harmful substances or microbes into the body and plays a role as a supportive component of the immune system. In order to fulfill this number of various roles, the organ exhibits a unique anatomy with localized functionalization characterized by the spatial division into the proximal duodenum, the connecting jejunum, and the distal ileum (see also Box 11.1). Digestion of food, including the resorption of nutrients, is efficiently performed by the proximal parts of the small intestine. The distal part, particularly the terminal section of the ileum, is of main importance for the immune system and resorption of special nutrients such as vitamin B12, bile salts or remaining nondigested remnants that passed the jejunum. Structurally, the small intestine represents a hollow tube built up by multiple layers (Fig. 11.1B and C), each with a distinct function. The tunica serosa represents the outer sheet of the tube, separating the small intestine from the peritoneal space of the body. It comprises a secretory mesothelium (lamina epithelialis serosae) and a layer of connective tissue (lamina propria serosae), intervened with vessels and nerve fibers, which segues into the tela subserosa. Adjacent to the tela subserosa, the tunica muscularis, an external muscular layer consisting of a thin longitudinal and a thick circular muscle layer, regulates the peristaltic movement by coordinated contraction patterns, thereby ensuring directed transportation of the chyme. The tela submucosa, a highly vascularized collagenous structure, connects the outer muscular layer with the inner part of the intestinal wall. It carries arterial, venous, and lymphatic vessels to supply the overlying Mucosa. The luminal Mucosa itself consists of three layers: The lamina muscularis, lamina propria, and the luminal epithelium. The lamina muscularis is a thin layer of smooth muscle cells, which ensures mixing of the luminal content by local contractions. The overlying lamina propria consists of loose connective tissue and immune cells. It further comprises lymphatic and vascular capillaries, formed by branches of the submucosal plexus. This network ensures nutrient transport and energy supply of the overlying epithelium that lines the inner surface as a prismatic cell monolayer. The small intestinal epithelium is specialized on the secretion of mucus, enzymes, and ions required for food digestion and the efficient absorption of nutrients such as sugars, amino acids, lipids, fatty acids, and carbohydrates [8,9]. Another important role of the epithelium is its barrier function that is crucial for separating the inner body from the intestinal environment, while simultaneously allowing trans- and paracellular transport [10,11]. On the cellular level, barrier integrity is built and maintained by the expression of tight

pathway (WNT3A) and R-Spondin 1 (RSPO) maintain stemness. In contrast, differentiationspecific factors including BMP, HH, and EFNB1 regulate differentiation along the villus. Different milieus of oxygen levels along the crypt-villi axis contribute to stem cell maintenance and differentiation. (F) Overview of the main cell types found in the intestinal epithelium with its main functions, average abundance, and turnover rate. BMP, Bone morphogenic protein; EFNB1, ephrin B1; EGF, epidermal growth factor; HH, hedgehog; SC, stem cell; TAC, transit amplifying cell.

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Box 11.1 Characteristics of the small intestine segments The duodenum is the most proximal part of the small intestine and builds the transition to the gut. On average, the duodenum of a human adult extends over a distance of 25 cm with a tube diameter of 2.5 cm. Next to nutrient resorption, a mayor task of the duodenum is the processing of the gut chyme prior to digestion, which involves neutralization by duodenal bicarbonate secretion and mixing of the chyme. Submucosal glands called glandulae submucosae (Brunner’s glands) protect the intestinal epithelium and promote food digestion by secretion of mucus, bicarbonate, and enzymatic precursors [1]. They are solely present in the duodenum. Chyme neutralization and digestion are further supported by pancreas juice, which is secreted into the duodenum via the papilla duodeni mayor. The main resorption takes place in the jejunum, which makes up 2/5 of the small intestine. It seamlessly follows the duodenum with continued surface enlargements for maximized nutrient resorption. In distal direction the jejunum segue into the ileum which represents 3/5 of the small intestine with a slightly smaller tube diameter of 2 cm. Epithelial cells of the ileum are specialized in the resorption of B12 and bile salts. In distal direction the histological architecture of the ileum increasingly approaches the structure of the large intestine. Peyer’s patches (PPs) are a special feature of the ileum [1]. These aggregates of single lymphoid follicles are in close contact with M cells of the overlying epithelium. By this, PPs act as an immunological sensor of the intestine milieu and play a crucial role in the decision making between tolerating and eliminating intestinal microorganisms. In adaption to the solidification of the chyme with increased transit time the distal part of the small intestine exhibits fewer and shorter plicae circularis, while they are absent in the distal part of the ileum [1
3]. In addition, average height of villi intestinalis decreases in distal direction (636 μm vs 537 μm) whereas crypt depth increases (80
176 μm vs 132
219 μm) [4,5]. Duodenal villi structures appear leaf like, while villi of the jejunum and ileum resemble more tongue- or finger-like morphologies. In rats, thickness of the mucus layer increases from proximal to distal (170
480 μm) but is thinnest in the jejunum (123 μm) [6]. Furthermore, pore size of tight junctions which defines the leakiness of the epithelial barrier in transcellular transport decreases in distal direction (0.75
0.8 nm vs 0.3
0.35 nm), which goes along with increasing electrical resistance and potential difference (0
3 mV vs 1
6 mV) [7].

and adherence junction proteins including E-cadherin, a range of claudin and occludin proteins as well as the junctional adhesion molecule A [12
15]. A remarkable feature of the small intestinal epithelium is the characteristic microstructure represented by crypts and villi (Fig. 11.1D). Crypts are tube-shaped invaginations of the intestinal epithelium harboring multipotent stem cells (SCs)

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under nearby normoxic (pO2B80
100 mmHg) conditions in vivo (Fig. 11.1E) [16,17]. Intestinal SCs (ISCs) are able to proliferate and differentiate into all intestinal cell types, thereby regenerating the epithelium every 3
5 days [18
21]. Adult ISCs have been first described as leucine-rich repeat-containing G-protein coupled receptor 5 positive (LGR51) cells [19]. Tian et al. provided evidence for a second “reserve” SC population expressing the oncogene “Bmi1” (B lymphoma Mo-MLV insertion region 1 homolog) that can act as an alternative SC pool under diseased or stressed environmental conditions [21]. Both cell types reside within the SC niche at the bottom of the crypt (Fig. 11.1D) where they undergo asymmetric cell divisions [22,23]. One daughter cell retains SC characteristics and resides within the SC zone, while the second progeny represents a “transit amplifying cell” (TAC) that proliferates within the TA zone of the crypt (Fig. 11.1D). After several rounds of proliferation, TACs migrate along the villus structure, thereby differentiating into specialized intestinal cell types that undergo anoikis at the tip of the villi [23
25]. Further localized within the crypt are Paneth cells, which represent longliving defensive cells secreting antimicrobial peptides, cytokines, and proteases [26,27]. In addition, Paneth cells secrete important growth factors such as Wnt3a (WNT), Noggin (NOG), R-Spondin 1 (RSPO), or epidermal growth factor (EGF) that sustain stemness within the SC niche [28
30]. While the concentration of these factors is high within the crypt, protein gradients of these factors decrease toward the villi (Fig. 11.1E) [16,31]. In contrast, the concentration of bone morphogenetic proteins, hedgehog protein and ephrin B1 is high within the villus region regulating differentiation of TACs [16,32,33]. In contrast to the crypts the villus structures represent elongated projections of the small intestinal lumen to enlarge the resorption surface, thereby maximizing the nutrient uptake capacity of the epithelium. Size and morphology of these structures vary within the small intestine (see Box 11.1) [34]. In contrast to the normoxic conditions within the crypt, cells located within the villi grow under near hypoxic conditions (pO2 , 10 mmHg) [16,17,35]. Enterocytes represent the most prominent cell population of the villi epithelium that secrete digestive enzymes playing a role in nutrient digestion [36]. Furthermore, enterocytes are responsible for nutrient absorption mediated by specialized transporters [37]. Their luminal surface is equipped with microstructures called “microvilli” that further increase the uptake of luminal nutrients, vitamins or water [34,38]. The absorptive enterocytes form a polarized cell layer, which is interspersed by a range of other specialized cells with distinct functions including goblet cells, enteroendocrine cells, Tuft cells, and microfold (M) cells [39]. Goblet cells produce mucus that consists mainly of Mucin-2 protein and antimicrobial peptides serving as a physical and biological protective layer against pathogenic microbes [40,41]. Enteroendocrine cells comprise several different cell subtypes, playing distinct roles, for example, for appetite regulation, blood glucose homeostasis or gut contractility by releasing a range of hormones, such as serotonin, somatostatin, glucose-dependent insulinotropic polypeptide or glucagon-like peptide (Worthington et al., 2018). Compared to other organs, the small intestine is confronted with a tremendous amount of environmental microbes including pathogenic and commensal ones. On the one site, beneficial microbiota must be kept viable to

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sustain nutrient supply and protection. On the other, overgrowth and pathogenic infections must be avoided. The main contributor to this microbial homeostasis is the mucus-secreting goblet cells that actively respond to invading microbes by flush exocytosis of mucus [40]. Tolerating the versatile intestine flora while protecting the body from invading harmful pathogens requires a concerted immunological system within the small intestine. Tuft cells and M cells located in the villi represent immunological active cells and play an important role in protection against pathogenic microbes. Tuft cells are able to initiate type 2 immune responses in context of intestinal protozoa and helminth infections [42,43]. In addition, they play a role as chemosensory cells [43]. M cells are concentrated at regions called Peyer’s patches describing immune-related parts of the ileum, where they are in close contact with the innate immune system in the underlying basal lamina [44]. They participate in phagocytosis and transcytosis of gut lumen molecules, antigens, and microbes as well as their subsequent presentation to leukocytes situated at the basolateral site of the intestinal epithelium (Mabbott et al., 2013).

11.3

Modeling the small intestine—biology meets engineering

Tissue-engineered in vitro models of the small intestine hold great promise to understand organ function and physiology under healthy and pathologic conditions. They are of great importance regarding regenerative strategies for intestinal diseases such as cystic fibrosis or IBDs. In the following parts, we review the main cell sources used for small intestinal tissue engineering and highlight current approaches to develop next-generation models of the human small intestine.

11.3.1 Modeling the small intestine in vitro by two-dimensional monolayer cell cultures In the past, small intestinal tissue engineering was represented by two-dimensional (2D) monolayer systems for the short-term culture of primary intestinal cell types. Modifying cell culture conditions such as media compositions and cell culture surfaces improved the survival and growth of human primary cell types. However, establishment of long-term cultures could not be achieved mainly due to the lack of important cell
cell or cell
matrix interactions. In addition to primary cells isolated from healthy tissue, cell lines derived from intestinal adenocarcinomas such as Caco-2, HT29, or T84 were generated in the past. Caco-2 and HT29 cells represent two of the most widely used in vitro models of the human small intestinal epithelium. However, each of these cell lines harbor specific characteristics that need to be considered when using these cells in different applications. Caco-2 cells are able to spontaneously differentiate into a tight epithelial barrier consisting of polarized enterocyte-like cells displaying a brush border membrane under common 2D monolayer culture conditions (Fig. 11.2A) [46
49]. Caco-2 cells further express digestive enzymes similar to the native small intestine but their

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Figure 11.2 Small intestinal cells cultured under 2D or 3D conditions. Overview of common 2D or 3D culture conditions for small intestinal cells. (A) Caco-2 cells are routinely grown as 2D monolayer cultures in standard tissue culture plastic as shown by a representative microscope picture. Scale bar 5 400 μm. (B) Transwell culture platforms represent state-of-the-art systems to grow Caco-2 cells in 3D on, for example, biological scaffolds such as the SIS [45]. After single-cell formation, Caco-2 cells form a multilayer on the biological SIS scaffold as shown by representative microscope pictures of H&E-stained in vitro models (B, upper panel). Immunohistochemistry-stained sections of 3D Caco-2 models demonstrate the expression of intestinal markers such as villin or pCK (B, lower panel; villin is shown in green, pCK is shown in magenta, DAPI counterstaining is shown in blue). Scale bar 5 100 μm. (C) Primary small intestinal cells can be grown as 3D organoid cultures enclosed in Matrigel that preserves an in vivo
like extracellular matrix surrounding. Scale bar 5 400 μm. (D) Similar to the Caco-2 cells, primary cell
derived organoid cultures can be grown on Transwell-like culture systems based on the biological SIS scaffold. Those cultures show a distinct monolayer (H&E-stained models shown in the upper panel in D) with the expression of several intestinal epithelial markers such as villin, lysozyme, mucin-2, mucin-1, or chromogranin-A (not shown). Representative microscope pictures of in vitro models stained for Villin or pCK are shown in the lower panel in (D) (villin is shown in green, pCK is shown in magenta, DAPI counterstaining is shown in blue). Scale bar 5 100 μm. (E) The complexity of the Transwell-like 3D models built on the SIS with primary organoid cells allows the investigation of, for example, epithelial barrier function or infection studies to investigate host
pathogen interactions. In (E), representative microscope pictures are shown for an organoid-based Transwell model that was infected with Salmonella enterica (upper panel: H&E-staining; lower panel: immunohistochemistry staining). Infection with S. enterica led to morphological chances of the monolayer shown in the upper panel in (E) with specific localization of the bacteria inside epithelial cells [shown by the yellow LPS signal in the lower panel in (E); DAPI counterstaining is shown in blue, pCK staining is shown in magenta]. Scale bar 5 100 μm. 2D, Two-dimensional; 3D, three-dimensional; pCK, pan-Cytokeratin; SIS, small intestinal submucosa.

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tight junction protein expression profile is more colon-specific, a characteristic that puts its use as small intestinal in vitro model into question [48,50,51]. Nevertheless, Caco-2 monolayer cultures are still widely used for absorption or transport studies, especially when cultured in the Transwell system that is described in more detail next (see also Fig. 11.2B) [52
58]. In contrast to Caco-2 the colon cancer
derived HT29 cell line represents undifferentiated cells that are not able to form a tight barrier under standard cell culture conditions [46,59,60]. However, the adaption of cell culture media triggers cell differentiation defining HT29 cell cultures as a suitable model to study smallintestinal differentiation processes in vitro [61
63]. While Caco-2 cells exhibit enterocyte-like phenotypes, HT29 cell cultures generally demonstrate a goblet cell
specific phenotype [64,65]. Subcultivation of HT29 cells in presence of methotrexate (HT29-MTX) stabilizes the mucus-secreting goblet cell-like phenotype demonstrating the HT29-MTX cell line as a valuable model to study diffusion of food or drug compounds across the intestinal mucus layer [64,66
68]. In view of their specific phenotype and associated functional features, Caco-2 and HT29/HT29-MTX cells are often cultured together. Such coculture systems combine a monolayer of absorptive enterocyte-like cells with mucin-secreting goblet cells [69
76]. Recently, Ferraretto et al. developed a Caco-2/HT29 coculture model that closely mimics structural and functional features of the human intestinal epithelium by using parental cells, which made the use of cellular subclones or growth factor
mediated in vitro differentiation redundant [71]. In the last years, small intestinal coculture models have been improved by integrating more cell types with relevant roles like fibroblasts, enteric glia or immune cells [77,78]. Lozoya-Agullo et al. recently studied transport mechanisms of diverse drug formulations using a triple-culture model comprising B-lymphocytes and Caco-2/HT29-MTX cells cultures [79]. In summary, studies on monolayer cultures of the small intestine led to important findings regarding physiology and function. However, the use of tumor-based models is limited, as they do not represent all cellular subtypes of the native intestinal epithelium and exhibit artificial gene and protein expression profiles which do not adequately reflect the in vivo situation.

11.3.2 Small intestinal organoids—artificial mini organs grown in vitro The small intestinal crypt hosts multipotent SCs with the ability to generate all types of the intestinal epithelium. In view of this key feature, ISCs represent the functional unit for tissue regeneration in vivo. In addition, ISCs harbor this remarkable capacity also in vitro showing the ability to proliferate, differentiate, and selforganize into three-dimensional (3D) structures called organoids closely mimicking in vivo
like tissue organization and cellular composition (Fig. 11.2C) [80
82]. Intestinal organoids also known as “mini-guts” or “mini-intestines” are currently of high interest in basic, pharmacological, and translational research.

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The fascinating era of organoid technology started with the ability to isolate organoid-forming units from the native intestinal epithelium, a milestone in small intestinal tissue engineering. Evans and Potten first achieved isolation and in vitro growth of the organoid-forming units in the late 1990s. The authors established monolayer cultures from the rat intestinal epithelium containing polarized epithelial cells including the ISC population (Evans, 1992). However, long-term culture of intestinal organoids was not feasible until innovative experimental techniques were developed. Ootani et al. established air
liquid interface cultures that allowed the formation of organoids from the murine neonatal intestinal epithelium when cocultured with mesenchymal fibroblasts as stromal compartment [83]. In the same year, Sato et al. invented a 3D culture system based on the extracellular matrix (ECM) “Matrigel” and a specified culture medium composed of the key growth factors Wnt3a, EGF, Noggin, and R-Spondin 1 that maintain stemness in vivo and in vitro [29,30]. At this, the 3D Matrigel environment provides cell
ECM contacts and represents an in vitro stromal unit that prevents extensive anoikis and supports epithelial cells in maintaining an apical-basolateral cell polarity [29,84
86]. Simultaneously, this culture condition allows the generation of budding protrusions from the central spherical domain of the organoid containing LGR51 SCs and differentiated epithelial cell types [29,30,82,87
89]. Nowadays, in vitro growth of intestinal organoids is routinely performed and protocols have been successfully converted to humans [29]. Long-term human organoid culture generally requires the growth factors Wnt3A, EGF, Noggin, and R-Spondin 1 as well as nicotinamide, N-acetylcysteine, B27, and small molecules for the inhibition of p38 kinase (SB202190) and a TGFβ pathway inhibitor (usually ALK4/5/7 inhibitors) [90]. Nevertheless, protocols and growth conditions are still modified and adapted. Withdrawal of Wnt3A, nicotinamide and the p38 kinase inhibitor “SB202190” result in a more complex structure of the organoid including the existence of mature differentiated cell types [29]. Recently, Fujii et al. published a refined culture system based on high-throughput single-cell RNA profiling data that supports cellular diversity in human intestinal organoids [88]. The authors showed that replacement of EGF and the small molecule p38 inhibitor “SB202190” by insulin-like growth factor-1 and fibroblast growth factor-2 maintains proliferation of ISCs in vitro [88]. In addition, this refined media stimulates multilineage differentiation of ISCs, thereby generating organoids that represent a cellular diversity more closely to the native tissue composition [88]. Compared to the classical growth media cocktail, this advanced organoid culture also improves CRISPR-based genome engineering efficiencies [88].

11.3.2.1 Organoid cell sources The technology of organoid culture enables the re-creation of valuable in vitro models of the small intestine as cyst-like structures with a central lumen surrounded by a highly polarized epithelium close to the intestinal villus [90]. In general, there are two different cell sources for the generation of human small intestinal organoids comprising primary intestinal tissue (intestinal crypts or isolated SCs) and human pluripotent SCs such as embryonic SCs or induced pluripotent SCs (iPSCs).

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Primary intestinal crypts are isolated from surgical or endoscopic biopsy. Organoids grown from primary tissue are known as enteroids [91] that retain a donor-specific identity including a precise genetic background. Therefore this cell source is beneficial for the establishment of healthy- and disease-related biobanks (van de Wetering et al., 2015), but clearly complicate standardization and reproducibility for in vitro cell culture applications [92]. Also, there are ethical concerns when using primary tissue
derived intestinal organoids and further worries if they maintain genetic stability in long-term cultures (Dutton et al., 2019). In 2011 two protocols were published to derive intestinal organoids from human iPSCs by stepwise differentiation [93,94]. Human iPSC
derived organoids (HIOs) are multilayered structures that contain major cell types of the small intestinal epithelium mimicking basic architecture and physiological functions in vitro. A significant drawback of HIOs is the immature, fetal-like phenotype of the generated intestinal cell types [95]. Maturation of HIOs was only achieved after implantation in vivo until Jung et al. described the in vitro maturation based on coculture of HIOs with T lymphocytes [96]. Further, the authors identified the STAT3activating interleukin-2 as a major factor for in vitro maturation that leads to the development of adult-like phenotypes [96].

11.3.2.2 Organoids as in vitro tools to model or study intestinal diseases Patient-specific organoids often reflect disease-associated pathologies characterizing them as highly interesting tools for translational research purposes. In this context, Dekkers et al. published an elegant study in 2013 that describes the development of a quantitative assay to study CFTR (cystic fibrosis transmembrane conductance regulator) function in intestinal organoid cultures [97]. The assay function is based on the capability of Forskolin to induce swelling of intestinal organoids, a feature that is absent in CFTR mutant organoids derived from patients suffering from cystic fibrosis [97,98]. Today, this so-called swelling test is widely accepted to prove CFTR function. Recently, Schwank et al. performed CRISPR/Cas9-mediated genome correction within patient-derived CF organoids and used the Forskolinstimulated swelling test to prove the correction of the CFTR mutation [99]. Nowadays, the assay further represents a potential drug screening platform and is of great advantage when testing individualized medication for cystic fibrosis patients. IBDs, such as colitis ulcerosa or Crohn’s disease (CD), are chronic inflammatory diseases affecting the colon and the small intestine. Molecular and cellular mechanisms underlying those diseases are poorly understood. Patient-derived organoid cultures therefore represent powerful research platforms to study disease-associated pathologies or pathomechanisms. In this context, long-term culture of IBD patient
derived organoids show similar mRNA expression profiles for inflammatory factors such as IL-1β compared to the corresponding initial biopsy [100,101]. This proves the preservation of the inflammatory state in vitro and hence indicates patient-derived organoids as potential tools for IBD research. In 2018 Suzuki et al. established CD-specific organoid cultures and performed single-cell RNA analyses

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to investigate the correlation of an inflammatory environment and alterations of the small ISC pool [102]. Finally, the authors could demonstrate a modified gene expression profile of the disease-specific ISC population that was associated with altered SC properties represented by the high capacity of CD-associated ISCs to rebuild organoids in vitro [102]. Another interesting approach in the context of translational research defines the organoid structure as “Trojan horse” that is in principal able to deliver drug-loaded nanoparticles to inflammatory sites thereby enabling local drug-release in future in vivo studies [103]. Besides disease modeling or studying disease-related pathomechanisms, organoid technology enables further host
pathogen studies in a more physiological manner which have been so far difficult to investigate (reviewed in more detail by [104]). Here, we want to mention one elegant study that was recently published by Heo et al. demonstrating the organoid culture as a relevant model to study Cryptosporidium infection in vitro [105]. However, besides all benefits, a major drawback in using organoid cultures in host
microbe interaction studies lies in organoid structure. In vivo, microbial growth, and microbiotic infection take place within the intestinal lumen that is represented in vitro by the luminal compartment of the organoid surrounded by the epithelium. Technically, the pathogen of interest has to be microinjected into the lumen of the organoid, which still represents a difficult and low-throughput methodology. In this context, Williamson et al. invented a next-generation approach by developing an automated platform for highthroughput organoid microinjection that can be monitored by fluorescent and brightfield microscopy [106]. The use of organoids as tools for SC-based therapeutic approaches is still under investigation. Recent studies demonstrated the successful transplantation of human colon-derived organoids orthopically into murine recipients, while maintaining selfrenewal and multipotency of the LGR51 population [107,108]. Particularly, Sugimoto et al. demonstrated the successful integration and self-organization of human intestinal epithelial cells into the epithelium of the murine-host by tracing ISC-derived cell derivatives [107]. Together, these studies and others indicate a potential use of ISC-derived intestinal organoids in context of future regenerative therapies.

11.4

Small intestinal tissue engineering in the Transwell—when cells meet scaffolds

Classical 2D monolayer cultures are of great advantage for standardization and scalability of in vitro cultures; however, these models clearly lack in vivo
like cell characteristics and only marginally represent functional and physiological features of the native organ. In contrast, intestinal organoids resemble the in vivo situation more closely [29,30]. However, both systems mainly miss mechanical or biophysical cues of the native ECM. Furthermore, especially in view of the morphological characteristics of the small intestinal epithelium, 2D monolayers grown in standard

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tissue culture plates fail to provide a luminal and basal compartment, which limits the applicability of these systems in terms of absorption or transportation studies. Similar to the monolayer cultures, there are also technical limitations in performing absorption or transportation studies on small intestinal organoids when grown in a 3D Matrigel
based environment. An elegant, TE-based concept uses diverse biomaterials as scaffolds to culture intestinal cell types within so-called Transwell systems (Fig. 11.2B, D, and E; see also Fig. 11.3). On the basis of a synthetic or biological membrane, this cell culture format provides a suitable microenvironment supporting cell adherence and further allow proliferation and differentiation of cells. In addition, Transwell-based culture systems provide an apical and basal compartment mimicking the luminal and basal side of the small intestinal epithelium in vitro. Especially this feature characterizes the Transwell platform as a suitable format in the context of absorption or transport studies. Current Transwell systems are based on either synthetic matrices, including polyester (PE), polycarbonate (PC) or polyethylene terephthalate, which are available with various pore sizes and thicknesses as well as biological scaffolds such as decellularized small intestinal submucosa (SIS) [45,109,110]. While standardized synthetic matrices harbor defined characteristics, they require additional coating procedures with ECM
proteins to enable cell attachment or proliferation, to sustain carrier-mediated transporter expression and to maintain cellular morphology or permeability characteristics [110]. Furthermore, synthetic membranes might influence diffusion and distribution of biological or chemical compounds. In contrast, biological matrices resemble a biocompatible 3D microenvironment; however, they often fail in the context of standardization and/or reproducibility [111]. To our knowledge, a detailed comparison of synthetic and biological matrices focusing on cell viability, differentiation, transport properties, etc. is yet not given in the community.

11.5

Next-generation models—integration of microenvironmental factors

Modeling the small intestine with functional and physiological relevance requires the integration of in vivo
like microenvironmental factors as depicted in Fig. 11.3. Engineered culture systems must incorporate the tissue-specific 3D structure as well as biomechanical and biophysical cues. State-of-the-art Transwell cultures often utilize scaffolds comprising either ECM-coated synthetic membranes or biological ECMs. In this regard the choice of scaffold is of high importance. Besides allowing proper cell attachment as well as proliferation and differentiation of ISCs [112], the scaffold must enable and support the self-formation of organized monolayers and further maintain long-term stability of the grown tissue. To meet these demands, TE-based concepts are currently applied to design next-generation culture systems that mimic natural environmental cues in vitro. Those approaches include biomimetic scaffolds and microfluidic devices.

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Figure 11.3 Designing small intestinal in vitro models with physiological relevance. Overview of five important factors that need to be considered when engineering optimized small intestinal in vitro models. In vivo
like cellular diversity is essential to incorporate diverse functions in artificial model systems and should mimic the cellular composition of the native epithelium comprising stem cells and differentiated cell types as well as immunological components and the microbiome. Biomechanical stimuli like the application of flow or the simulation of contraction as well as the integration of native-like mucus or the use of extracellular matrix scaffolds have to be considered to properly resemble the native environment and improve functional characteristics. 3D culture systems reflecting the native spatial organization including the crypt-villi structure or the stem cell niche are a prerequisite to enable spatial distribution of cell types. Application of Gradients of bioactive factors or oxygen is further important to mimic signaling within artificial model systems. In parallel, the supply of cells in vitro with nutrients, oxygen, energy but also growth and differentiation factors could be improved by vascularization strategies. 3D, Three-dimensional.

Commonly used scaffolds comprise hydrogels build by collagen or Matrigel as well as electrospun scaffolds composed of synthetic substrates such as polylactic glycolic acid or biological components such as spin silk or SIS. Wang et al.

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reported in 2017 the development of a microengineered collagen-based scaffold that enables the formation of human small intestinal epithelium with key structural features such as a crypt-villus architecture and associated cell type compartmentalization [113]. This approach combines an engineered scaffold displaying a characteristic microstructure with biophysical cues and a chemical gradient to mimic the human small intestine in vitro. Recently, Kasendra et al. reported another stateof-the-art example of a microfluidic-based system that represents a human small intestine on-a-chip model [114]. Microfluidic systems often consist of mainly two compartments, which are tightly connected to each other and further associated with a pump system to apply mechanical forces such as shear stress to the in vitro
grown epithelium. The model developed by Kasendra et al. is based on human primary organoids cultured underflow and cyclic deformations together with a microvascular endothelial component on a porous synthetic membrane (PDMS coated with collagen I and/or Matrigel). The authors reported that the intestinal epithelium forms villi-like protrusions covered by a polarized layer of differentiated epithelial cell types. Further, this chip system allows sequential analyses of fluid-collected samples in terms of digestion, absorption or transportation studies and provides the opportunity to include immune cells or a microbial compartment. A slightly different, socalled droplet-based microfluidic system was reported by Pajoumshariati et al. in 2018 [115]. The specificity of this system lies in the possibility to separate the crypt cells from the Peyer’s patch immune cells, a feature that is given in vivo but was so far not reported for an artificial in vitro system. By encapsulating both cell types into distinct biomaterials, the authors were able to coculture crypt and Peyer’s patch immune cells within one microfluidic device. In summary, this droplet-based microfluidic system ensures the spatial separation of both cell compartments while simultaneously enabling cell growth, proliferation, and differentiation. In contrast to synthetic scaffolds, biological scaffolds hold great promise to allow tissue modeling in a physiological relevant fashion. Schweinlin et al. published an elegant study in this context, which describes the design of a novel human small intestinal in vitro model based on human enteroids cultured on top of a biological scaffold represented by the decellularized small intestinal epithelium (SIS, small intestinal scaffold) [45]. The scaffold comprises a conserved ECM structure and basal lamina contents [45,116] that resemble an in vivo
like 3D microenvironment (see also Fig. 11.2). Another group used decellularized intestinal tissue to fabricate polymeric structures that consist of the polymer Parylene C that was vaporized onto decellularized SIS scaffolds [117]. Subsequently, those biomimetic molds were used to generate PDMS platforms containing tissue-like structures that mimic the crypt-villi architecture in combination with features of a basement membrane. Although decellularization of organs is a time-consuming process that fails in terms of standardization and reproducibility, biological scaffolds still harbor great potential for small intestinal tissue engineering approaches. Of further importance is the integration of the in vivo
like cellular diversity into current models, to mimic a mucus layer covering the intestinal epithelium or to study host
microbial as well as host
pathogen interactions. In this context, a continuous and direct cross-talk between the intestinal epithelium, microbiota, immune

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cells, or intestinal metabolites is required [118]. Technically, cellular diversity is achieved by the combined culture of different cell types. Early studies addressing this issue used Caco-2/HT29-MTX cocultures to model an epithelial cell layer covered by a simple mucus coating. However, this system fails to resemble the diversity of an in vivo
like mucin expression profile and appropriate thickness of the mucin layer [69
76]. A more advanced 3D model with an improved formation of mucus was built by the combined culture of Caco-2/HT29-MTX cells together with stromal compartments such as fibroblasts or immunocytes in collagen-based Transwell systems [119]. Schimpel et al. cultured Caco-2 cells with goblet and M cells in a triple-culture system to generate efficient models for absorption or transport studies [120]. In addition, Leonard et al. described in 2010 the design of a 3D coculture model of the inflamed mucosa comprising enterocytes, monocytes and dendritic cells [121]. Susewind et al. established a similar Transwell-based model in 2016 that was used for toxicity tests of nanoparticles [122]. As the microbiome plays regulatory roles during development, maturation, and function of the small and large intestine, a detailed understanding of precise regulatory mechanisms as well as a profound knowledge on host
microbiome or host
pathogen interactions within the gut is of great interest. Ordinary attempts in designing suitable in vitro models in this context enabled the cocultivation of intestinal epithelial cells together with anaerobic bacteria only in the short-term, mainly due to bacterial overgrowth [123]. Recently, Calatayud et al. described the development of a host
microbiome model for long-term culture of enterocytes, goblet and immune cells exposed to a microbial environment [124]. Integration of the mucosal immune system into a human primary cell model of the small intestinal epithelium was for example reported by Noel et al. in 2017 to study gut physiology and host
pathogen interactions in vitro [125]. Further interesting models are based on a microfluidic gut-on-a-chip system where Caco-2 cells were cocultivated under aerobic conditions for more than 1 week with intestinal microbiota [126,127]. Tissue homeostasis and regeneration in vivo are regulated by spatial and temporal cascades of signaling factors with distinct functions (see also Fig. 11.1) [25,128]. Therefore defined biochemical gradients have to be considered in order to establish functional in vitro models [45,114,126]. So far, growth factors are added to the culture medium to mimic signaling pathways in vitro; however, all cells are exposed to a uniform concentration of growth factors [29,30]. An elegant technical solution to solve this problem would be a biofabricated functionalized gradient scaffold that enables the controlled release of certain growth factors when culturing cells in vitro. Recently, Liu et al. described such a device generated by 3D bioprinting; however, the suitability of such a spatiotemporal release scaffold for modeling the small intestine still needs to be proven [129]. Besides the biochemical gradient, there is also an oxygen gradient within the small intestine [130,131]. Due to its distinctive structure, the subepithelial mucosa is provided by the bloodstream with higher oxygen levels, whereas the gut lumen is less oxygenated to enable the growth and survival of anaerobic and aerobic microbes [16,130,131]. In terms of mimicking natural oxygen environments in vitro, Chen et al. described in 2015 a bioengineered 3D porous scaffold composed of silk protein to establish human-like

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intestinal models that reflect in vivo
like luminal oxygen levels [132]. To recapitulate the in vivo situation more closely, Marzorati et al. developed the HMI (Host
Microbiota Interaction) module that provides a physiologically relevant environment to culture intestinal cell types in combination with microbiota under microaerophilic conditions [130,133]. Recently, Shah et al. reported a more advanced and modular microfluidic system that enables the cocultivation of Caco-2 cells and obligate anaerobes under their specific oxygen needs within a perfusion microchannel system called HuMIX (human
microbial cross-talk) that represents the in vivo situation more closely [130]. In summary, the here mentioned approaches for studying influences of oxygen gradients in small intestinal in vitro models are the first steps toward studying host
microbiome interactions. Further challenges in designing representative in vitro models of the natural small intestine are presented by integrating a functional vascular system [134]. Kasper et al. studied the role of vasculature in IBD using a Transwell model of Caco-2 cells combined with the microvascular endothelial cell line ISO-HAS-1 [135]. Further, Bertassoni et al. published an elegant work in 2014 in which they generated an artificial vessel network in a 3D micromolding approach with the ability for endothelial repopulation [136]. By combining the vessel-like networks with cell containing hydrogels, functionality of these artificial vessels was demonstrated by improved mass transport as well as survival and differentiation of cells cultured within these hydrogel models [136]. In addition to these works, Kim and Kim used a collagen-based bioink in combination with bioprinting technology to develop 3D models of the intestine that not only include 3D architecture by the presence of villi structures but also incorporated blood capillaries [137].

11.6

Outlook

The ultimate goal of small intestinal tissue engineering is the establishment of a standardized in vitro model representing the full functional diversity of the human gut. Two promising approaches have emerged on the way to meet this ambitious goal. Organoid cultures highly represent the in vivo situation but their standardization is complicated. Contrary, engineered TE-based systems meet the requirement for a high reproducibility, but for now, fail to resemble full intestinal function. Combining the advantages of both culture systems represents one of the main challenges for the upcoming decades. Great varieties of applied TE-inspired systems are used in the field. However, common standards need to be defined and set to combine biological cell sources such as primary cell types or organoids with biomimetic scaffolds and evaluation methods. Besides the standardization of existing methods, there is still a need for the development of new approaches, which might overcome hurdles of the current platforms. Along with the standardization, efforts should focus on model complexity and functional diversity. In contrast to organoid cultures, engineered culture systems harbor the advantage of a high modulation capacity, which, in the long-term,

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makes them a superior candidate in terms of implementing new features and the establishment of an optimized in vitro model. Ideally, complexity of engineered models could be increased by stepwise integration of relevant factors, for instance, the microbiome, immune cells, vascular structures, biochemical gradients, and mechanical stimuli. In this regard, the evolution of current and new techniques in both, the engineering field (3D printing, development of new surface materials, etc.) and biology (microbial culture, SC proliferation, differentiation, etc.) will be of great support and will further result in an interdisciplinary exchange more precious than it currently is. Certainly, the establishment of a single platform, incorporating both, functional complexity and required reproducibility within a fully functional in vitro model of the small intestine is not an achievement of the near future. Until then, the use of the available platforms will greatly depend on distinct research questions. For instance, patient-derived organoid culture is predestined for modeling pathogenic mechanisms and disease progressions, whereas iPSC-derived organoids can become a powerful tool for investigating early intestinal development or genetic diseases. Due to their high standardization, microengineered models are ideal to investigate the influence of separate factors. Based on a fundamental platform, the addition of individual elements (microbiome, scaffold, etc.) would enable the design of modular Systems. Therefore choice and optimization of appropriate in vitro models require the evaluation of pros and cons depending on the distinct research background.

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From insulin replacement to bioengineered, encapsulated organoids

12

Elisa Maillard1 and Se´verine Sigrist2 1 Strasbourg University, Strasbourg, France, 2Defymed S.A.S., Strasbourg, France

12.1

Introduction

Glucose regulation is a complex process involving not only pancreas for hormones secretion, and liver, fat, and muscle for glucose storage but also the gastrointestinal tract and the brain for coordination [1]. The pancreas is the only organ able to produce insulin, a unique hypoglycemic hormone. When insulin production, secretion, or use are impaired, diabetes develops. To treat and limit the complications related to chronic glucose fluctuations, insulin replacement is often required and different treatments exist. The present chapter aims at explaining glucose regulation and the actors involved, the diabetes problem, and current treatment. We will then discuss the future for diabetic patients.

12.2

Pancreas

The pancreas (Fig. 12.1) is a glandular organ, which has two sections with specific functions. The pancreatic exocrine compartment comprises acinar cells and duct cells in charge of the production and transport of digestive fluid. The digestive fluid contains water, salts, bicarbonate, and several different digestive enzymes (including lipases, proteinases, and amylases) and is transported to the intestine through the duct system. This mixture breaks down fats, proteins, and carbohydrates to facilitate their absorption. The exocrine compartment represents more than 80% of the pancreas and as such is the main component. The endocrine tissue comprises pancreatic islets, also called islet of Langerhans or pancreatic islets that are scattered and embedded throughout the exocrine tissue. They represent only 1%
5% of the pancreas [2]. The role of this component, which will be detailed further, is the release hormones into the bloodstream, crucial for glycemia regulation [3,4].

12.3

Pancreatic islet

Pancreatic islets are complex mini-organs (Fig. 12.2) surrounded and held together by a specific extracellular matrix (ECM). Different collagen fibers are connected by Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00015-5 © 2020 Elsevier Ltd. All rights reserved.

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Figure 12.1 Pancreas.

Figure 12.2 Human islet structure. Mini-organ organization of the islets comprises α-cells in blue, β-cells in pink, δ-cells in pink, PP-cells in purple, ε-cells in orange, vessels in red, ECM in yellow. ECM, Extracellular matrix; PP, pancreatic polypeptide.

laminins, fibronectin, nidogens, and other proteins [5
7]. ECM plays not only a structural role but also a functional role by transmitting signals through the focal adhesion. Integrin is the protein responsible for the anchorage of cells and initiates crucial survival pathways [8]. ECM is also a reservoir of growth factors beneficial for islet survival and function [9,10]. It is the cement of the islets that holds together different types of cells that collaborate closely for the maintenance of glycemia homeostasis. Each cell type, specialized in the production of one specific hormone that has an impact on the neighboring cells’ capacity of secretion, will be detailed further on. They are able to detect slight glucose variations and respond appropriately by rapid secretion of hormones. In order to coordinate and transfer hormones into the bloodstream, numerous vessels, and nerves penetrate the inner core of those organoids and ensure optimal function. In addition to endothelial and nerve cells, islets contain less specific cells such as resident macrophages and fibroblasts, the latter of which are involved in the building of the ECM.

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Cells composing islet are called α, β, δ, ε, and ghrelin cells, responsible for the secretion of glucagon, insulin, somatostatin, pancreatic polypeptide (PP), and ghrelin, respectively. Somatostatin, PP, and ghrelin are less specific pancreatic hormones since they are also produced by other cells in other organs and have consequently different roles that will not be developed in this chapter.

12.3.1 Composition of pancreatic islets 12.3.1.1 α-Cells: glucagon production α-Cells represent around 30% of human islets and are responsible for glucagon production [11]. Glucagon is a hormone with a length of 29 amino acids. Glucagon gene is translated into a precursor called preproglucagon, which is cleaved in the endoplasmic reticulum (ER) to form the proglucagon peptide, and it undergoes a final maturation step providing the active form of glucagon stored in the vesicles [12,13]. When the level of blood glucose is low, glucagon is secreted to exert its action mainly on the liver. It antagonizes insulin by stimulating glycogenolysis, or the breakdown of glycogen (form of glucose storage), by activation of adenylyl cyclase and protein kinase A, which promote phosphorylation of many regulatory enzymes. It has the same action on the muscles and activates triglyceride lipase to form fatty acids and glycerol in adipose tissues [14].

12.3.1.2 β-Cells: insulin and amylin production β-Cells represent approximately 50% of human islet and are responsible for insulin production. Insulin is a peptide hormone with a length of 51 amino acids (molecular weight 5.8 kDa) [11] and is secreted in response to changes in plasma glucose concentration [15,16]. Insulin gene is translated into a precursor called preproinsulin and is cleaved into proinsulin when its signal peptide has been removed at the entrance of the ER. Proinsulin is a 9 kDa peptide containing insulin A and B chains joined by the C-peptide, which has 30
35 amino acids [17]. Within the ER and Golgi, proinsulin undergoes an excision releasing both the C-peptide and the mature form of insulin. After this final maturation step, both peptides are internalized into vesicles and secreted into the blood in a 1:1 molar ratio, which is why C-peptide is a useful marker for insulin secretion [14]. Amylin or islet amyloid polypeptide is also a hormone produced by β-cells and cosecreted with insulin. Amylin has a paracrine action on α-cells to inhibit glucagon secretion, stimulates the hypothalamus for triggering the satiety signal and stop food intake, and slows down the gastric emptying to delay glucose absorption [18].

12.3.1.3 δ-Cells: somatostatin δ-Cells represent around 10% of human islet and are responsible for somatostatin production and secretion [11]. It is a small cyclic protein that inhibits hormone release, negatively regulating α- and β-cells. Somatostatin binds to high-affinity G-protein-coupled receptors [19] that inhibit cyclic adenosine monophosphate and

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downregulate insulin and glucagon secretion. It is expressed throughout the body and has different functions, such as affecting neurotransmission in the central nervous system and cell proliferation.

12.3.1.4 Pancreatic polypeptide cells: pancreatic polypeptide production PP cells represent less than 2% of human islet and are responsible for the production and secretion of PP [15]. This peptide is 36 amino acids long and acts as a hormone released postprandially and regulated via the vagus nerve. PP secretion is independent from glucose, has no action on insulin secretion, and inhibits glucagon secretion at low glucose concentrations via PPYR1 receptors on α-cells [16]. In addition, PP has a role in the reduction of gastric emptying and intestinal motility. PP could act as an intraislet regulator of secretion [15].

12.3.1.5 ε-Cells: ghrelin ε-Cells represent 1%
2% of a human islet and are responsible for the production and secretion of ghrelin [20]. It is a 28 amino acid hormone produced mainly by the stomach. Ghrelin is a central regulator of appetite and inhibits β-cells insulin secretion via paracrine interaction between δ- and β-cells. It also stimulates growth hormone secretion [21,22].

12.3.2 β-Cells role and insulin function In response to postprandial nutrient influx, glucose enters β-cells via glucose transporter GLUT1, and glucose uptake increases adenosine triphosphate (ATP) production. This impacts the closure of the ATP-dependent potassium channel at the membrane and triggers membrane depolarization. The voltage-dependent calcium channels open and calcium enters cells, which promotes the fusion of insulincontaining granules to the membrane inducing the increase of insulin secretion. Insulin secretion is under the control of calcium flux, glucagon, and somatostatin all of which can amplify or inhibit this process. Moreover, it has been shown that throughout the pancreas, the islets are working in synchronization under the direction of β-cell coordinator within each islet [23] to respond to glucose stimuli. Insulin is released directly into the portal vein by the pancreas, and through there it reaches the liver. Insulin binds to its receptors, is internalized and degraded, and regulates insulin release to the peripheral tissue. This is called the hepatic insulin clearance and can “filtrate” between 50% and 80% of the insulin produced [23
26] and avoid hyperinsulinism (Fig. 12.3). Once the first hepatic pass has been completed, clearance continues in the peripheral organs such as fat and muscle. Insulin binds to specific cell membrane tyrosine kinase receptors in the target cells [14]. This interaction promotes glucose storage and glycogen synthesis and inhibits gluconeogenesis and glucose production. In muscle and adipose tissue, insulin stimulates the translocation of the

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Figure 12.3 Glucose homeostasis. After a meal, glycemia rises activating β-cells. They secrete insulin, which inhibits α-cells and activates glucose storage in the liver, the muscle or the adipose tissue. Insulin secretion can be inhibited by δ- and ε-cells. Glucose is transformed and stored as glycogen, reducing the global glycemia. During the fasting period, glycemia decreases activating secretion of glucagon by α-cells. Glucagon secretion can be inhibited by β-, δ-, and PP-cells. Glucagon triggers glycogenolysis to increase the release of glucose and rises glycemia. PP, Pancreatic polypeptide.

glucose transporter GLUT4 to the surface of cell membrane to increase glucose uptake, glycogen synthesis, lipids biosynthesis, and lipolysis inhibition [14,27]. Furthermore, insulin release is constant and pulsatile, corresponding to the socalled basal insulin. In response to an increase in glucose blood level, the release of insulin is biphasic, with a large amount of insulin released within 10 minutes after glucose increase (corresponding to the release of stored insulin in vesicles), and a second phase with a decreased insulin release. This biphasic response has a function: the first phase is addressed mainly to the liver to stop glycogen production, whereas the second phase of release triggers glucose uptake by the muscle [28]. For some years, this process was considered independent, but currently, numerous factors, such as the circadian clock, have been described to intervene in glucose storage, insulin clearance, and insulin production. The complexity of glycemia regulation renders the possibility to mimic it very difficult [29].

12.4

Diabetes

The definition of diabetes by the World Health Organization (WHO) is, “diabetes is a chronic disease that occurs either when the pancreas does not produce enough

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insulin or when the body cannot effectively use the insulin it produces. [. . .] Hyperglycaemia, or raised blood sugar, is a common effect of uncontrolled diabetes and over time leads to serious damage to many of the body’s systems, especially the nerves and blood vessels”. Diabetes affects more than 60 million people in Europe (400 million globally), and the complications link to the chronicity of the disease such as kidney failure, peripheral vascular disease, stroke, and coronary artery disease are responsible for more than 600,000 death annually with a significant health and economic burden. WHO estimates that diabetes was the seventh leading cause of death in 2016 [4]. The WHO guidelines describe criteria for diagnosis of diabetes: “Diabetes symptoms (e.g. polyuria, polydipsia and unexplained weight loss for Type 1) plus: a random venous plasma glucose concentration $ 11.1 mmol/L or a fasting plasma glucose concentration $ 7.0 mmol/l (whole blood $ 6.1 mmol/l) or two hours plasma glucose concentration $ 11.1 mmol/l two hours after 75 g anhydrous glucose in an oral glucose tolerance test (OGTT). With no symptoms, diagnosis should not be based on a single glucose determination but requires confirmatory plasma venous determination. At least one additional glucose test result on another day with a value in the diabetic range is essential, either fasting, from a random sample or from the two hour post glucose load. If the fasting random values are not diagnostic the two hour value should be used.”

12.4.1 Type 1 diabetes Type 1 diabetes (T1D) can affect people of any age but mostly occurs in children and adolescents (,20 years). According to the International Diabetes Federation (Diabetes Atlas, 8th Ed.), more than 1.1 million people of this group developed T1D globally in 2017. More than one quarter (28.4%) of children and adolescents with T1D live in Europe (IDF, 8th Ed.), and within Europe, the highest rates of childhood diabetes are found in Scandinavia and north-west Europe [30]. T1D is an autoimmune disease in which the immune system specifically destroys β-cells resulting in insulin deficiency. The production of autoantibodies specific to β-cells solely is triggered by a combination of genetic susceptibility, autoimmunity, and environmental insults such as viral infection, toxins, or some dietary factors [31]. This reaction leads to the production of autoreactive T cells responsible for the destruction of the β-cells and the loss of insulin production. As a result, this type of diabetes requires exogenous insulin intake as soon as clinical manifestation of the disease (already around 80% of the cells are destroyed) [14,32].

12.4.2 Type 2 diabetes Type 2 diabetes (T2D) accounts for 90% of the diabetes population (more than 420 million in 2014; WHO), and it is most commonly seen in older adults. However, children, adolescents, and younger adults are increasingly developing T2D due to rising levels of obesity, physical inactivity, and poor diet (IDF, 8th Ed.) promoting the expression of genes involved in diabetes. Indeed, obesity is the major

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potentially modifiable risk factor for T2D, with an exponential relationship between body mass index and the risk of T2D [33]. There is a progressive impaired production of insulin due to loss of pancreatic islet mass and functionality along with the inability of the body to use the hormone [14,32]. In order to cope with this improper use of insulin, its production is increased and the pancreas is slowly depleted. In that case, insulin treatment is not required at the beginning of the disease, and an adjustment of the lifestyle can be usually proposed. In the case of inefficiency of diet and exercises, antidiabetic agents can be prescribed, which decrease glycemia by enhancing the elimination of glucose in urine (by inhibitor of the sodium
glucose cotransporter responsible of the recycling of the glucose in the kidney), limiting glucose absorption in the gut (α-glucosidase inhibitors), and by increasing β-cells sensitivity to glucose (via incretins). The aim is to achieve normoglycemia and relieve diabetes symptoms, such as thirst, polyuria, weight loss, and ketoacidosis in type 2 diabetic patients. The long-term goals are to prevent the development or slow the progression of chronic complications of the disease.

12.4.3 Gestational diabetes Gestational diabetes (GD) is temporary in pregnant women and usually occurs during the second and third trimesters. This kind of diabetes is diagnosed when the level of blood glucose is higher than normal but still below the cutoff value of diabetes (WHO). Nowadays, 1 out of 10 women develops GD, and about half of those women will be diagnosed with T2D within 5
10 years after childbirth (IDF, 8th Ed.). Women who are in this situation are at high risk of complication during their pregnancy, childbirth, and for their babies, though it is usually a transitionary disorder. Concerning treatment, around 70%
85% of cases can be controlled with lifestyle modifications alone [34].

12.4.4 Other diabetes Beside the more prevalent diabetes, there are several forms of diabetes such as maturity onset diabetes of the young (diabetes MODY), latent autoimmune diabetes of adulthood (diabetes LADA), double diabetes with the coexistence of type 1 and 2 diabetes [35], and steroid-induced diabetes. In 2017 it has been estimated that half of all people 20
79 years old live with undiagnosed diabetes (IDF, 8th Ed.).

12.4.5 Poor glycemia regulation complications When diabetes is not controlled, it can lead to the development of life-threatening health problems, decreasing the quality of life and, in extreme cases, leading to death.

12.4.5.1 Acute complications Hyperglycemia can negatively impact body metabolism by increasing lipolysis and, thus, producing a large quantity of ketone bodies, which in turn lead to metabolic

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acidosis [36], and to a serious disease called diabetic ketoacidosis. This a medical emergency since the blood acidification can impair the all body. The signs are usually related to frequent urination, extreme thirst, abdominal pain, nausea, confusion, fatigue, and many others, and the characteristic that should alert a patient is the very specific fruity-smelling breath.

12.4.5.2 Chronic complications Among other complications, people with diabetes are 2
3 times more likely to have cardiovascular diseases by increased the risk of blood clots, blood pressure, and cholesterol levels. Moreover, the risk of developing end-stage renal disease is 10 times higher (IDF, 8th Ed.). Diabetes can also cause retinal capillaries damage, which in turn may lead to loss of vision or blindness. Furthermore, nerve damage is another side effect of the disease, especially in the distal nerves of the limbs. “Diabetic foot” is one of the most common consequences of neuropathy resulting of the amputation of the feet (IDF, 8th Ed.) (cf. Chapter 17: Tissue engineering in urology).

12.5

Insulin replacement for type 1 diabetes

12.5.1 Glucose measurements 12.5.1.1 Capillary blood measurement First tests were developed in the 1960s, but the first easy-to-use home glucose meters have been available since the 1980s [37]. The glucometer usable at home has improved significantly the quality of life of patients with diabetes. It is a handheld device used for measuring glycemia thanks to an enzymatic reaction on singleuse test strips. A blood sample is harvested on the strip after prickling a finger. Glucose present in the blood will be oxidized by glucose oxidase situated on the strip and releases electrons. The produced electric current is proportional to glucose quantity and transmitted to the device, thus accurately measuring glucose concentration [38]. It is essential to control glycemia level during an intensive insulin therapy because it is highly variable throughout the day depending on the carbohydrate amount in meals, exercise practice, and global state [39]. Thus the checkup has to be done between 6 and 10 times a day, every day, in order to control the efficiency and permit adjustment to the treatment and to avoid hypoglycemia, which can quickly create discomfort and pain.

12.5.1.2 Interstitial blood measurements Fairly recently, a revolution was made with the measurement of glucose in interstitial tissue, which opened the possibility to continuous glucose measurement without further pain or discomfort. Indeed, it became possible to measure glucose transdermally, for instance, and constantly have a sensor at the back of the arm. There are three different monitors on the market, all composed of a sensor, a transmitter, and

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a display device. The sensor is also based on glucose oxidation and electron production creating an electric current by reaching the electrode. The intensity of the electric current is then conducted (by the transmitter) to the reader. Two of the systems complete a measurement every 5 minutes and transmit the data, the other one acquires data only when the flash reader is close to the sensor. The combined software is able to measure glycamia evolution during the day and can alert in case of hypoglycemia, especially at night [40]. It is possible to connect these continuous glucose monitors with insulin pumps to create an augmented insulin pump, as the starting point of the artificial pancreas [40,41].

12.5.2 Exogenous insulin As a result of the insulin production deficiency in type 1 diabetic patients, exogenous insulin is needed. The aim of the insulin treatment is to reproduce the scheme of islet secretion “basal bolus” by the pancreas. Insulin release is sequential and depends upon the quantity of glucose in the blood. β-cells detect glucose increase and release insulin in order to maintain its blood level in normal range. Therefore high amounts of insulin are usually released by the pancreas during a meal in order to store the excess of glucose for later use. In parallel, between meals, there is a release of glucose triggered by glucagon in order to supply cells with energy. This increase in blood glucose level is under the control of insulin, which is released also during those phases permitting again to prevent an increase in blood glucose. The scheme of insulin secretion by the islets, and more particularly by the β-cells, finely regulates glycemia (Fig. 12.4).

12.5.2.1 Multiple daily injections The administration of insulin usually involves subcutaneous injection with syringes, pen-like devices, or pumps. Subcutaneous injection is a safe, rapid, and easy method for insulin administration, but the fact that insulin is found in high quantity in the plasma triggers antibody production [42].

Figure 12.4 Basal bolus scheme mimicking the physiological secretion of insulin. In blue the physiological release of insulin during the day. In green the flat line reflects the long-acting insulin and the red arrows indicate the meals and the injections of the rapid insulin as a bolus.

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To mimic the physiological release of insulin, multiple insulin injections (intensive insulin therapy) are performed during the day with several types of insulin. Thus longacting insulin (delayed action) analog is injected in order to mimic the basal secretion of the pancreas, and the fast-acting one, also called the bolus injection, is injected at mealtime (acting in minutes). The difference resides in pharmacokinetic and pharmacodynamic properties depending on their chemical modifications.

12.5.2.2 Continuous subcutaneous delivery insulin Since the late 1970s, continuous subcutaneous insulin administration by pumps has been developed, being closer to physiological release [43,44]. The pump is a device containing insulin connected to the patient via a catheter in the abdomen subcutaneously. There is a constant release of low quantities of insulin for prolonged durations, and the bolus is ordered by the patient via the display in the pump. The pump must be refilled periodically, and the catheter changed to prevent obstruction by insulin crystallization. Continuous insulin delivery is better for patients, and improved glucose control and less hypoglycemia were observed [45,46]. Despite the progress made in the domain of pumps and insulin engineering, severe side events are still observed in some patients. The presence of hypoglycemia is still important [41,47]. Recent improvements permit to stop the pump when the glucometer detects a decrease in glycemia. Connectivity between glucose measurement devices and pumps allowed the development of algorithms that predict decreases in glycemia under a threshold, and in response, trigger the suspension of insulin delivery. The time interval today is around 30 minutes, between the detection/prediction and the suspension. Once the glycemia returns to an acceptable level, the release of insulin starts again [48]. This connection opens the field for autonomy, called closed loop or artificial pancreas.

12.5.2.3 Artificial pancreas In order to mimic even more the physiological release of insulin, an autonomous system called artificial pancreas is being developed. This system can either be composed of an external or implantable pump (intraperitoneal release and hepatic insulin clearance), an implantable glucometer, and a mathematical model of control algorithms hosted in a smartphone. Algorithms predict glucose excursions and adjust insulin delivery in accordance with the measurement of a real-time sensor glucose level. Thus it is important to include algorithms that can adapt to the blood glucose levels of each patient lifestyle [32,48]. The idea is that technology can replace human control to be more precise and more predictive, which could impact hypoglycemia and regulation. Hence, adaptable algorithms are tested in clinics with promising results [49]. Tremendous efforts and progress were made the last couple of years on these technologies; however, the safety of the system needs to be further improved, and the delay in response of the pump relative to glycemia needs to be refined. The system works better during the night than in the day because predications are more accurate since there is variability in lifestyle [48]. Intraperitoneal delivery of insulin is one of the

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Figure 12.5 Disadvantages of insulin injections from the patient’s point of view.

advantages of the implantable artificial pancreas with respect to hepatic insulin clearance. The intraperitoneal route may be beneficial; nonetheless, the pumps have to be placed surgically, which is more invasive than the external pumps [41]. Despite all these advanced technologies, fine physiological regulation of glycemia seems out of reach, and the patient is always facing disadvantages related to exogenous insulin administration (Fig. 12.5). Insulin is not the only actor in glycemia regulation; the brain and the gut, which release incretin, play an important role, and an automated system does not take those additional pathways into account. Thus replacement of cells is the best option.

12.5.3 Endogenous insulin production Because of the interconnectivity between all the protagonists in glycemia regulation, recreation of collaboration between artificial tools becomes difficult. Alternatively, cells able to release insulin according to the physiological pattern described earlier and sensitive to signals from other organs for synchronized action and optimal function can be used. This is the actual autonomous system. Restoration of in situ insulin production has been the focus of medical science for centuries. The initial attempts to restore endogenous insulin secretion were carried out in 1894. At that time a British physicist, Dr. Williams implemented the injection of a pancreatic cell preparation from a sheep pancreas. Unfortunately, the effectiveness of the treatment was not proven, and the patient died 2 days later. In fact, the first proof of concept appeared during the 1960s and the simultaneous transplants of a segment of the pancreas and one of the kidneys of a cadaverous donor in a patient with renal insufficiency. The 26-year-old recipient had been insulin-independent for 6 days posttransplantation, and graft rejection was observed 60 days after transplantation. This first transplantation demonstrated the technical feasibility of this approach.

12.5.3.1 Pancreas transplantation The technique was then modified by Lillehei [50], who carried out a complete pancreas transplant. The function of the transplant was maintained for up to 1 year. This technique was then transposed in many countries with varying degrees of success

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[51]. Since that time, pancreas transplantation has become the treatment that can treat T1D. For instance, more than 28,000 pancreas transplants have been performed in the United States [52] and increased by 7.0% in 2016 over the previous year [53]. It has been estimated that the patient survival rate is more than 95% at 1 year and 88% at 5 years [52]. However, given the risks of the procedure associated with heavy immunosuppression, this therapy is only recommended for patients on immunosuppressive therapy, transplanted with a kidney, or waiting for a kidney transplant for which a simultaneous kidney/pancreas transplant can be done [54]. Despite the effectiveness of this method to treat T1D, pancreas graft could lead to severe surgical complications.

12.5.3.2 Pancreatic islet transplantation Pancreatic islet transplantation is a highly successful treatment for reversing the life-threatening hypoglycemic unawareness in a subgroup of people with T1D mellitus. Although insulin independence is achieved in 50%
60% of transplant patients, resolution of hypoglycemic unawareness is achieved in almost 100% of patients [55,56]. Technically, pancreatic islets are extracted/isolated from the pancreas of brain-dead donors. The approaches currently used for islet isolation are all derived from the semiautomatic method developed in 1988 by Dr. Ricordi [57] consisting of the separation of the pancreas by enzymatic digestion using collagenase coupled to mechanical dissociation (Fig. 12.6). The technical difficulty lies in controlling the separation of the pancreatic tissue. After a first phase, called digestion, a suspension of the pancreatic tissue

Figure 12.6 Islet isolation and transplantation process.

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is obtained; the islets must then be separated and purified from the other cells. During a second phase, called purification, the pancreatic tissue is then subjected to density gradient centrifugation to separate exocrine and endocrine tissues (islet). At the end of this second phase, a preparation enriched in islets of variable purity is obtained. Islets can then be transplanted. Implantation of the islets takes place in the liver where they are slowly infused through the portal vein. They then follow the venous system to lodge on the periphery of the liver for implantation. This administration is performed under local anesthesia percutaneously by a radiologist. The first transplants of the human Langerhans islets date back to the early 1970s [58]. The first transplants yielded disappointing results. After the intervention, insulin-independence at 1 year was only 9%. The success rate was improved in the 2000s owing to the work by Edmonton team [59] and the application of a new immunosuppressive protocol. They showed that it was possible to achieve 80% insulin independence at 1 year in transplant patients. In this approach, called the Edmonton protocol [59], various major modifications were made to the original protocol: (1) the selection of patients with T1D eligible for islet transplantation only (patients without renal complication were selected), but suffering from unstable diabetes with severe hypoglycemia; (2) the injection of an increased islet mass by the use of 2
4 donors per recipient with an average of 10,000 IEQ/kg; and (3) the establishment of a steroid-free and nondiabetogenic immunosuppressive therapy characterized by induction with anti-IL-2 receptor antibodies followed by low doses of tacrolimus combined with sirolimus. After more than 15 years, islet transplantation has shown its feasibility and effectiveness and has been transposed in many centers worldwide [60]. The latest Clinical islet Transplantation Registry (CITR) report of December 2016 (https:// citregistry.org/system/files/9AR_Report.pdf) reveals 1927 infusions for 2421 donors and 1011 recipients. After the last infusion, 50% of the patients were insulin independent at 1 year. However, a decline in insulin independence was noted since only 20% of these patients were insulin independent at 5 years. The failure rate of the graft remains low (,4%), while the maintenance of graft function that results in a positive peptide C (C peptide .0.3 ng/mL) is approximately 80% despite a decline over time (45% to 5 years). Maintenance of normal fasting blood glucose was also observed in more than 50% of patients after 5 years, associated with a drastic reduction in severe hypoglycemia in more than 90% of transplant patients. Finally, insulin requirement was greatly reduced. The unpredictability of the islet isolation outcomes remains one of the main challenges preventing this treatment from being used more widely [61].

12.6

Islet transplantation limits (Fig. 12.5)

12.6.1 Low isolation yield and high pancreas requirement Despite effectively improving the glycemic control, a major drawback of the procedure is the significant number of pancreata required to obtain the 10,000 IEQ/kg

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required for insulin independence [55], along with the risk of patient sensitization due to multiple transplants. Furthermore, during organ retrieval, blood supply arrest triggers hypoxia and abrupt biochemical/metabolic changes. The ischemic insult continues during the preservation period prior to islet isolation. It is known that islet yield is inversely correlated with cold ischemia time duration [62,63].

12.6.2 Extracellular matrix destruction During the isolation procedure, ECM is digested by the collagenase used to separate islets from the exocrine. The interaction between ECM and integrin is disturbed which deactivates the survival signals and triggers apoptosis [64]. The islets are then cultured in a 2D environment from 24 to 48 hours and only 60%
80% of the islet mass survives and undergoes infusion in the portal vein [65,66]. Once islets are transplanted, the surrounding tissue needs to be remodeled for proper colonization of the islets in the liver. Among these environment modifications, vascularization is another limiting factor in this procedure. Indeed, despite the fact that liver is highly vascularized with a mean oxygen partial pressure of 20 mmHg, it takes around 14 days before having a functional vascularization for the graft [67]. During the first days, islets have access to oxygen only by diffusion that triggers hypoxia in the islet cores.

12.6.3 Hypoxia Physiologically, islets are highly vascularized, permitting maintenance of their high demand for oxygen and nutrients [68,69]. Therefore there is no need for islets to express high level of antioxidant enzymes. Consequently, unlike most cells that are able to use antioxidant enzymes, islets are highly sensitive to oxidative stress generated during hypoxia. Under partial hypoxia, cells rely on anaerobic glycolysis to generate ATP, which enhances the generation of reactive oxygen species (ROS) [70]. Moreover, ROS stabilize HIF-1α, which results in the activation of a transcriptional program with effects on metabolism, redox homeostasis, vascular remodeling, tumorigenesis, inflammation, and other processes that facilitate reoxygenation. However, in isolated islets, apoptosis is triggered rapidly since the process of vascularization cannot be enhanced. Moreover, generation of ROS will not only activate DNA strand breakage and peroxidation of proteins and lipids, but also a number of signaling pathways involved in inflammation and apoptosis [71,72].

12.6.4 Instant blood-mediated inflammatory reaction During the infusion, the contact between the islets and portal blood induces an inflammatory and injurious thrombotic reaction called “instant blood-mediated inflammatory reaction” (IBMIR). This phenomenon is characterized by rapid activation of coagulation and complement, the recruitment, and infiltration of the islets by leukocytes along with the adhesion and activation of platelets [73]. These reactions negatively affect the integrity and morphology of the islets leading to their

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death. It has been described that 40%
50% of the islet mass infused is destroyed by the reaction within the first few hours posttransplantation. Moreover, islets posttransplantation themselves secrete cytokines and chemokines involved in inflammation emphasizing that the reaction is already activated [74
76].

12.6.5 Autoimmunity and alloimmunity Posttransplantation, islets have to encounter two types of immune reactions— autoimmunity (T cell) and alloimmunity—involving conventional host antigraft immune response [77]. Autoimmunity is the reaction at the origin of the development of diabetes and the specific destruction of the β-cells [17,78,79]. Anti-β-cell antibodies could be reactivated after transplantation and can be correlated to an early loss of the graft [77]. Alloimmunity is another potent immune reaction due to the differences between the donor and the recipient regarding the specific identity of each [human leukocyte antigen molecules in humans and major histocompatibility complex (MHC)]. The response is categorized into three phases—in the first phase, the peptide-MHC complex on the antigen-presenting cells; the second comprises costimulatory molecules; and in the third, signals are triggered by cytokines, which will activate the T-cell proliferation. The ultimate biological effect is the recruitment of immune cells to the grafts with function loss [79
82].

12.6.6 Immune suppressive regimen Despite the advancement in the Edmonton protocol, it is known that some immunosuppression is deleterious for islet survival in the long term. Calcineurin inhibitor (tacrolimus), which is a potent inhibitor of allorejection and autoimmune recurrence, has been shown to decrease insulin secretion and be toxic for β-cells. Alternatives have been proposed, but they are globally less efficient than tacrolimus is [83]. Thus several donors are often required to reverse diabetes in a single recipient, limiting a larger application of the therapy.

12.7

Improvements in islet transplantation (Fig. 12.7)

Despite the dramatic decrease in islet mass during the complete process of islet transplantation, the positive results obtained in clinical settings encourage the research community to seek solutions for optimization and decreasing the islet damage. A global solution should be provided to put an actual step forward in terms of graft efficiency. Islet preservation within the pancreas during ischemia is the first target. Pancreas preservation is investigated. Despite the abundance of studies in the literature using different solutions [126
129], machines [130], or molecules [131
133], the gold standard remains the University of Wisconsin solution for decades.

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Figure 12.7 Islet transplantation limitations and solutions. The process of islet isolation decreases functional islet mass. Some of the main limitations and the solutions for clinical protocol as well as long-term solution for organ shortage and immune suppression are mentioned. Source: Adapted from Shapiro AM, Lakey JR, Ryan EA, Korbutt GS, Toth E, Warnock GL, et al. N Engl J Med 2000;343(4):230
8; Gamble A, Pepper AR, Bruni A, Shapiro AMJ. Islets 2018;10(2):80
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8; Komatsu H, Kandeel F, Mullen Y. Pancreas 2018;47 (5):533
43; Ramnath RD, Maillard E, Jones K, Bateman PA, Hughes SS, Gralla J, et al. Cell Transplant 2015;24(12):2505
12.; Vivot K, Benahmed MA, Seyfritz E, Bietiger W, Elbayed K, Ruhland E, et al. Int J Biol Sci 2016;12(10):1168
80; Vivot K, Langlois A, Bietiger W, Dal S, Seyfritz E, Pinget M, et al. PLoS One 2014;9(10):e107656; Li X, Meng Q, Zhang L. J Immunol Res 2018;2018:2424586; Couri CEB, Malmegrim KCR, Oliveira MC. Front Immunol 2018;9:1086; Dadheech N, James Shapiro AM. Adv Exp Med Biol 2018; Rodriguez-Brotons A, Bietiger W, Peronet C, Langlois A, Magisson J, Mura C, et al. Tissue Eng, A 2016;22(23
24):1327
36; Kim HI, Yu JE, Park CG, Kim SJ. J Korean Med Sci 2010;25(2):203
10; Schaschkow A, Sigrist S, Mura C, Dissaux C, Bouzakri K, Lejay A, et al. Cell Transplant 2018;27(8):1289
93; Berman DM, Molano RD, Fotino C, Ulissi U, Gimeno J, Mendez AJ, et al. Diabetes 2016;65(5):1350
61; Bertuzzi F, Colussi G, Lauterio A, De Carlis L. Eur Rev Med Pharmacol Sci 2018;22(6):1731
6; Boettler T, Schneider D, Cheng Y, Kadoya K, Brandon EP, Martinson L, et al. Cell Transplant 2016;25(3):609
14; Bottino R, Knoll MF, Knoll CA, Bertera S, Trucco MM. Front Med (Lausanne) 2018;5:202; Buitinga M, Janeczek Portalska K, Cornelissen DJ, Plass J, Hanegraaf M, Carlotti F, et al. Tissue Eng, A 2016;22(3
4):375
85; Calafiore R, Basta G, Montanucci P. Methods Mol Biol 2017;1479:283
304; Carlsson PO, Espes D, Sedigh A, Rotem A, Zimerman B, Grinberg H, et al. Am J Transpl 2018;18(7):1735
44; Chhabra P, Brayman KL. J Transplant 2011;2011; Cross SE, Hughes SJ, Partridge CJ, Clark A, Gray DW, Johnson PR. Transplantation 2008;86(7):907
11; Cross SE, Vaughan RH, Willcox AJ, McBride AJ, Abraham AA, Han B, et al. Am J Transpl 2017;17(2):451
61; Delaune V, Lacotte S, Gex Q, Slits F, Kahler-Quesada A, Lavallard V, et al. Transpl Int 2019;32(3):323
33; Denner J, Scobie L, Schuurman HJ. Xenotransplantation 2018;25(4):e12403; Gabr MM, Zakaria MM, Refaie AF, Ismail AM, Khater SM, Ashamallah SA, et al. Cell Transplant 2018;27 (Continued)

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L

Supplying oxygen during ischemia using the two-layer method has been challenged in clinics, but the results were disappointing (CITR report, [134]). Recently, an oxygen carrier, which is hemoglobin from marine worm called HEMO2life, has been developed that seems promising as it can carry up to 156 oxygen molecules and possesses antioxidant properties. Efficacy of this agent has been demonstrated against ischemia in the lung, kidney, or heart for transplantation in preclinical and clinical models [135,136], and on islets in culture [84]. The second target is the restoration of a suitable microenvironment promoting cell survival. Biomaterials are developed to mimic the natural environment of islets inside the pancreas and to possibly promote islet oxygenation. Different matrices, such as collagen [137] or fibrin [138], have shown improvement in islet insulin secretion and were assumed to reinforce islet structure enhancing the resistance

(6):937
47; Gołe˛biewska JE, Bachul PJ, Wang L-j, Matosz S, Basto L, Kijek MR, et al. Cell Transplantant 2019;28(2):185
94; Hogan AR, Doni M, Molano RD, Ribeiro MM, Szeto A, Cobianchi L, et al. Cell Transplant 2012;21(7):1349
60; Izadi Z, Hajizadeh-Saffar E, Hadjati J, Habibi-Anbouhi M, Ghanian MH, Sadeghi-Abandansari H, et al. Biomaterials 2018;182:191
201; Kanak MA, Takita M, Kunnathodi F, Lawrence MC, Levy MF, Naziruddin B. Int J Endocrinol 2014;2014:13; Kourtzelis I, Magnusson PU, Kotlabova K, Lambris JD, Chavakis T, editors. Regulation of instant blood mediated inflammatory reaction (IBMIR) in pancreatic islet xeno-transplantation: points for therapeutic interventions 2015; Cham: Springer International Publishing; Langlois A, Bietiger W, Seyfritz E, Maillard E, Vivot K, Peronet C, et al. Cell Transplant 2011;20(9):1333
42; Langlois A, Mura C, Bietiger W, Seyfritz E, Dollinger C, Peronet C, et al. PLoS One 2016;11(3):e0147068; Maillard E, Juszczak MT, Clark A, Hughes SJ, Gray DR, Johnson PR. Biomaterials 2011;32 (35):9282
9; Maillard E, Juszczak MT, Langlois A, Kleiss C, Sencier MC, Bietiger W, et al. Cell Transplant 2012;21(4):657
69; Maillard E, Sanchez-Dominguez M, Kleiss C, Langlois A, Sencier MC, Vodouhe C, et al. Transpl Proc 2008;40(2):372
4; Min BH, Shin JS, Kim JM, Kang SJ, Kim HJ, Yoon IH, et al. Xenotransplantation 2018;25(1); Mohseni Salehi Monfared SS, Larijani B, Abdollahi M. World J Gastroenterol 2009;15(10):1153
61; Montazeri L, Hojjati-Emami S, Bonakdar S, Tahamtani Y, Hajizadeh-Saffar E, NooriKeshtkar M, et al. Biomaterials 2016;89:157
65; Omori K, Kobayashi E, Rawson J, Takahashi M, Mullen Y. Cryobiology 2016;73(2):126
34; Pepper AR, Bruni A, Shapiro AMJ. Curr Opin Organ Transpl 2018;23(4):428
39; Pepper AR, Pawlick R, Gala-Lopez B, MacGillivary A, Mazzuca DM, White DJ, et al. Transplantation 2015;99(11):2294
300; Rajab A. Curr Diab Rep 2010;10(5):332
7; Rheinheimer J, Bauer AC, Silveiro SP, Estivalet AA, Boucas AP, Rosa AR, et al. Arch Endocrinol Metab 2015;59(2):161
70; RodriguezBrotons A, Bietiger W, Peronet C, Magisson J, Sookhareea C, Langlois A, et al. J Diabetes Res 2016;2016:3615286; Safley SA, Kenyon NS, Berman DM, Barber GF, Willman M, Duncanson S, et al. Xenotransplantation 2018;25(6):e12450; Samy KP, Butler JR, Li P, Cooper DKC, Ekser B. J Immunol Res 2017;2017:8415205; Sneddon JB, Tang Q, Stock P, Bluestone JA, Roy S, Desai T, et al. Cell Stem Cell 2018;22(6):810
23; Stokes RA, Cheng K, Lalwani A, Swarbrick MM, Thomas HE, Loudovaris T, et al. Diabetologia 2017;60 (10):1961
71; Tatum JA, Meneveau MO, Brayman KL. Diabetes Metab Syndr Obes 2017;10:73
8; Wang LJ, Kin T, O’Gorman D, Shapiro AMJ, Naziruddin B, Takita M, et al. Cell Transplant 2016;25(8):1515
23; Wojtusciszyn A, Branchereau J, Esposito L, Badet L, Buron F, Chetboun M, et al. Diabetes Metab 2018 [84
125].

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toward transplantation-inherent mechanical stress. In addition, an oxygen-producing matrix was shown to be beneficial for islet survival and function [139]. Fibrin has emerged as the ideal microenvironment for islets in vivo; this could be combined with an oxygen carrier to prevent hypoxia of cell inside the matrix [65]. However, fibrin participates in the IBMIR reaction, and the islet infusion after disruption of the fibrin matrix forms residues, which potentiate inflammation and graft destruction. Thus it is necessary to consider an alternative site outside the vascular system to host the graft to avoid IBMIR, with the opportunity to cotransplant the islets using biodegradable and biocompatible biomaterials. The omentum, which offers a large surface area [140] outside the vascular network, can accommodate a large islet volume, cotransplanted materials, or other cell types with islets, even at low purity [85]. Ongoing clinical trials are studying the omentum. The omentum is highly plastic and vascularized and has the capacity to generate new blood vessels, providing high oxygen availability for the graft [141]. Similar to that in the liver, the omentum permits insulin portal venous drainage [142], which is mandatory for proper glycemic control [143]. Thus a new implantation technique, hOMING (for h-Omental Matrix Islet filiNG), combining the environment, sites, and graft, was developed yielding encouraging results [86]. Finally, the omentum has been described as an immune-privileged site [144] that could permit to reduce immunosuppression. Taken together, improvement of islet isolation yield, islet insulin secretion, islet survival in vitro and in vivo could be the key to decrease donor requirement for diabetes reversion. However, donor shortage and immunosuppression remain to be addressed, and this would always be a problem for wider application. Alternatives for these two limitations represent the future of the cell therapy for patients with diabetes.

12.8

Other sources of insulin-secreting cells

The need for an inexhaustible source of insulin-secreting cells has been prompting the scientific community for several years with the ultimate goal to cure diabetes.

12.8.1 Cells of animal origin To compensate for the insufficient availability of islets, xenotransplantation may be an alternative to allotransplantation [144]. In this context, pigs seem to be the most appropriate candidates to provide the xenogeneic cells given their anatomical and physiological similarities to humans. However, several issues need to be addressed for their actual transfer to clinical settings.

12.8.1.1 Pig islet function Pig insulin differs in only one amino acid compared to the human insulin, and the insulin contents of porcine and human islets are comparable (616 and 750 IU/IEQ),

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respectively [145]. Physiological differences in terms of cell composition exist. For example, a pig islet contains 8% of α-cells as compared to 30% for a human islet. Moreover, for insulin secretion, the response of pig islets to glucose stimulation is four times lower than that of human islets [146], referring to the necessity to implant a larger number of pig islets for the production of a sufficient level of insulin to regulate blood glucose levels [146]. Several transgenic pig models have been developed to improve the insulin secretory function of the islets [147], such as those showing transgenic expression of a GLP-1 (glucagon-like peptide 1) resistant to DDPIV enzyme (dipeptidyl peptidase IV) combined with a constitutive activation of the muscarinic type 3 receptor which increases insulin secretion by a factor of 4 [147]. Islets from pigs of three donor ages—neonate, juvenile, or adult—were currently analyzed and compared [148]. In fact, membrane integrity, β-cell proliferation, and kinetics of insulin secretion were found to differ in an age-dependent manner. The highest proportion of compromised cells was found in neonate pig islets, reflecting a high degree of cell turnover [149
151]. β-cell proliferation was also higher in neonate and juvenile pig islets than in adult pig islets. However, insulin content and number of β-cells were higher in adult pig islets, associated with a greater secretion of insulin after a glucose stimulation test, than in neonate and juvenile pig islets [152]. On the other hand, in vitro maturation of juvenile or neonate pig islets could improve the glucose responsiveness [151] indicating that these young pig islets were more suitable in terms of viability and functionality. Numerous clinical trials have been performed with neonate pig islets but have failed to produce insulin independence in patients with diabetes [153,154]. To improve the efficiency of pig islets, a recent study demonstrated that a combination of mesenchymal stem cells with neonate pig islets could accelerate islet maturation and improve engraftment and function [155]. Porcine islets continue to be tested in clinical trials [156]. Defining an accurate and reproducible protocol is necessary to expand and mature the neonate or juvenile pig islets to further consider this option in the near future.

12.8.1.2 Risk of zoonosis Xenotransplantation of pig islets is still limited due to a risk of zoonosis (infections occurring due to agents that are naturally transmitted from animals to humans) [157]. Indeed, the pig genome has several endogenous active retroviral sequences (porcine endogenous retrovirus, PERV). Several studies have demonstrated the ability of PERV to infect different cell lines of human origin, with characteristics similar to those of retroviruses (gammaretrovirus) [98]. Until now, no porcine pathology has been associated with PERV, nevertheless, the expression of PERV particles has been observed notably in lymphomas [158,159]. These PERV sequences, therefore, theoretically induce a health risk of retroviral transmission of a porcine organ grafted into the recipient. In terms of biosafety, the production of animals free of all endogenous retroviral sequences (serum pathogen-free) could be the solution. However, despite the possibilities offered by enzymes, such as transposases or in vitro recombinases [160],

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this goal remains difficult to achieve in vivo and could induce a very high extra cost. However, some studies demonstrated the lack of virus detection during islet culture even if the animals were positive to this virus before islet culture [156,161,162]; these data suggest that the hygiene status of the herd may not reflect the status of the product. An easier alternative could be the selection of animals that do not transmit PERV. This capacity has been demonstrated in mini-pigs, some of which produce only PERV-C viral particles that are noninfectious to humans [163]. However, the risk of recombination, although low, persists in these animals [164].

12.8.1.3 Xenotransplantation and specific immune response A major difference exists in pigs in the synthesis of the carbohydrate galactosealpha-1,3-galactose (α-Gal) that was found in glycoproteins and glycolipids from all mammals except primates, including humans. Consequently, immediately after transplantation, pig islets elicit a hyperacute rejection induced by the α-Gal epitope [165]. Therefore transgenic pigs deficient in α-Gal have been created [166], although the effectiveness of the islets obtained from these pigs has not been demonstrated yet. One of the significant barriers to porcine islet transplantation is the early xeno-specific response [120,167]. Few studies have compared xeno-specific immune responses with allogenic responses [168
170]. However, a recent study [167] reported an increase in macrophage and antibody responses toward xenografts compared with those toward allografts, indicating that immune or genetic manipulations will be required to improve xeno-islet engraftment. Thus despite the numerous studies conducted to obtain humanized pig islets with increased functionality and resistance, biosafety and xenografts rejection continue to remain the major limitations to the clinical use of this cell source. Thus encapsulation seems to be the only method that can prevent rejection and ensure long-term survival of the transplanted islets without the risk of infection for the patient.

12.8.2 Surrogate cells 12.8.2.1 Embryonic stem cells Several other strategies and alternative cell sources have been developed to generate β-like cells, including direct differentiation from embryonic stem (ES) or induced pluripotent stem (iPS) cells. In recent years, there has been a drastic increase in the number of laboratories involved in stem cell research. ES cells are obtained at early stages of development from the internal cell mass of a blastocyst. These cells are pluripotent and could be differentiated into any embryonic cell type. They also have the ability to self-renew. Methods for establishing immortalized cell lines from mouse and human cells have been in existence since 1981 and 1998, respectively. The first trial for establishing insulin-producing cells from murine ES cells was conducted in 2000 [171]. These cells, grafted in a

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diabetic mouse, were able to improve their blood glucose. However, insulin production from these cells remained extremely low. Researchers then focused their studies on to obtain a definitive endoderm from ES cells. The first step toward this goal was taken by Kubo et al. [172], who developed a culture medium that induced endoderm production. D’Amour et al. [173], using this culture medium, developed a multistep protocol in which the first step involved enrichment of endodermic cells, followed by a second step that allowed the production of endocrine pancreatic hormone-producing cells, especially those producing insulin and C-peptide. One of the disadvantages of this method was the simultaneous secretion of several hormones by the resulting cells, which did not appear to be of any specialized type. Subsequent research efforts were focused on obtaining more “mature” β-cells, with a better controlled identity. Further progress in the production of insulin-secreting cells was enabled by the characterization of transcription factors involved in the development of the pancreas, particularly, the development and maturation of β-cells. Two of these transcription factors, pancreatic duodenal homeobox 1 (PDX1) [174] and neurogenin 3 (NGN3) [175], play key roles in β-cell maturation. Other signals also regulate the development of β-cells, such as epigenetic [176] and environmental factors (e.g., oxygen) [177]. Thus a method for in vitro generation of insulin-producing cells was established using endodermal cells that were engineered to express the cellular machinery that allowed insulin secretion following in vivo engraftment [122]. In 2014 Pagliuca et al. [178] demonstrated for the first time that it was possible to generate insulin-producing cells in vitro. They established a cell differentiation protocol that generated more than 100 million insulin-producing cells from ES cells. These cells expressed mature cell markers, responded to glucose stimulation, stored insulin within granules, and secreted insulin levels comparable to those secreted by human islets even after repeated glucose stimulation tests. Once transplanted, these cells were also able to improve the glycemic balance of diabetic mice. The proof of concept was, therefore, realized and then spread to several laboratories worldwide. The biggest challenge faced by stem cell-based insulin production techniques, which requires extremely expensive growth factors, is to increase the yield (scaleup) while limiting the cost of production and adhering to good manufacturing practices [179]. According to the stages of embryonic development, iPS cells can also find applications in regenerative medicine.

12.8.2.2 Induced pluripotent stem cells The procedure involves reprogramming specialized cells of an adult into stem cells. Maehr et al. [180] successfully reprogrammed fibroblast cells from the skin of a patient with diabetes to iPS cells. Takahashi and Yamanaka [181] demonstrated that expression of the genes OCT4, SOX2, c-MYC, and KLF4 can lead to the reprogramming of fibroblasts from an adult into iPS cells. Moreover, Tateishi et al. [182] successfully generated

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insulin-producing ILCs from the iPS cells from skin fibroblasts by using a serum-free in vitro protocol. This suggests that iPS cells have the potential to differentiate into both islet cells and areas of the inner lining of the pancreas, similar to human ES cells. These cells could avoid the use of an immune response, thereby solving ethical problems and offering a promising option for diabetes treatment. In 2012 Yamanaka and Gurdon were awarded the Nobel Prize for Medicine for discovering these specialized adult cells that can be reprogrammed into stem cells. The team used a combination of genes to reprogram differentiated cells first from mice, and then from humans 1 year later, into pluripotent stem cells [181,183]. The resulting iPS cells were then able to differentiate into all cell types from almost any organ. iPS cells are similar to ES cells and are inducible ex vivo. iPS cells can be reprogrammed by defined factors from somatic cells. Despite the growing interest of the scientific community in iPS cells, there are several issues that need to be addressed before considering their application in clinical settings. iPS cells retain the epigenetic memory of the specialized cells they are derived from [184] and in some cases undergo accelerated aging. Moreover, a major challenge is the risk of induction of mutations and activation of oncogenes during the reprogramming process, regardless of the method used [184]. Therapeutically, it is therefore crucial to determine the survival of cells used in vivo after transplantation, their level of insulin secretion in response to glucose, and especially the potential complications that could be induced by their transplantation. In the case of ES cells, the most severe complication would be the formation of teratomas and their possible malignancy [185
187]. In this context the concept of a bioartificial pancreas, which could extend cellular therapy to a large number of patients, could be a solution used in combination with these cells. Patients could then have access to an unlimited source of cells secreting insulin in response to glucose stimulation, along with a protection from these cells.

12.9

The bioartificial pancreas

12.9.1 Definition The strategy for developing a device called bioartificial pancreas, using immunoisolation or encapsulation, consists of placing insulin-secreting cells between semipermeable membranes composed of biocompatible and nondegradable inert materials. These materials must allow the passage of small molecules, such as insulin (  6 kDa) and glucose (180 Da), while preventing the entry of immune cells (  7 μm) and antibodies (  150
900 kDa). Thus the encapsulation of insulinsecreting cells would prevent the rejection of the graft in the absence of immunosuppressive regimen. The bioartificial pancreas must thus have three prerequisite functions: (1) protecting the transplanted cells from the recipient’s immune system, (2) protecting the recipient from the transplanted cells [stem cells, genetically modified cells (teratoma)], and (3) allowing rapid exchanges between the transplanted cells and circulating blood. The three fundamental factors in the development of a

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Figure 12.8 Bioartificial pancreas (BAP) properties.

bioartificial pancreas are the materials used for encapsulation, size of the capsule, and implantation site. Strategies for encapsulation of devices are classified into two broad categories: microencapsulation and macroencapsulation. The selected implantation sites are the peritoneum and under the skin (Fig. 12.8).

12.9.2 Microencapsulation Microencapsulation involves encapsulation of up to two islets in a capsule of less than 500 μm in size. Pancreatic islets are mostly immobilized in hydrogels, or in a threedimensional network of hydrophilic and cross-linked polymer chains. The most used and studied hydrogel for cell encapsulations is alginate [188], and the preferred implantation site is the peritoneal cavity [189]. The most commonly used encapsulation technique called drip approach was described for the first time in 1980 by Lim and Sun [190]. Drip approach involves rapid crosslinking of alginate in the presence of the islets in a solution of CaCl2 or BaCl2, to form the capsules containing the islets. Owing to the small size of the capsules, it is possible to optimize the diffusion phenomenon. Indeed, the distance between the external environment and the encapsulated islet is small [191]. The spherical shape limits immune responses because of the absence of a surface angle [192]. Because the total volume of the implant is small, the surgery is minor, and other implantation sites are easily accessible with capsules of 300 μm [193]. Other types of biocompatible hydrogels, including agarose or polyethylene glycol, have also been used for microencapsulation [194]. Coating materials such as poly-L-lysine for microcapsules have also been investigated to improve biocompatibility and mechanical stability. The benefits of microencapsulation could be attributed to the structure. Moreover, microcapsules have a high surface-to-volume ratio, allowing insulin secretion, cell respiration, and nutrient entry [195]. However, the major challenge is the biocompatibility of the encapsulation material, determining the performance of the microcapsules not only for immunoprotection but also for free fibrotic overgrowth, which in turn permits long-term islet cell survival [196].

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The first human clinical trial using encapsulated human islet cells was performed in 1994 by Soon-Shiong et al. [197]. The use of encapsulated human islets resulted in glycemic control and insulin independence for 9 months, but more than 15,000 IEQ/ kg was required to achieve this. Another clinical trial, performed in 2006 by Calafiore et al. [198], showed that a positive C-peptide was associated with the long-term reduction in the requirement for exogenous insulin without an immune response. Other clinical trials involved xenogeneic material using neonatal porcine islets [199,200]. Elliott et al. [200] reported a case of long-term survival (9.5 years) using microencapsulation of neonatal porcine islets using commercial microencapsulated islets from the Living Cell Technologies company. While no clinical trials involving porcine tissue showed sufficiently induced metabolic control, these initial clinical trials did demonstrate the feasibility of such an approach [201]. The capsules created have low mechanical resistance and stability. In fact, the rupture of a single capsule can lead to the loss of the entire graft due to the activation of the immune system. Recovering these capsules from all the cells in such cases is also impossible. Currently, the challenge is to find an encapsulation process that preserves the encapsulated islets while avoiding an immune response in the host. However, the implanted grafts failed and the reasons were as follows: a lack of biocompatibility of the capsules, which lead to chronic inflammation and then fibrosis; incomplete immunoprotection of the islets; and acute inflammation following surgical implantation (40% of the implanted islets died within the first few weeks of implantation, as the distance between the islets and the blood capillaries triggered hypoxia). The lack of reproducibility of results obtained using various encapsulation processes and between laboratories has been observed. Many laboratories are now seeking to develop standardized manufacturing processes for microencapsulation. To this end, for the last 20 years, microfluidics technology has been used to develop techniques to encapsulate cells.

12.9.3 Macroencapsulation Macroencapsulation consists of encapsulating all the islets in the graft in macroscopic pocket. This system is made of a polysaccharide (e.g., alginate or agarose) or a thermoplastic (e.g., polycarbonate or polyimide). The geometry of the chamber is variable, and it can be a cylindrical fiber or a sheet. The pouch can be intravascular or extravascular. However, given the high risk of thrombosis in intravascular systems, the extravascular system is preferred. The first macroencapsulation devices were developed in the 1950s. Although the safety of such an approach has been demonstrated, numerous studies carried out in animals have shown that insufficient diffusion of oxygen and nutrients caused a massive loss of cells inside the device. However, many areas of improvement could be identified, namely, the biocompatibility of the encapsulation membranes, the growth of host cells in the device, rejection time of encapsulated cells, and prevention of rejection. New tests were successfully performed in rats using the fiber technology used in renal dialysis, and these tests validated the use of the macroencapsulation model. However, the size of the device combined with the low cell density limits it from being considered as a treatment for diabetes.

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Therefore another approach involves the use of a flat sheet device with two membranes welded together to form a flat pocket. This configuration allows greater stability than the fibers and better oxygen diffusion. Baxter Healthcare was the first to develop such a device in the 1990s and published encouraging results of studies involving a rat model [202]. In their model the lack of oxygenation was partially solved by the induction of a capillary network around the device. However, despite the encouraging results obtained in a rat model, the results could not be confirmed in a model of a larger animal. Since then, many groups have worked to improve this module. Today, this module is called Encaptra (formerly known as Theracyte) and is developed by the company Viacyte. The results obtained on rodent, although encouraging, are limited by the size of the device, which allows implantation of only a small number of islets, and by the subcutaneous implantation site, which does not allow physiological restoration of insulin secretion [203,204]. This device was in clinical phase I/IIa at the end of 2014. In this trial, stem cells were differentiated into insulinsecreting cells once implanted in humans but with mixed clinical results. A new study is underway, but it will be necessary to implant between four and six devices in the patient to provide a sufficient number of islets to restore the insulin independence of patients. In addition, this study will be performed along with immunosuppressive therapy to validate both the safety of the device and the ability to secrete insulin based on blood glucose while preventing diverse immune responses. This new clinical approach leaves questions about the immunoisolation abilities of the device. Another macroencapsulation device called MailPan was launched in France by Defymed. It was the result of close collaboration between many laboratories for more than 20 years. The technology of the device is based on the same prerequisites as the ENCAPTRA device, using a polymer-based membrane that is flexible in terms of thickness, porosity, and density of pores to provide the required diffusion properties [205
209]. Furthermore, this device also has rigid support that can modulate the shape of the custom system and thus permit injection of a sufficient number of cells [210]. In addition, it is the only device that provides the ability to fill and empty the encapsulation module, as needed without any surgical intervention. The device can also be preimplanted, which allows prevascularization of the device before the injection of cell [208,209], thus leading to optimized oxygenation of these cells (Fig. 12.9). It has therefore been shown that it is possible to implant a sufficient number of pancreatic islets in such a device. The device could be emptied and filled by the introduction of septum, which makes it possible to replace the islets when they are no longer functional. The complete system has been validated in rats in terms of functionality and immunoprotection. The device is perfectly biocompatible. To evaluate the safety of the device, studies have been carried out in pigs after implantation of the device in the peritoneum. The team is now working to select the candidate stem cells that will be functional in the device to enable rapid clinical entry. Another macroencapsulation device is in its advanced developmental phase: the βairs device developed by Beta-O2s. In the various studies presented previously, a crucial point that has been highlighted is the importance of effective oxygenation of the islets, with hypoxia being one of the major limitations for cell survival in the

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Figure 12.9 Bioartificial pancreas implantation n in vivo.

device. Although vascularization can allow the formation of an oxygen gradient, in case the device is not preimplanted, the establishment of effective cell oxygenation in the device is crucial to preserve cell viability. The βairs device offers an alternative approach by delivering oxygen directly to the islets encapsulated in the device. The device is therefore composed of an encapsulation chamber where the islets are protected in alginate. Through a subcutaneous implantable chamber, oxygen can be injected into the device to oxygenate the islets and protect against cellular hypoxia [211]. The device is implanted in the peritoneum and has been demonstrated in rats and pigs that maintenance of normal blood glucose levels is possible by daily injection of oxygen into the module [211
214]. The first clinical studies carried out in 2015 with human pancreatic islets demonstrated that it was possible to maintain a functional transplant in the patient, in the absence of immunosuppressive treatment. However, the device does not contain a sufficient number of islets to achieve insulin independence in the patient. Although this device demonstrates the key role of oxygen in the survival of islets, the compliance of patients to use this therapeutic approach could be limited, since insulin injections are replaced with oxygen injections. Cell Pouch, the final microencapsulation device discussed here, has been recently developed by Sernova Corp. This device aims to induce the formation of a neovascularized chamber under the skin where pancreatic islets can subsequently be placed. This device is specifically designed to promote only vascularization around the subcutaneous pocket; however, an immunoprotective function could be added thereafter (either by the addition of Sertoli cells, or by the setting up of a device in the vascularized space). Studies in a murine model have already demonstrated the effectiveness of such an approach in optimizing the viability of the transplanted islets [115]. Clinical studies have also been performed, but its use in patients is not yet sanctioned. New clinical studies are currently underway. The major differences between the macroencapsulation devices that are today in the development phase are based on the following:

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Figure 12.10 Different strategies for diabetes treatment: from the present to the future. G

G

G

G

G

The implantation site. The peritoneum is the site that allows the most physiological delivery of insulin. The possibility of replacing the cells and not the device in all. The possibility of preimplanting the device before filling to optimize the exchange between the cells (oxygenation) and the recipient. The number of cells that can contain the device, this number must allow to restore insulin independence. Finally, the immunoprotection the final goal being to use this therapeutic approach without immunosuppressive regimen.

12.10

Conclusion

Despite considerable technological advances over the last 20 years, fine physiological regulation of glycemia using insulin administration seems out of reach, and the patient still faces disadvantages related to exogenous insulin administration. Pancreatic islet transplantation is currently the most promising therapeutic modality for treating patients with T1D. However, the number of pancreas available limits this approach. The considerable progress made in recent years in the generation of other sources of insulin-secreting cells (cells of animal origin or stem cells) could permit the use of this technique in a larger number of patients in the future. However, the risks that these cells present (e.g., biosecurity, infections, teratomas) prevent them from being considered for transplantation. The bioartificial pancreas provides new hope for treatment, by providing a secure cell therapy approach without the need for an immunosuppressant. Current and future clinical trials will help achieve this goal (Fig. 12.10).

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50. [214] Ludwig B, Rotem A, Schmid J, Weir GC, Colton CK, Brendel MD, et al. Proc Natl Acad Sci USA 2012;109(13):5022
7.

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Diabetic wound healing with engineered biomaterials

13

Laura E. Castellano, Jorge Delgado, Arturo Vega-Gonza´lez and Birzabith Mendoza-Novelo Department of Chemical, Electronics and Biomedical Engineering, Science and Engineering Division, University of Guanajuato at Leo´n, Guanajuato, Mexico

13.1

Introduction

The nonhealing ulcer in the foot of a person who suffers from diabetes mellitus is a multifactorial disease, which can lead to lower limb amputation. These factors include persistent inflammatory phase, poor epithelialization, reduced angiogenesis and blood flow, epigenetic changes, and misbalanced level of enzymatic activities. Diabetic foot ulcer (DFU) is defined as a foot affected by ulceration that is associated with neuropathy and/or peripheral arterial disease (PDA) of the lower limb in diabetic patients [1]. Five key fields of DFU management were delignated in 2015 by the International Working Group on the Diabetic Foot: prevention, footwear and offloading, peripheral artery disease or peripheral neuropathy, infection, and wound healing [2]. Over the years, it has been observed that evidence-based management of DFU can significantly reduce hospitalization, amputation, disability, mortality, and cost burdens. Standard therapies applied to treat peripheral neuropathy, infection, or wound healing, such as moist dressing, debridement, medicines, wound offloading, creams, sprays, and solutions such as platelet-rich plasma, have been ineffective [3]. Engineered biomaterials have emerged as key players in the restoration of molecular events or factors, which are generally synchronized in normal wound healing, but disturbed or delayed in DFU. In addition, new therapies supported by biomaterials are expected to improve wound recovery without a longterm hospital stay or high costs and to be easily applied by nurses and physicians.

13.2

Impaired wound healing under condition of diabetes

13.2.1 Diabetic foot ulcer complications The principal risk factors associated with foot ulcers are peripheral neuropathy, foot deformity, and PDA [4,5]. The most predominant diabetic neuropathy is distal symmetric polyneuropathy (DSPN) [6]. DSPN is characterized by altered sensation, pain, weakness, or symptomatic autonomic dysfunction. Motor neuropathy leads to foot deformities, and sensory neuropathy generates the absence of sensation, Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00005-2 © 2020 Elsevier Ltd. All rights reserved.

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resulting in constant abnormal pressure on the foot. Due to recurring pressure, increasing keratinization turns into callus formation, which induces ulcers [4,7]. PDA is a well-known manifestation of atherosclerosis that is associated with the formation of foot ulcer in almost 50% of diabetic people [4,8]. The impairment of the microcirculation is also observed in diabetes, which includes reduction of capillary size, thickening of basement membrane, and arteriolar hyalinosis [4,9]. The recurring hyperglycemia in diabetic people results in endothelial dysfunction that is implicated in the generation of microcirculation impairment [4,10]. These alterations in the diabetic microcirculation can explain the poor wound healing that is usually observed in diabetic patients.

13.2.2 Cellular and molecular events in diabetic foot ulcer The normal wound healing process is a well-orchestrated and overlapping of four phases (Fig. 13.1): 1. The hemostasis phase involves the accumulation of platelets that play an essential role in clot formation. This phase is mediated by cytokines and growth factors (GFs) such as platelet-derived GF (PDGF), transforming GF-β (TGF-β), vascular endothelial GF (VEGF), and fibroblast GFs (FGFs) [11]. 2. In the inflammatory phase the polymorphonuclear neutrophils (PMNs) are first to arrive, followed by the monocytes that migrate into the wound site and differentiate into macrophages, which play a crucial role in both inflammatory phase and tissue repair. Macrophages accomplish phagocytosis and the removal of cellular debris, bacteria, other foreign materials, and damaged tissue. In addition, these cells produce inflammatory cytokines and GFs such as tumoral necrosis factor alpha (TNF-α), interleukin (IL)-1, IL-6, and bFGF [11]. 3. Proliferative or tissue formation phases in which inflammation resolves include angiogenesis, the migration of fibroblasts, endothelial cells (ECs), and keratinocytes [11]. Fibroblasts produce new extracellular matrix (ECM) molecules such as collagen in the injury, which is increased by epidermal GF (EGF), keratinocyte GFs (KGFs), and TGF-β [12,13].

Figure 13.1 Outline of the wound healing process in nonpathological and diabetic conditions.

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4. The remodeling phase takes place over a much longer period of time, and the newly formed tissue is rearranged for higher mechanical strength [12]. During this phase, macrophages release matrix metalloproteinases (MMPs) in order to break down the temporary ECM and induce apoptosis of cells, generating mature skin to normal structure [14].

The failure of these phases to occur in a well-timed sequence promotes pathologic wound healing (Fig. 13.1). In chronic wounds the proliferative and remodeling phases do not occur and are characterized by dysfunctional cellular events and abnormal cytokine and GF function [15]. Wound tissue remains in the inflammatory phase, which does not promote tissue regeneration, and consequently, the wound could not heal [16]. For example, levels of neutrophils, serine elastase, and increase in proinflammatory macrophage activity result in the differential expression of GFs and elevated level of protease activity (i.e., MMP-2, -8, and -9), therefore extending inflammation process [11
13,17]. Neutrophils also release free oxygen radicals that induce significant oxidative stress, thus promoting chronic inflammation [18]. On the other hand, hypoxia-inducible factor (HIF) is a transcription factor that is a crucial mediator of the macrophage response to hypoxia [19], which induces the expression of inflammatory cytokines such as TNF-α, IL-6, IL-1β, in macrophages in a HIF-1α-dependent and -independent manner [20]. Macrophages participate in host defense, the initiation and resolution of inflammation, GF production, phagocytosis, cellular proliferation, and wound tissue healing. Macrophages have been classified into M1 phenotype for classically activated by interferon (IFN)- γ or microbial products, and M2 phenotype for alternatively activated by IL-4, IL-13, and IL-10. The M1 activation results in a highly proinflammatory macrophage phenotype, which is regulated by Toll-like receptor (TLR)4 ligands and IFN- γ and is characterized by Th1 (T helper type 1) response that is crucial in the elimination of pathogens. The M2 phenotype is characterized by Th2 (T helper type 2) response, necessary to promote wound healing and resolution of inflammation by releasing antiinflammatory cytokines such as IL-10 and GFs such as macrophage-derived GFs, PDGF, aFGF, bFGF, TGFα, and TGFβ [11,21,22]. The polarization of M1 to M2 phenotype is a very important process to prevent the delay of the wound healing, which is the outcome of chronic inflammation. Diabetic wounds, in both humans and mice, have shown that macrophages with an inflammatory phenotype that secreted IL-1β blocked the activation of the M2 macrophage phenotype [23]. Evidence suggests that IL-12 delays angiogenesis and the wound healing [24]. In diabetic wound healing, IL-1 and IL-12 play a negative role.

13.2.3 Implication of advanced glycation end products on cell function There are a variety of harmful substances that affect the molecular biology of diabetic wound healing, such as high glucose, high free lipids, and advanced glycation end products (AGEs). AGEs are proteins or lipids that become glycated as a result of prolonged exposure to sugar. These harmful compounds glycated proteins, such

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as collagen in skin, result in the formation of glycosylated collagen that functions distinctly compared to normal collagen in the skin. Long-term hyperglycemia can generate the production of a huge amount of AGEs, affecting the nonhealing or chronic wound in the skin and cell that participate in such process, for example neutrophils, fibroblasts, keratinocytes, and endothelium cell.

13.2.3.1 Neutrophils Neutrophils are cells that participate in the normal wound healing process, but abnormal neutrophils may contribute to the pathogenesis of nonhealing wound existing in diabetic people. AGEs can have a high affinity for human neutrophils AGE receptor (AGER), leading to increased intracellular calcium and actin polymerization, which affect cell migration [21]. It has been reported that neutrophils disseminated in skin tissue could be activated by AGE and then released H2O2, damaging normal tissue [25]. Tian et al. found that in in vitro and in vivo experiments, the exposure to AGE inhibited the viability of PMNs, promoted AGER production and cell apoptosis, and prevented the migration of PMNs. Also, AGEtreated neutrophils showed increased secretion of inflammatory cytokines and increased stress. These results suggest that an interaction between AGE and its receptors inhibits neutrophil viability and functions in the diabetic rat burn model [26]. Likewise, neutrophils from diabetic patients increase apoptosis, which is probably related to a high glucose environment. The decrease in neutrophil presence in the wound site leads to increased susceptibility to infections in diabetic patients [27].

13.2.3.2 Fibroblasts Fibroblast is a cell that produces the ECM and collagen, and it participates in the wound healing process. In in vitro studies, AGEs inhibit the proliferation of fibroblast and induce apoptosis [28]. Also, it has been reported that when fibroblasts are cultured in high glucose or AGEs’ media, they inhibited proliferation, decreased collagen synthesis, reduced synthesis of hyaluronic acid, present abnormal expression, or activity of proinflammatory cytokines or GFs and MMP-2, -3, -9, -13 [21]. In diabetic mice, fibroblasts present a severe impairment in VEGF production under normoxic and hypoxic condition [29].

13.2.3.3 Keratinocytes Keratinocytes are the predominant cells in the epidermis layer. The regulation of their proliferation and apoptosis are mediated by nuclear factor-κB (NF-κB), a transcription factor. In mice, NF-κB has been involved as a critical regulator of keratinocyte proliferation and differentiation. In in vitro studies the presence of AGEs diminishes the activity of the keratinocyte, inhibiting their proliferation, and it is also suggested that AGEs can prevent the cell transition from S to G2/M phase by activating NF-κB signaling pathway [21].

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13.2.3.4 Endothelial cells ECs, in vitro under high glucose or AGEs, increased apoptosis, upregulated secretion of adhesion molecules such as ICAM-1 (intercellular adhesion molecule1) and VCAM-1 (vascular cell adhesion protein-1), increased reactive oxygen species (ROS), reduced superoxide dismutase level, and activated cell-signaling pathway of mitogen-activated protein kinase and NF-κB [21,30]. High level of glucose can activate oxidative stress in ECs by AGEs [21].

13.2.4 Epigenetic changes related to diabetic foot ulcer Diabetes is a genetic disease; although the environment plays a key role in its development and evolution. The environment stimuli can cause changes in the gene expression without any change in the DNA sequence, this referred to as epigenetic modifications [31]. Epigenetic mechanisms may include DNA methylation, histone posttranslational modifications, and gene expression control through noncoding micro-RNAs (miRNAs) [32]. It has been reported that chronic diabetic complications such as hyperlipidemia and hyperglycemia can induce epigenetic changes that promote a proinflammatory macrophage phenotype [33]. The hyperglycemia is due to insulin resistance, which causes a sustained inflammatory condition that significantly reduces wound healing by macrophages in type 2 diabetic patients, resulting in foot ulcers [34,35]. Diabetes induces epigenetic modifications in macrophages, which are carried out by the following mechanisms: 1. Histone acetylation is catalyzed by histone acetyltransferases (HATs) and is associated with gene induction, histone deacetylation is catalyzed by histone deacetylases (HDACs) and is linked with transcriptional repression [33,36]. 2. Methylation and demethylations of histones are catalyzed by histone methyltransferases (HMTs) and histone demethylases (HDMs), respectively [33,37]. Histone methylation can induce both transcriptional activation and repression [33]. 3. DNA methylation at CpG nucleotides is associated with DNA methyltransferases (DNMTs) causing gene silencing. DNA demethylation is related to active chromatin [33]. Numerous studies suggest that hyperlipidemia induces a DNA methylation-dependent change to an inflammatory macrophage phenotype, through the suppression of antiinflammatory-related genes [33,38
40]. On the other hand, some reports imply that hyperglycemia promotes macrophage activation through cross talk between HMTs and HDMs in NF-κB signaling pathway (revised in Ref. [33]). 4. miRNA is a small no-coding RNA molecule of 21
25 nucleotides in length that negatively regulates gene expression posttranscriptionally. It usually binds to 30 -untranslated region of its mRNA target. It has been reported that miRNAs play a role in diabetes and its complications [41]. Madhyastha et al. reported differential expression of miRNAs related to cell development and differentiation, during wound healing in diabetic mice. This preliminary study implicates a significant role of miRNA in the pathogenesis of diabetic wounds [42]. In human diabetic foot, it was demonstrated that the expression profile of miR-203 has a positive correlation with the severity of DFU [43].

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In conclusion, hyperlipidemia and hyperglycemia promote inflammatory macrophage phonotype by epigenetic modifications, affecting the wound healing process of DFU. Considering this, macrophage epigenetics can be a target for DFU therapy.

13.3

Physicochemical aspects and fabrication of biomaterials in diabetic wound healing

Strategies related to moisture balance, protease sequestration, GF stimulation, antimicrobial activity, oxygen permeability, and capacity to promote autolytic debridement that facilitates the production of granulation tissue and the reepithelialization process are usually weighted in a biomaterial-based dressing [44,45]. The materials involved in the process of diabetic wound healing demand, in general, superior properties compared to those used in normal wounds. Broadly speaking, materials placed into the wound have to stay longer due to longer times needed to heal diabetic wounds [46]. This is also envisaged to avoid unnecessary manipulation of the wound.

13.3.1 Hydration The control of hydration via a material must be more accurate: an excessive amount of water can facilitate the proliferation of microbes, but the lack of it due to a longer time of exposition hinders a proper diffusion of necessary cells and biofactors needed to reconstruct the damaged tissue. Dehydration of the wound can be prevented using the formulation with hydroxy groups of the natural gellan gum and/or polyethylene glycol derivatives as cross-linkers [47,48]. Those gels show an extended swelling ratio and drug release. Another positive effect is an antioxidant effect and better mechanical properties. This kind of gels, with natural gums, have been also used to encapsulate cells more effectively, for instance, chondrocytes [49], providing pores that can be tuned to host the cells and make them more viable. Silk fibroin (SF)
poly(vinyl alcohol) mat dressing and electrospun SF mats have been highlighted as dressing that prevents water loss, have good oxygen permeability, and enhance fluid drainage [50,51].

13.3.2 Oxygenation Another consequence of the diabetic condition is the need for a biomaterial that is able to help neovascularization and to improve the number of cells involved in the wounded place. In such a sense, several more factors can play an important role: the amount of oxygen present in the tissue and the modulation of inflammatory cofactors. An interesting approach in wound healing is the use of oxygen via a microfluidic wound bandage [52], which marginally helps collagen maturation and wound closure. Nonetheless, the treatment is still considered under development, this study points out the benefits of proper oxygenation in the wound to obtain a more homogeneous and mature collagen matrix.

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13.3.3 Infection control The foot infection is a common complication in diabetes mellitus and the most recurrent cause of nontraumatic lower member amputation. The principal risk factor for the development of a diabetic foot infection (DFI) is due to the sustained foot wound. The worldwide incidence of DFI has been reported to range between 25.2% and 58%. Close to one in two diabetics with a DFU will acquire a DFI (revised in Ref. [53]). The control of DFI is essential to avoid lower limb amputation in the diabetic patients. Since most bacteria are resistant to antibiotics, this makes necessary the use of different antibacterial treatment to enhance the antibacterial
antifungal activity, for example, the use of metal nanoparticles (NPs). Silver (Ag) NPs are by far most commonly described in the literature because it is understood the mechanism that they use to prevent bacterial proliferation, and are cheap and easy to synthesize by green chemistry or in-situ methods. Derivatives of cellulose and natural organic acids with keto groups are used to reduce silver nitride with or without silver NPs seeds. As a result, a matrix that can form part of a hydrogel is obtained. It is reported that hydrogels with silver NPs reduce the time to heal diabetic wounds [54]. Other metals used include gold, copper, palladium, iron, titanium dioxide, and zinc oxide. [55]. Most of them can also be obtained from natural reducers and in such a way which is friendly with biomedical applications (Fig. 13.2).

13.3.4 Nanoparticle synthesis and its bioactivity A series of solid NPs commonly included in hydrogels are those with silica, providing the use of them to promote mainly angiogenesis. The poly(ε-caprolactone)/gelatin nanofibers dressing containing silicate-based bioceramic particles was produced by electrospinning [56]. Silicic acid has been proposed as the bioactive form of nanosilica that is capable of conferring positive charge to the NP surface to facilitate the internalization by cells [57]. Rapid angiogenesis, collagen deposition, and reepithelialization were stimulated by the silicon ions released from an electrospun membrane dressing composed of bioactive glasses and biodegradable polyesters in diabetic mice wounds [58]. Bioceramic particles were synthesized from tetraethyl orthosilicate, triethyl phosphate, and calcium nitrate by a sol
gel method. The chemical reactions on the surface of bioactive glasses stimulate the release of critical concentrations of ionic products, including calcium, silicon, and phosphorous, which have exhibited favorable intracellular and extracellular responses for promoting rapid angiogenesis [58]. As can be seen, the components from where the bioparticle is obtained are innocuous and the method is technically simple. Perhaps the most challenging aspect to attend the synthesis is the diverse shapes of the particles obtained by the synthesis that can affect its healing potential [59].

13.3.5 Cross-linking of biopolymers There are multiple kinds of materials used on diabetic wounds based on cellulose, chitosan, dextran, starch, alginate, elastin, etc. Mixtures of these raw materials are common, and chemical modification also. Most of them can form or react with

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Figure 13.2 Outline of the process of obtaining metallic NP intended to interact with cells. Reduction, nucleation, aggregation, and stabilization are key steps to adapt NP properties. NP, Nanoparticles.

aldehydes. Once they are formed or included, amines are also commonly used to cross-link the material [60]. Chemical bonds are then formed, and the resulting material has different, enhanced properties like a better uptake of water due to the inclusion of polar groups such as hydroxy, amine or acidic groups. The second important property is the enhancement of mechanical properties due to the precise

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cross-linking with the aim of increasing the elastic modulus at least one order of magnitude. Finally, using cross-linkers it is possible to modulate the porous size, which is a necessary condition to help cell proliferation. Among matrices to produce hydrogels, maybe the use of chitosan is the most manageable, since it can be easily linked via the free amine groups, which can also be used to change the pH-dependence of chitosan [61], making it more soluble in basic media inserting acidic groups and tuning its degradation time. Chitosan is a biopolymer obtained from deacetylation of chitin, which is the second most abundant polysaccharide in nature [62]. It can be found in the exoskeleton of crustaceans and insects, and in the cell wall of some fungi [63]. It is constituted by N-acetyl glucosamine and D-glucosamine units, linked by β-(1
4) glycosidic bonds. Chitin is not easy to dissolve due to its highly extended hydrogen-bonded semicrystalline structure [64], but once it is deacetylated, the structural conformation of the polymer is not preserved, it can be easily dissolved in acidic pH and human body can hydrolyze it as chitosan [65]. One of the most interesting properties of chitin/chitosan material for biomedical applications is attributed to the hemostatic behavior of the biopolymer because platelets are activated by chitin. This property can be preserved tuning the degree of deacetylation in the chitin/chitosan system: usually, commercial preparations are around 75% deacetylated. In addition, the activation of macrophages with chitin/chitosan induces the controlled production of VEGF [66]. Chitosan has been considered the main material for wound healing due to its biodegradability, biocompatibility, nontoxicity, antimicrobial, and biologically adhesive and hemostatic activity [67]. Chitosan scaffolds have been also produced using a 3D printer [68]. Those scaffolds have shown excellent properties in terms of biocompatibility, cytocompatibility, and toxicity toward two different skin-associated human cell lines, Nhdf and HaCaT, and have been used in the treatment of diabetic wounds, promoting the regeneration of a tissue with an improved functionality with respect to wounds treated with a commercial product. An alternative when collagen has to be used is the obtention of scaffolds via electrochemical deposition: [69] applying an electrical potential between two plates, a collagen dressing with excellent mechanical and biocompatible characteristics can be obtained. The dressing has been used in wounds, showing a significant increase in granulation tissue formation at 7, 14, and 21 days compared to untreated wounds. This approach can be certainly proposed to treat chronic wounds as suggested by the authors. Other strategies used to enhance hydrogel interaction in the diabetic wound healing include, for instance, the use of gels with L-arginine, which inhibits macrophage’s nitric oxide synthase: a polyester amide linked with arginine can be synthesized and evidence support faster wound diabetic closure and enhancement of angiogenesis [70].

13.4

Biomaterials supporting the administration of bioactive agents

The loading of therapeutics into biomaterials usually offers enhanced efficacy relative to the plain biomaterials to overcome the multiple aspects disrupted in the

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Figure 13.3 Outline of certain strategies based on biomaterials intended to interact with the cells in the wound bed. Material composition and microstructure, agents that prevent pathogen colonization, and vehicles that deliver genes are emphasized.

diabetic wound. The engineering of multifaceted bioactive dressings promises to induce synergistic effects on wound healing in diabetic conditions (Fig. 13.3). The different strategies based on the administration of GFs, drugs, natural extracts, genes, and cells, which use biomaterials with a role beyond as a simple vehicle, are discussed herein.

13.4.1 Therapy based on growth factors The local administration of GFs involved in wound healing is supported by biomaterials in order to improve dose- and time-dependent efficacy and costeffectiveness. Biomaterials are intended to prevent degradation of GFs from the proteolytic microenvironment of chronic wounds [71]. Table 13.1 shows the combinatorial therapy strategies based on GFs and biomaterials that evidenced the restoration of wound healing in diabetic rodents. Proteoliposomes, formed by the heparan sulfate proteoglycan syndecan-4 and liposomes, showed immunomodulation properties that improved the efficacy of delivered PDGF from alginate dressing [75].

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Table 13.1 Combinatorial growth factor (GF)-based therapy showing induction of wounds healing in experimental diabetic conditions. Growth factors

Adjuvants

aFGF Growth hormone PDGF EGF PDGF VEGF KGF EGF EGF/bFGF bFGF

Hyaluronate Proteoliposomes Lactate Cellular protective peptide AgNP LL-37 Chitosan/dextran

Dressing materials

Refs.

Cross-linked polyacrylic acid (carbomer) gel Methacrylate and vinylpyrrolidone copolymer films Collagen
PLGA nanofibrous mesh Chemically conjugated hyaluronate patch Alginate gel PLGA particles Elastin peptides in colloidal suspension

[72]

Polyurethane foam Silk fibroin
poly(vinyl alcohol) mat Reactive chitosan/dextran gel

[78] [50] [79]

[73] [74] [71] [75] [76] [77]

EGF, Epidermal growth factor; FGF, fibroblast growth factors; KGF, keratinocyte growth factors; PDGF, plateletderived growth factor; PLGA, poly(lactic-co-glycolic acid); VEGF, vascular endothelial growth factor.

Lactate, as a degradation product of the poly(lactic-co-glycolic acid) particles (polyester PLGA), was combined with the VEGF action to promote the healing of both nondiabetic and diabetic wounds [76]. Furthermore, the modification of VEGF with a sequence of collagen-binding domain favored the activity of this GF loaded into a collagen scaffold [80]. Other approaches include the use of cellular protective peptides that prevent apoptosis [77], AgNP, [78] or cathelicidin (LL-37) [50] with antibacterial activity in combination with the delivery of KGF, EGF, and bFGF, respectively. The chemical conjugation between EGF and hyaluronate film dressing seems to retain the activity of the GF [71]. The antifouling properties of materials such as thiolated chitosan/maleic acid dextran gel dressing have been harnessed to enhance the prohealing activity of bFGF [79]. The cross-linked polyacrylic acid (carbomer) gel dressing was proposed as a biostable system for 24 months capable of retaining the activity of FGF [72]. Despite bioactivity of GFs and adjuvant molecules/materials to manipulate host cells in the wound site, this therapy remains expensive and opens the doors to alternative bioactive agents, as discussed below.

13.4.2 Pharmacological treatment alternative to growth factors The manipulation of host cells for therapeutic purposes includes the delivery of small molecules as an alternative to GFs, as shown in Table 13.2. Immunomodulatory drugs that promote antiinflammatory macrophages or decrease proinflammatory macrophage phenotype have been loaded into biomaterials intended as delivery systems in the diabetic wound [81,82]. This includes the

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Table 13.2 Molecules as alternative to growth factors in induction of wounds healing in experimental diabetic conditions. Therapeutics

Agents

Dressing materials

Refs.

Immunomodulatory agents

MCP-1

Gelatin/poly(glycolic acid) electrospun mat Sprayable gelatin gel

[81]

Glutaraldehyde-cross-linked collagen foam Pyrrolidinone
chitosan foam Poly(ε-caprolactone) electrospun fibrous mesh Poly(L-lactic acid) electrospun fibrous mesh Lipid nanoparticles in chitosan/ collagen foam

[83]

Bioactive neuropeptide

Competitive inhibitors of enzymes

IL-8/MIP3a Neurotensin

DMOG

Simvastatin

[82]

[84] [85] [86] [87]

DMOG, Dimethyloxalylglycine; IL, interleukin; MIP-3a, macrophage inflammatory protein-3a; MCP-1, monocyte chemoattractant protein-1.

monocyte chemoattractant protein-1 [81], the chemotactic IL-8, and macrophage inflammatory protein-3a [82]. In this last case, gelatin was modified with hydroxyphenyl propionic acid groups in order to react in the presence of horseradish peroxidase and hydrogen peroxide. These systems were able to induce the recruitment of cells into the wound site and trigger the healing. In macrophages, HDAC inhibitors diminish the TLR-mediated inflammatory response [88]. HDAC inhibitor decreases macrophage IL-1β, IL-6, and TNF-α model of type 1 diabetes [33,89]. The design of selective inhibitor or activators of HATs, HDACs, HMTs, HDMs, and DNMT, with the modes of actions, described earlier, may be an appropriate therapy for the treatment of DFU. Also, the antihyperglycemic metformin showed an ability to attenuate macrophage-derived inflammation [33,90] and after eluding from biodegradable polyesters electrospun fibrous dressing effectively accelerated the wound healing [91]. The ability of the neurotensin to interact with immune cells was rationalized as a system that affects the early inflammation step on diabetic wounds [83,84]. This delivery system used a foam dressing based on chitosan derivatized with 5-methyl pyrrolidinone groups and collagen cross-linked with glutaraldehyde. The competitive inhibitor of prolyl hydrolases dimethyloxalylglycine (DMOG), which stabilizes HIF-1α and shows proangiogenic activity, has motivated its incorporation into the mat dressings [85,86]. The mesh morphology, DMOG, and even Si ions were associated with a synergistic effect on diabetic wound healing. The same idea has been assessed with the incorporation of simvastatin (SV) in nanostructured lipid carriers and then in a chitosan scaffold [87]. SV is an effective 3-hydroxy-3-methyl-glutaryl-coenzyme A reductase inhibitor is commonly prescribed to reduce plasma cholesterol levels and is also effective in wound healing because of its immunomodulatory, antiinflammatory, and proangiogenic effects.

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The selective inhibition of MMP-2 and MMP-9 by ND-336 in combination with the application of MMP-8 accelerated the wound healing by affecting inflammation, angiogenesis, and epithelialization [92].

13.4.3 Treatment with natural extracts The bioactivity of natural extracts has led to their incorporation into the dressing material. The antibacterial and antioxidant activities of flavonoids (polyketide derivatives synthesized exclusively in plants) have been related to the healing effect of dressings on diabetic wounds. Such is the case of apigenin (extracted from Morus alba) incorporated into gellan gum
chitosan gel dressing [47] or morin loaded into psyllium
keratin gel dressing [93]. Stilbenes are another plant-derived polyketide with antioxidant and prohealing activity in the diabetic wound. Resveratrol (RSV) is a polyphenolic antioxidant compound that was proposed for the treatment of diabetic wounds due to its unique properties to modulate tissue regeneration, microcirculation, function of peripheral nerves, production of cytokines, and insulin sensitivity. RSV loaded into microparticles comprised hyaluronic acid and dipalmitoylphosphatidylcholine showed a synergistic effect with collagen
laminin dermal matrix to promote healing [94]. The antiinflammatory and antioxidant activities of polyphenols such as curcumin (diferuloylmethane) have been tested with dressing biomaterials [95,96]. Gelatin microspheres were loaded with curcumin and then loaded into a thermosensitive pluronic F127 gel dressing. This dressing allowed the delivery of curcumin in a way that is responsive to the collagenase activity of the wound, while providing an efficient wound fluid absorption. The curcumin loaded into chitosan NP impregnated into collagen
alginate dressing were proposed as synergistic inductors of epithelialization and granulation tissue formation. Film dressing comprised Aloe vera (AV) extracts and SF induced the collagen synthesis and maturation and wound contraction under experimental diabetic condition by modifying the fibroblast behavior [97]. SF can support attachment, spreading and proliferation of epidermal cells and fibroblasts, while AV gel components exhibit antimicrobial and antiinflammatory (antibradykinin, an inflammatory mediator). Avena sativa extracts containing the alkaloid avenanthramides and the soluble fiber b-glucan were combined with keratin and konjac glucomannan (extracted from the plant Amorphophallus konjac) to produce an antibacterial, antioxidant, and prohealing gel dressing [98].

13.4.4 Gene therapy
based approaches Gene therapy represents an alternative approach to treat or prevent any diseases that cannot be addressed with conventional treatments. The gene delivery application can be made through biological, chemical, and physical approaches. In biological approaches, viruses are the most common gene vectors such as adenovirus serotype 5, adeno-associated virus type 2 and 5, Moloney murine leukemia virus
derived retrovirus, and human immunodeficiency virus-1
derived lentivirus [99]. Physical and chemical approach such as peptides, liposomes, nanomaterials, microneedles,

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electroporation, and iontophoresis have been employed for gene delivery as well (revised in Ref. [100]). These types of alternatives restore the cell function by the transiently silencing or activating genes that are important for wound healing [101,102]. These systems should generally be smaller than cells allowing them to be internalized. Delivery of Kelch-like ECH-associated protein 1 (Keap1) siRNA complexed in a liposome and supercharged coiled-coil protein hybrid NP delivery system (namely lipoproteoplex) showed to activate endogenous antioxidant mechanisms, normalizing the ROS imbalance and tissue regeneration in murine diabetic wounds with severe oxidative stress [103]. Keap1 is the repressor of a central regulator of antioxidant pathways, the transcription factor, and nuclear factor erythroid-2-like 2 (Nrf2). Nrf2 dissociated from its repressor, Keap1, is translocated from the cytoplasm into the nucleus and binds antioxidant response elements in the promoter region of a wide array of genes involved in protection against oxidative stress, protein stability, proteasome integrity, autophagy, senescence, and inflammation. Polyplexes of bFGF (pbFGF)-encoding plasmid with poly(ethylene imine) were incorporated into electrospun fibers with a core-sheath structure, and poly(ethylene glycol) was included into the fiber sheath to allow a sustained release of pbFGF for 4 weeks. The gradual pbFGF release revealed significantly higher wound recovery rate with improved vascularization, enhanced collagen deposition, and maturation, complete reepithelialization and formation of skin appendages [104]. A star-branched cationic polymer β-CD-(D3)7, which consists of β-cyclodextrin and poly(amidoamine) (PAMAM) acts as low-toxicity gene carrier into cells. The injection of β-CD-(D3)7/MMP-9 in a lesion in the diabetic rat skin improved the wound healing process. Thus β-CD-(D3)7 acts as a siRNA carrier for decreasing MMP-9 expression and improving wound healing in diabetic rats [105]. Exosomes are one of the most important secretory products of mesenchymal stem cells, mediating intercellular communication [106]. Exosomes are also considered as drug-delivery vehicles for the treatment of diseases and naturally occurring RNA carriers, which can even deliver therapeutic short interfering RNA to target cells. Exosomes overexpressing the miRNAs miR-126 with angiogenic activity were loaded into chitosan dressing. This composite promote wound surface reepithelialization, accelerate angiogenesis, and expedite collagen maturity by a synergistic effect [106].

13.4.5 Cell therapy
based approaches Cell-based therapy is a promising strategy for the management of DFU considering the paracrine mechanism to affect host cells. The transplantation of allogenic stem cells is considered a suitable option because of their capacity to differentiate into multiple cell lineages. To avoid the invasive collection of bone marrow cells, adipose-derived stem cells (ADSCs) are being assessed because they can be easily isolated from the adipose tissue and are capable of in vitro expansion [107]. The autologous transplantation of hypoxically preconditioned peripheral blood mononuclear cells and fibroblasts has also been evaluated [108]. Cells are expected to increase epithelialization and granulation tissue formation, antiinflammatory and antiapoptotic effects, and release of angiogenic cytokines [107]. This task has been

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addressed by cell sheet engineering. In this approach, native ECM secreted by the cells is preserved, and cells can be harvested as a contiguous cell sheet with intact cell-to-cell connections suitable for wound treatment [109]. Sheets of autologous fibroblasts and peripheral blood mononuclear cells supplemented with fibrin sealing glue increased the in vitro secretion of VEGF, hepatocyte GF, and TGF-β, as well as angiogenesis and fibroblast migration. In diabetic mice, refractory cutaneous ulcers healed more quickly after transplantation of fibrin-supported cell sheets than untreated ulcers or ulcers treated with only cell sheets [108]. Biomaterials are intended to provide favorable biomechanical support and biochemical environment for cells during transplantation [110]. Hydrogels, featured by a three-dimensional network swollen by the water, confer a template for easy manipulation and application of cells, while it is suitable for their survival. Table 13.3 shows the injectable hydrogel platforms acting as a cell delivery and retention system. The thermosensitive ethylene oxide/propylene oxide block copolymer (Pluronic F-127) gel-encapsulated ADSCs enhanced the expression of angiogenic and wound healing factors and promoted the cell proliferation in the wound site and significantly accelerated wound closure, which was accompanied by the facilitated formation of granulation tissue [111]. Synthetic hydrogel endowing pendent acrylate groups renders an injectable system with a rapid gelation rate, tunable mechanical properties, and nonswelling and antifouling properties, while ADSCs’ stemness and secretion abilities were maintained after encapsulation and diabetic wound healing process was accelerated [112]. ADSCs transplanted via hyaluronic acid
based hydrogels into diabetic mice with full thickness increased the number of intraepidermal nerve fibers, suggesting a beneficial effect in the treatment of peripheral neuropathy [113]. Table 13.3 Biomaterial technology supporting the cell-based therapy for wound healing in experimental diabetic conditions. Cells

Technology

Dressing materials

Refs.

Fibroblasts and peripheral blood mononuclear cells Umbilical cord perivascular stem cells Adipose-derived stem cells

Cell sheet

Fibrin

[109]

De/recellularized ECM

Dermal matrix

[110]

Hydrogel

Thermosensitive ethylene oxide/ propylene oxide block copolymer (Pluronic F-127) Hyperbranched multiacrylated poly(ethylene glycol) and thiolated hyaluronic acid Hyaluronic acid Alginate

[111]

Placenta-derived stem cells ECM, Extracellular matrix.

Hydrogel

[112]

[113] [114]

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Perivascular stem cells derived from the human umbilical cord recellularizing a dermal matrix enhanced wound closure, angiogenesis, reepithelization, and granulation tissue formation, with minor collagen deposition, suggesting a controlled scarring [110]. Placenta-derived stem cells suspended in alginate hydrogel and applied topically in a patient with DFU suggested the safety and clinical efficacy of biomaterials and stem cells therapy [114]. Certain limitations associated with the implementation of cell therapies have been identified due to high cost, need for specialist expertise to administer, specific storage requirements, and short shelf lives [115].

13.5

Biomaterials with prohealing activity

13.5.1 Natural extracellular matrix biomaterials The use of the living cellularized matrix is clinically limited as wound healing products because it is associated with high costs and potential immune reactions if the tissue is allogenic or xenogeneic. Thus biomaterials based on processed ECM have been developed to replace ECM molecules and critical GFs for the normal healing process in difficult-to-heal or chronic wounds [116]. The processing of mammalian tissues has emerged as a suitable method to render bioactive materials that owe their bioactivity to the maintenance of high concentrations of cell-derived cytokines and GFs. Soluble GF can be impregnated into the ECM materials to increase the bioactivity and induce the wound healing in the early stage of the wound in diabetic mice [117]. ECM scaffolds should be able to withstand the harsh proteolytic wound microenvironment and act as a temporary provisional matrix that allows infiltration of cells so that normal cellular activities can take place to promote matrix deposition, angiogenesis, and ultimately wound healing [118]. Using a decellularized dermal matrix derived from genetically engineered mouse (thrombospondin-2 knockout), it was demonstrated that structure and mechanics of the material are the keys to recruit cells, remodeling, and integration in both a subcutaneous implant and diabetic wound setting [119]. The application of a porcine small intestinal submucosa
derived matrix (in form of a membrane that is slightly larger than the wound, moistened with saline, and covered with a secondary dressing) in mixed arterial/venous ulcers resulted in a useful and well-tolerated treatment by patients, with a complete wound healing after 6 weeks [116]. Allogeneic micronized amniotic membrane, that is, amnion processed as received from donors in 300
600 μm microparticles accelerated wound healing in a diabetic mouse mainly by secreting factors related to growth, inflammation, and chemotaxis, that regulated macrophage migration and phenotype switch, recruited progenitor cells, and increased neovascularization. In this case, the micronized ECM material also served as a long-term dermal scaffold, whereas it was considered “living” due to the omission of tissue decellularization [120]. The reconstruction of a neuropathic ulcer in a patient with diabetic foot syndrome with a vascularized medial femoral condyle flap resulted in the wound healing, pain reduction, and improvement of gait. However, the use of autografts seemed to be limited due to the risks of infection at the donor site, the

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availability of tissue, and surgical costs. The solubilized ECM biomaterials applied in the gel state, either alone or combined with other biopolymers, have shown that nondiabetic wound healing did not exhibit scars [121]. This strategy is aimed at developing products that are easy to handle and place into wounds while taking advantage of the dual pH- and thermal-response of the collagen protein that composes the ECM materials. To the author’s knowledge, there are no studies that show that ECM hydrogels can be used to induce diabetic wound healing. The bioactivity of peptides isolated from natural sources has been tested in diabetic wound healing. The soft body of terrestrial snail (Cryptozona bistrialis) was processed to produce protein hydrolysates [122]. Keratin-derived powder was shown to accelerate the healing of diabetic wounds [123]. The biopolymers such as collagen, oxidized regenerated cellulose, bacterial cellulose, chitosan, or alginate in wound care products have shown the ability to form natural bonds with the surrounding tissue and therefore promote the healing process, exhibiting binding capacities for inflammatory mediators such as cytokines, proteases, and free radicals, which were found in high concentrations in chronic wounds [124].

13.5.2 Peptide-based biomaterials Peptide amphiphiles are self-assembled biodegradable materials that can be provided with multiple domains that are able to signal cells or to increase the therapeutic benefit of drugs (Fig. 13.4) [125
127]. Elastin-like peptides and heparin mimic peptide

Figure 13.4 Outline of biomaterials based on bioactive amphiphiles containing peptide building blocks that render a variety of nanostructures because intra- and intermolecular interactions. Self-assembled nanofibers or nanotubes produce gel networks that can be combined with other materials or therapeutic drug.

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were able to self-assemble into NPs (in combination with KGF) and nanofibers, respectively [128,129]. After administration in diabetic wounds, wound closure and granulation tissue formation were accelerated. Peptides with bioactive sequences such as an angiopoietin-1-derived or a laminin-derived have been combined with chitosan/ collagen gel dressing [130] or poly(ethylene glycol-co-citric acid-co-N-isopropylacrylamide) [131], respectively. A peptide with the amino acid sequence K2(SL)6K2 was proposed as a gel dressing that allows the infiltration of host cells in a subcutaneous in vivo model and the healing in diabetic wounds [132].

13.5.3 Inorganic agent
containing dressings Wound healing materials based on silicate and other inorganic materials have been shown to accelerate the wound healing process, improve the therapeutic efficacy, and prevent scarring [133]. Silicate materials can dissolve in a short period of time once they encounter the body fluid and absorb the surrounding biological molecules and therefore accelerate the penetration of fibrocytes into the wound site and the construction of collagen matrix [57]. The ionic dissolution products of bioactive glasses have been shown to activate fibroblasts through upregulated secretion of GFs such as bFGF, VEGF, and EGF and increase the production of collagen I and fibronectin as well as stimulate their migration [134]. Accordingly, this dressing material could act as a temporary scaffold that supported the proliferation of cells during the regeneration of soft tissue in the wound site [133]. In order to be in direct contact with the wound bed, inorganic agents are mainly used in the form of particles, ointments, gels, and membranes for wound healing. A similar idea was hypothesized using the silicate-based bioceramic particles (nagelschmidtite) embedded in conducive poly(ε-caprolactone)/gelatin nanofibrous composite scaffold [56]. Authors proposed that the acceleration of the diabetic wound healing by the application of nanofibrous composite scaffolds is related to the activation of epithelial to mesenchymal transition and endothelial to mesenchymal transition pathway. Zn21 and Ca21 ions released from silicate ceramic hardystonite embedded in the injectable alginate hydrogel were associated with the bacterial growth inhibition, stimulation of cell proliferation and migration, blood vessel, and epithelial formation during wound healing [135]. The effect of bioglass embedded in an injectable alginate hydrogel on the repair of rat diabetic chronic skin defects was complemented by using deferoxamine—a drug that promotes the secretion of HIF-1α—thereby upregulating the expression of angiogenic GFs and facilitating revascularization [136]. The combination of agarose and alginate resulted in the thermosensitive hydrogel that is cross-linked with the ions released from bioactive bioglass particles. The material was capable of stimulating the blood vessel and epithelium formation and the enhance healing of chronic wounds in a rabbit ear ischemic wound model [137]. Other inorganic compounds and ions used in skin wound healing as Ag1, B31, Ce41, Cu21, Fe31, Ga31, Ti41, Si41, and Zn21 [138,139]. Metal oxide NPs have been discovered to act as free radical scavengers because of their dual oxidation state, meaning that these NPs have oxygen vacancies with a role as direct antioxidants [139].

Diabetic wound healing with engineered biomaterials

13.6

353

Final remarks

Wound healing in the diabetic condition is featured by persistent inflammatory phase, poor epithelialization, reduced angiogenesis and blood flow, epigenetic changes, and misbalanced level of enzymatic activities, among others. Standard therapies poorly address these events or factors, and alternative therapies are required to reduce hospitalization, amputation, disability, mortality, and cost burdens in the management of DFU. Biomaterial technology offers a way to settle the hydration, oxygenation, or infection control requirements in the diabetic wound bed. The NPs synthesis and biopolymer cross-linking are procedures that allow the design of biomaterials with properties on demand. Microstructure and surface chemistry are fundamental material properties to balance moisture, sequestrate proteases, promote antimicrobial activity, permeate oxygen, and promote autolytic debridement. The bioactivity of biomaterials derived from the decellularized ECM, selfassembled peptides or inorganic materials—mainly silicates—in the diabetic wound healing has been related to the residual composition and preserved microstructure, cell-signaling domains, and dissolution products, respectively. The loading of therapeutics into biomaterials increases the efficacy relative to the plain biomaterials to restore the molecular/cellular events or factors disturbed or delayed in DFU. A synergistic effect on diabetic wound healing is induced by the administration of GFs, drugs, natural extracts, genes, or cells using biomaterials with a role beyond as a single vehicle. It is envisioned that engineered biomaterials will be the preferred option to improve diabetic wound recovery without prolonged hospitalization or high costs and be applied them easily by nurses and physicians.

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Bone morphogenetic protein
assisted bone regeneration and applications in biofabrication

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Naomi C. Paxton, Cynthia S. Wong, Mathilde R. Desselle, Mark C. Allenby and Maria A. Woodruff Institute of Health and Biomedical Innovation, Queensland University of Technology (QUT), Brisbane, QLD, Australia

14.1

Background

The underlying physiological processes for bone regeneration are well understood [1
3] involving intramembranous and endochondral ossification [4] (Fig. 14.1). For very small fractures, bone can naturally remodel using direct bone formation stimulated from the uninjured bone at either end of the defect without the presence of scar tissue. Ultimately, the function of the bone is rapidly restored so that the new tissue cannot be distinguished from the surrounding tissue. However, “critical-sized” defects, which tend to be large, are defined as injuries in which the bone does not have the capacity to regenerate over the lifespan of the patient or animal, and therefore surgical intervention is required to restore the function of the bone [5]. A variety of treatment options are well established for critical-sized defects as a result of significant injury, tumor resection or skeletal abnormalities in orthopedic, and maxillofacial surgery. However, these treatments remain a significant clinical challenge and are a major burden on global healthcare [6]. The current gold-standard treatment for large bone defects is grafting either through the use of harvested tissue from the patient (autografting) or another donor (allografting) surgically implanted into the defect site to assist bone healing [7]. Using natural bone as an implant to facilitate bone healing has a number of benefits. New bone and vasculature growth are readily assisted by the inflammation process in conjunction with the presence of native tissue, key growth factors, and other biological stimuli that are already present within the defect to signal bone remodeling. In addition, native bone provides physiologically relevant structural and mechanical support. However, donor demand significantly exceeds supply and a range of complications, including high rates of infection, severe pain, and refractures, limit treatment efficacy [8]. Bone morphogenetic proteins (BMPs) are growth factors within the transforming growth factor-ß (TGF-ß) family, first discovered in 1965 by Urist, an American orthopedic surgeon [9]. Also known as cytokines and metabologens, there are 25 Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00016-7 © 2020 Elsevier Ltd. All rights reserved.

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Figure 14.1 Schematic overview of the three primary stages of natural bone healing, inflammatory stage, endochondral bone formation, and coupled remodeling. Source: Reprinted by permission from Springer Nature Customer Service Centre GmbH: Springer Nature, Nature Reviews Rheumatology Einhorn TA, Gerstenfeld LC. Fracture healing: mechanisms and interventions. Nat Rev Rheumatol 2015;11:45
54. Available from: https://doi.org/10.1038/nrrheum.2014.164, ©(2015).

different BMPs known. Of these, several have been shown to play a significant role in bone formation and regeneration and have been actively used in tissue engineering and regenerative medicine (TERM) and biofabrication for bone healing [10]. Specifically, BMP-2 and 7 have been used in a wide range of bone condition treatments, including spinal fusion and nonunions, due to their demonstrated role in osteoblast differentiation [11]. BMPs have been utilized in TERM and biofabrication research to engineer functional biological substitutes that replace and restore tissues and organs. Biomaterials can provide a BMP carrier matrix for controlled delivery, essential to scaffold bone regrowth over weeks and months to bridge critical size defects. The use of resorbable, biodegradable materials comprises grafting substitutes that not only replace the missing tissue but also instigate bone regeneration and slowly undergo degradation resulting in complete resorption [12]. This facilitates partial or complete restoration of the tissue without the use of permanent implants [13]. Biomaterials and the use of BMP in TERM and biofabrication strategies for the treatment of bone loss will now be discussed.

14.2

Bone morphogenetic protein for bone regeneration

The process for bone healing in critical-sized defects comprises physiological mechanisms stimulating the recruitment of sufficient cells and nutrients to initiate

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bone regeneration and remodeling (Fig. 14.1). In long bones, for example, pooling of blood from exposed marrow and ruptured vasculature first creates a fracture hematoma, whose coagulation forms a fibrin network to provide a vital healing microenvironment consisting of a fibrous network to support cellular adherence and growth factor release. This environment assists with the migration of mesenchymal stem cells (MSCs), adhesion of platelets, and the release of growth factors including BMP that are crucial to the healing process, known as the inflammation stage where chemotactic signaling mechanisms attract the cells necessary to induce healing [14]. Next, the formation of a fibrocartilage callus begins through the differentiation of the MSCs into chondrocytes and infiltration of fibroblasts. Subsequently, endochondral ossification systematically converts soft callus to woven bone and blood vessel ingrowth. These deliver the required perivascular cells that instigate the formation of woven bone and resorption of the calcified cartilage. Finally, remodeled bone is formed through osteoblast/osteoclast activity and biomechanical loading [15]. Given the sensitivity of the healing process both spatially within the defect as well as the sequence of biological events, the delivery of BMPs in bone healing devices is critical to their efficacy. BMPs, suspended in a water-based solution such as phosphate buffered saline or cell media, have been directly injected into defect sites as a minimally invasive administration technique. However, rapid diffusion of BMP away from the defect area following a burst release after implantation significantly decreased the efficacy of the proteins during the healing process. Also, for the treatment of critical-sized bone defects, mechanical stability is crucial for the bone healing process to both stabilize the defect area and facilitate tissue regeneration. As such, BMP has been incorporated into a number of scaffold systems to localize and contain the delivery of BMP to the defect area whilst minimizing protein diffusion, to increase effectiveness [16]. Delivery systems have focused on two primary techniques: physical encapsulation and scaffold surface binding [17].

14.2.1 Bone morphogenetic protein delivery via carrier materials The motivation for encapsulating BMP in a physical carrying device or scaffold stems from BMP’s short half-life and lack of specificity to bone tissue [18]. As a result, BMP-containing devices have been required to deliver very high doses of BMP to a specific location to facilitate adequate bone regeneration [19]. There are four main categories of materials traditionally used in these systems, summarized in Fig. 14.2. In particular, synthetic polymers and metals have been fabricated into scaffolds for bone TE techniques [13,24]; however, their largely bioinert properties have been a limitation in optimizing cell attachment and integration. Therefore surface coatings have become a useful technique to modify both the geometry and biological properties of scaffolds to enhance cell attachment, proliferation, migration, and interaction between the host tissue and implant. Furthermore, surface binding of BMP to a scaffold localizes the distribution of growth factor within the defect to encourage bone formation at the host bone-implant junction.

Figure 14.2 Examples of ceramics, natural polymers, synthetic polymers, and composites used as BMP-carrier systems. (A) Porous β-TCP scaffolding for BMP-2 delivery. (B) BMP6-loaded alginate microspheres for incorporation in chitosan scaffolds for periodontal tissue engineering. (C) BMP-loaded PLGA microparticles electrosprayed into an electrospun PCL mesh. (D) 3D printed PCL
TCP composite scaffold into which rhBMP-7 is transferred into the center of the scaffold prior to implantation in a sheep tibial defect. BMPs, Bone morphogenetic proteins; PLGA, polylactic-co-glycolic acid; TCP, tricalcium phosphate. Source: (A) Reprinted by permission from John Wiley & Sons, Inc. from Sohier J, Daculsi G, Sourice S, de Groot K, Layrolle P. Porous beta tricalcium phosphate scaffolds used as a BMP-2 delivery system for bone tissue engineering. J Biomed Mater Res, A 2009;9999A: NA-NA. Available from: https://doi.org/10.1002/jbm.a.32467 [20]. (B) Reprinted by permission from Taylor & Francis Ltd, http://www.tandfonline.com from Soran Z, Aydın RST, Gu¨mu¨s¸derelio˘glu M. Chitosan scaffolds with BMP-6 loaded alginate microspheres for periodontal tissue engineering. J Microencapsul 2012;29:770
80. Available from: https://doi. org/10.3109/02652048.2012.686531 [21]. (C) Reprinted by permission from John Wiley & Sons, Inc. from Bock N, Woodruff MA, Steck R, Hutmacher DW, Farrugia BL, Dargaville TR. Composites for delivery of therapeutics: combining melt electrospun scaffolds with loaded electrosprayed microparticles. Macromol Biosci 2014:14;202
14. Available from: https://doi.org/10.1002/mabi.201300276 [22]. (D) Reprinted from Cipitria A, Reichert JC, Epari DR, Saifzadeh S, Berner A, Schell H, et al. Polycaprolactone scaffold and reduced rhBMP-7 dose for the regeneration of critical-sized defects in sheep tibiae. Biomaterials 2013:34;9960
8. Available from: https://doi.org/10.1016/j.biomaterials.2013.09.011 [23] with permission from Elsevier.

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Hydroxyapatite (HA) and other calcium-based materials mimic the natural composition of bone. HA is a calcium phosphate, comprises Ca10(PO4)6(OH)2, and can be processed into a range of forms including powders, blocks, and granules providing osteoconductivity with a high affinity with BMP [25]. BMP can either be adsorbed onto the surface of HA or covalent binding between HA, and BMP molecules can occur as a result of the calcium sites of HA binding to COO2, OH, and NH2 functional groups on BMP molecules [26]. Due to the intrinsic brittleness of HA as well as most calcium-based materials, additional materials are incorporated into composites to improve the mechanical properties of the BMP carrier to avoid repeat fractures. For example, Sawyer et al. [27] developed a composite polycaprolactone (PCL)/tricalcium phosphate (TCP) 3D-printed scaffold and investigated its potential for bone regeneration in a rat cranial defect model (Fig. 14.3). Bone formation volume was quantified using micro-computed tomography and visualized with histology, demonstrating significantly higher bone volume in the defects after 4 and 15 weeks with BMP incorporated into the scaffold system compared to empty and scaffold-only treatments [27].

Figure 14.3 Sawyer et al. [27] demonstrated improved bone regeneration in rat calvarial critical-sized defect using a 3D-printed PCL/TCP scaffold with rhBMP-2 compared to empty and scaffold-only experimental groups. (A) μCT images display a significant difference in regenerative capacity of the difference groups over the 15-week implantation which was (B) quantified using μCT bone volume analysis showing the efficacy of the rhBMP-2 treatment. (C) These results were also confirmed by histological analysis showing the infiltration of calcified bone tissue (stained black in MacNeal/von Kossa stain against blue ECM) within the entire defect area after 15 weeks of treatment with 5 g rhBMP-2. Goldner’s Trichrome indicates bone in greed and osteoid in red; Masson’s trichrome differentiates trabecular bone in blue and host cortical bone in red. μCT, Micro-computed tomography; TCP, tricalcium phosphate; ECM, extracellular matrix. Source: Reprinted from Sawyer AA, Song SJ, Susanto E, Chuan P, Lam CXF, Woodruff MA, et al. The stimulation of healing within a rat calvarial defect by mPCL-TCP/collagen scaffolds loaded with rhBMP-2. Biomaterials 2009;30:2479
88. Available from: https://doi. org/10.1016/j.biomaterials.2008.12.055 with permission from Elsevier.

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Hydrogels are a class of material characterized by high water-content polymer solutions cross-linked using biocompatible methods to form 3D polymer networks in which cells can be suspended. Natural polymers have been a focus due to their high biocompatibility. Specifically, alginate, gelatin, and hyaluronic acid have been selected as excellent polymer candidates for bone hydrogels and as BMP carriers. For example, hydrogels containing 5 wt.% hyaluronic acid cross-linked with polyethylene glycol tetra-thiols (PEH-SH4) have been loaded with human MSCs and BMP-2 and implanted into rat calvarial defects for 4 weeks. Histological analysis demonstrated significantly improved bone regeneration for BMP-containing hydrogels combined with MSCs and expression of vascular markers indicated mature bone formation [28]. Hydrogels have also been employed as slow-release carriers for BMP by employing tailorable polymer degradation to release suspended BMP in a controlled manner. Most notably, gelatin hydrogels have been developed for this purpose [29
31]. Naturally occurring extracellular matrix components, including glycoproteins such as collagen, fibronectin, and fibrinogen, as well as proteoglycans, are harnessed in TE strategies to deliver BMP to bone defect sites [32,33]. Of these, the structure and orientation of collagen plays a significant role in determining the bone’s mechanical properties [12,34,35]. Collagen can be processed in an aqueous phase, making it suitable for processing of soft materials including injectable materials, sponges, and sutures [36]. Of these, sponges initially received significant attention due to their natural bone biomimicry, biocompatibility, suitable surface properties for cell and molecule interactions, controlled degradation rates, and biocompatible end products [36]. Porous collagen sponges, made from animal- or human-derived collagen, can be fabricated via lyophilization and cross-linked via ultraviolet irradiation to create a porous structure that allows for cell infiltration [37]. For the loading of BMP into collagen sponges, several factors must be considered. For the formation of the sponge, its density and porosity impact the biomechanical efficacy of the device, including strength, stiffness, and other mechanical properties, while sterilization techniques must also be considered. BMP is then loaded onto the surface of the sponge with high surface-tovolume ratios via soaking. Time, BMP concentration, solute formulation, and the type of BMP used have also been identified as key factors that require significant optimization to improve implant efficacy [36]. In addition, collagen has been used in a wide range of coatings [38
40] and hydrogels [41,42] for TE applications. Despite long-standing preclinical success, relatively few BMP-carrier materials have successfully proceeded through clinical trials. Two successful products, INFUSE (Medtronic, Dublin, Ireland) and OP-1 Implant (originally Stryker, Kalamazoo MI, United States, also owned by Olympus Biotech, Hopkinton MA, United States), which proceeded rapidly through preclinical trials and successful clinical trials were commercialized in the early 2000s and feature the use of collagen sponges as BMP delivery vehicles to instigate rapid bone formation.

14.2.2 Clinical products Following the successful trials of BMP-loaded collagen sponges for use in spinal fusion in 2002, the Food and Drug Administration (FDA) and European Medicines

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Agency approved Medtronic’s INFUSE bone graft substitute for use in anterior lumbar spinal fusion procedures. Additional approvals were granted for the use of INFUSE to treat open tibial fractures as well as sinus and alveolar ridge augmentation in 2004 and 2007 [43] (Fig. 14.4). In addition, Stryker’s OP-1 putty was also approved for clinical use, using a similar system [45]. The INFUSE Bone Graft (Medtronic, MN, United States) was developed to be used with a titanium cage, known as the LT-Cage Device, for anterior lumbar interbody fusion. The cage is intended to provide structural support to maintain the space between degenerated spinal discs while the collagen and rhBMP-2 sponge facilitate bone fusion [46]. Seminal publications demonstrated the efficacy of the INFUSE products for a range of applications and sample sizes [47
49]. A landmark publication in 2002 reported the evaluation of 450 patients undergoing treatment for open tibial fractures, finding a significant decrease in risk of failure, requirement for secondary invasive interventions and faster fracture healing compared to control patients [48]. This study led to FDA approval of the INFUSE Bone Graft for this application. Stryker’s OP-1 Implant contains BMP-7 mixed with bovine bone collagen into an implantable device [50]. Initially given FDA approval in 2001 as a humanitarian device, whereby it may only be used where all other treatment options, including autograft, have been tested or are unavailable. Cahill et al. [45] reported that in 2002 only 0.69% of all spinal fusion procedures involved BMP-containing products. By 2006 this number increased to 24.3% of all primary spinal fusion procedures and 36.6% of revision surgeries [45]. This significant spike in the use of BMP products throughout the spinal care industry represents a significant peak in the use of BMP for bone regeneration treatments. Despite the widespread use and significant interest in BMP for clinical treatments, recent limitations have been reported leading to clinical approval being withdrawn for several BMP-containing treatments.

Figure 14.4 Preparation of INFUSE Bone Graft Large II Kit (8.0cc), as recommended by the Medtronic Instructions for Preparation and Handling [44]. Reconstituted rhBMP-2 is deposited by hand throughout the absorbable collagen sponge. Source: © 2018 Medtronic. All rights reserved. Reprinted with permission.

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14.3

Biomaterials for Organ and Tissue Regeneration

Bone morphogenetic protein limitations

In 2011 Carragee et al. published a vital review addressing the safety concerns raised in a number of FDA data summaries and analyzed several new insights into BMP-related complications reported in follow-up publications and other databases [51]. Key safety concerns such as implant collapse and infection rates were identified in the data reports that had been omitted from discussion in the original industry-sponsored publications, including the identification of a collapse of disc space by up to 50% and large osteolytic cystic lesions in the original radiographic data [52]. In fact, all 13 original publications relating to the clinical efficacy of INFUSE and AMPLIFY BMP products concluded that there were no adverse events that could be related to the use of the products. However, from 2006, a number of key publications reported significant adverse effects of BMP with complication rates between 20% and 70%. As a result, the FDA released a public health notification detailing the significant risks related to swelling of the neck, throat, and dysphagia. Furthermore, a number of lawsuits were filed against Medtronic by former employees for damages on behalf of the federal government [51]. Paper retractions, research misconduct, and investigations into inducements to the doctors reporting INFUSE’s success contributed to a tarnished perception of the efficacy of BMP in the spinal care industry [53
56]. The effectiveness and advantages of BMP were thoroughly reviewed in 2013 from clinical data across the Yale University Open Database Project, Medtronic reports, and FDA reports. The authors could not conclude any significant advantages of BMP in spinal fusion products compared to bone grafts and also highlighted significant disadvantages associated with BMP use, including potential carcinogenicity and ectopic bone formation, which had otherwise been excluded from previous reports [57,58]. These will be discussed in further detail in the subsections next.

14.3.1 On- and off-label use While the efficacy and complications regarding Medtronic’s BMP-containing spinal fusion products were thoroughly investigated over a half-decade, over 85% of procedures involving BMP-containing products were reported to be used off-label in 2010 [59]. Their off-label clinical use was likely promoted due to perceived onlabel success and further highlights the need for effective communication and reporting within the clinical sphere [60]. The reader is referred to a number of thorough articles that describe in detail and summarize the evidence for side effects as a result of BMP presence in bone regeneration products [45,51,61,62].

14.3.2 Complications and risks Increased rates of complications in BMP-containing procedures associated with the local defect site, including edema, infection, implant collapse, and other abnormal

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radiographic findings compared to controls, have been identified in a number of follow up studies [51]. For example, loss of stability and disc collapse were identified in 2008 in procedures using BMP-containing devices for spinal fusion by Smoljanovic and Pecina [63] based on the original publications supporting the clinical translation of INFUSE [47]. However, the findings were not identified by the authors at the time of publication in 2002. In addition, loss of alignment and high rates of infection have been reported based on retrospective clinical evidence, including up to five times the rate of deep wound infections for treatments using BMP-2 compared to those without for posterolateral spinal fusion procedures [51]. Uncontrolled ectopic bone formation is one of the most widely recognized complications associated with BMP use, characterized by the diffusion of BMP away from the defect site and subsequent ossification of tissues [64]. It is estimated that ectopic bone formation occurs six times more often in patients with BMPcontaining devices [65]. This is considered unsurprising by James et al. [65] due to the significant number of cells such as fibroblasts, adipocytes, and myoblasts that undergo osteogenesis following BMP-2 exposure. Ectopic bone formation as a result of spinal treatments can lead to significant neurological impairment via the compression of nerves, and challenging revision surgery is often required [62]. Nerve damage can also occur as a result of neurocompressive seromas and epidural hematomas, which poses a significant side effect of BMP [61]. Carcinogenicity has also been identified as a significant potential side effect of BMP products. However, the processes involved in BMP inducing cancer in a range of cell types have caused significant confusion in the underlying basic science. In a recent review, Schmidt-Bleek et al. identified this incongruity in the research surrounding the carcinogenic properties of BMP devices; studies have demonstrated both the reactivation and suppression of cancer cell growth as a result of BMP [10]. Notwithstanding, the FDA indicated that BMP-containing devices should not be used for patients with a history of cancer as a precautionary measure. The cancerinducing side effects of spinal treatments using BMP were brought to attention in an FDA report that reported incident rates of up to 5% compared to 1.8% for control treatments after 60 months [66]. The cause of such high rates of cancer has yet to be thoroughly investigated; however, the use of supraphysiological dosage has been widely proposed as the catalyst [10], which will be discussed in further detail next. The reader is further referred to a number of thorough articles that describe BMP-related side effects [45,51,61,62].

14.3.3 Implant considerations It is unclear in the literature as to why such high doses (.10 mg) are required in BMP products [67,68]. Factors such as diffusion of BMP away from the defect site and short biological half-life have been proposed as significant contributors to the requirement for supraphysiological doses, culminating in implants that do not adequately facilitate the natural biological role of BMP in the bone healing process [10]. Second, the efficacy of the devices themselves poses a significant concern for

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the safety of the treatment process and delivery of BMP to the defect site. The BMP-carrier sponge comprises bovine-derived collagen, however, since the collagen is derived from animals, the risk of immunogenic reactions and transmission of disease has been a source of concern [69]. In addition, the mechanical and degradation properties of the carrier devices must be adequately matched to the physiological environment in which they are implanted. Finally, the economic impact of bone regeneration devices containing BMP has been under significant investigation to determine the viability of such products in the spinal care landscape. Garrison et al. reviewed the surgical costs for patients undergoing spinal fusion and tibial fracture treatments using products with and without BMP in the United Kingdom and reported significantly higher costs for patients using BMP [70]. However, a number of studies have concluded that while BMP use may incur higher initial costs, long-term savings offset this cost increase through the reduction in follow-up surgical treatments, pain medication, and management of other complications associated with grafting and alternative treatment options [71
73]. These concerns have driven the development of the next generation of BMPcarrying devices that seek to develop more advanced bone graft substitute materials that regenerate bone in a more rapid, low risk, cost-effective manner in comparison to grafting and the existing products available, overcoming the challenges listed previously.

14.4

Current strategies

Recently, the academic and clinical communities have recognized the significant challenges still to overcome in bone regeneration through the large body of TERM and retrospective clinical research that has reported the complications of BMP use in bone regeneration devices. While reviewing the current trends in bone tissue engineering in 2014, Mravic et al. suggested that “a transition away from an interest in BMP-2 may be expected due to an increasing side effect profile” [74]. The next generation of bone regeneration devices will endeavor to better control the therapeutic and adverse properties of BMP. Three significant areas of TERM research will be discussed demonstrating the range of future strategies to overcome the concerns with BMPs using biofabrication techniques, converging TERM innovation with additive manufacturing technology. Lowering the dosages of BMP in bone implants has been highlighted as a crucial step in future research, in combination with finding more suitable delivery mechanisms to promote the efficacy of BMP within the healing process, localized to the defect site. Also, by controlling the release of BMP within the defect, the endogenous cascade of healing can be directly stimulated, enhancing bone regeneration. Finally, advanced biomanufacturing strategies offer enhanced biomimicry that could incorporate the minimized dose and controlled release systems or ultimately see a transition away from BMP altogether due to the ability to fabricate complex, hierarchical bone tissues using biofabrication strategies.

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14.4.1 Minimizing dose Bone tissue regeneration products have required up to a million times the natural concentration of BMP to justify efficacy, based on preclinical data from primates [75]. The FDA has approved 1.5 mg/mL as the recommended dose of BMP, and products such as INFUSE products contain up to 12 mg in a variety of sizes all at this concentration [44]. The scientific rationale for such high doses is well justified; compromised clinical outcomes have been widely reported for patients receiving lower doses of BMP to treat a range of bone defects compared to the clinical standard dose [48,76,77]. However, establishing a balance between efficacy using higher dosages and increased safety risks remains a significant challenge that bone TERM researchers continue to try to address [78,79]. In addition, as discussed in Section 3.3, the direct costs of BMP in bone regeneration devices are substantially higher than other treatments; lowering doses would provide a more cost-effective and accessible treatment option for patients. A number of recent studies have investigated the use of other growth factors in combination with lower doses of BMP to enhance the osteoinductive and angiogenic capacity of BMP (Fig. 14.5). Herberg et al. [81] have investigated the use of stromal cell
derived factor-1ß (SDF-1ß), which is known to be involved in bone formation, to regulate BMP signaling at lower doses than clinically used. The dosedependent relationship between the two growth factors was investigated in a critical-sized in vivo rat calvarial defect model. The results demonstrated that the addition of SDF-1ß to defects with insufficient BMP to induce significant bone formation alone displayed comparable osteoinduction to the treatments using BMP at clinically used amounts [81]. Similarly, Refaat et al. reported the successful use of cartilage oligomeric matrix protein with a low-dose BMP treatment to enhance

Figure 14.5 (A) μCT images of bone formation after silk fibroin/nano-hydroxyapatite scaffolds containing low doses of BMP-2 and VEGF were implanted for 8 and 12 weeks in rat calvarial defects. (B) The combination of both growth factors was found to induce significantly higher bone volumes within the defect after 12 weeks. μCT, Micro-computed tomography; BMP, bone morphogenetic protein; VEGF, vascular endothelial growth factor. Source: Reprinted with permission from the Royal Society of Chemistry from Wang Q, Zhang Y, Li B, Chen L. Controlled dual delivery of low doses of BMP-2 and VEGF in a silk fibroin
nanohydroxyapatite scaffold for vascularized bone regeneration. J Mater Chem B 2017;5:6963
72. Available from: https://doi.org/10.1039/C7TB00949F [80].

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bone formation in an 8-week rat spinal fusion model [82]. Charles et al. [83] also demonstrated the use of fibroblast growth factor 2 in combination with BMP in mice calvarial defects. In both studies, lower doses of BMP in combination with the chosen growth factors still produced similar bone regeneration compared to the high-dose BMP treatments. These studies demonstrate the efficacy of achieving adequate bone regeneration by lowering the dose of BMP required with the addition of other proteins [82,83]. Dosage customization and patient-specific therapies, a goal of biofabrication research, would ensure higher accuracy in delivery and dosage administration than would otherwise occur with off-the-shelf treatments [84]. However, a substantial amount of research would need to be devoted to understanding the effects of dosages for the treatment of a range of conditions as well as creating diagnostic tools to assess the required dosage necessary. In addition, strategies have been developed to improve the efficacy and controlled spatiotemporal delivery of BMP; these will be discussed in the next section.

14.4.2 Controlled release systems The controlled release of proteins was first demonstrated by Langer and Folkman in 1976 who developed a method of encapsulating high molecular weight proteins in biocompatible polymers that undergo controlled degradation to slowly release the therapeutic proteins [85]. Since then, a range of strategies have been employed in TERM devices and drug delivery to mitigate the impact of single, high-dose administration of high concentration proteins (Fig. 14.6). These strategies have been readily applied to BMP delivery for bone regeneration and control the delivery of BMP to the localized defect site. BMP diffusion into surrounding tissues is minimized by encapsulating BMP in hydrogel constructs or microparticles that both degrade over time, thereby releasing BMP into the local environment to stimulate the healing cascade rather than a burst release. Here, strategies in BMP encapsulation in microparticles, BMP encapsulation in biomaterial matrix, and surface coatings on biomaterial matrices will be discussed.

14.4.2.1 Microparticle bone morphogenetic protein encapsulation A considerable amount of literature has been published on the fabrication of microparticles encapsulating BMP for sustained and controlled release. A range of both synthetic and naturally derived polymers with biodegradable properties have been employed to provide a physical barrier between the BMP and defect area; as the polymer degrades, BMP is then released into the defect site. The rate of protein release can therefore be controlled through tailoring the encapsulation volume, shell thickness, and encapsulation efficiency, thereby allowing the rate and volume of BMP diffusing into the defect area throughout the healing process to be carefully controlled. Additives and composites are also widely used to enhance the surface and biological properties of the microparticles for more successful bone regeneration.

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Figure 14.6 The use of (A) microparticles, (B) hydrogels, and (C) surface coatings has been proposed to control and sustain the release of BMP (green) into the defect site. BMP particles in (B) not to scale. Bioprinted hydrogel scaffold image courtesy of Dr. Tomasz Ju¨ngst, Department for Functional Materials in Medicine and Dentistry (FMZ), University of Wu¨rzburg, Germany. BMP, Bone morphogenetic protein.

Natural polymers are readily adaptable, biocompatible, and able to be processed easily via a range of fabrication techniques to form microparticles and have been the focus of a number of studies on BMP encapsulation. Mantripragada et al. have reported the development of BMP encapsulated in chitosan microparticles with diameters of approximately 700 μm to treat rat femoral defects [86
88]. The microparticles were successfully fabricated and characterized before undergoing extensive in vitro and in vivo analysis, concluding that the presence of the microparticles contributed to improved bone remodeling, although they did not degrade over the 12week implantation [88]. The feasibility of loading and controlling the release of BMP within gelatin microparticles has been extensively reported with in vitro experiments validating techniques for sustained BMP release behavior [89
91] as well as demonstrated ectopic bone formation in a subcutaneous small animal model [92]. A number of gelatin-based microparticle systems have also outperformed microparticle delivery system-free controls in a number of rat defects [93,94]. In the field of synthetic polymers, polylactic-co-glycolic acid (PLGA) has been favored for encapsulating BMP in microparticles due to its low melting point and manipulability. Since PLGA is bioinert, PLGA microparticles are often administered in conjunction with an additional biomaterial or scaffold to support the defect site and facilitate osteoconduction while BMP is slowly released into the defect site stimulating osteoinduction [95] (Fig. 14.7). PLGA-coated BMP microparticles have been tested in combination with collagen
HA scaffolds [96], calcium phosphate scaffolds [97] as well as a range of hydrogels [98
101], which will be discussed in more depth in the next section. Finally, complex multimaterial microparticles have also been developed to control the release of both BMP and vascular endothelial growth factor (VEGF) to promote angiogenesis in parallel with bone formation [102].

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Figure 14.7 SEM images of PLGA microparticles loaded with BMP-7, fabricated via electrospraying, after 7, 14, and 21 days culture with MC3T3-E1 preosteoblast cells [95]. BMP, Bone morphogenetic protein; PLGA, polylactic-co-glycolic acid.

14.4.2.2 Matrix bone morphogenetic protein encapsulation Entrapment of BMP in biomaterial matrices such as hydrogels has also demonstrated significant efficacy through the slow and controlled release of proteins via hydrogel degradation. Furthermore, since the BMP is suspended in a mechanically robust hydrogel network, diffusion of BMP away from the implant is significantly minimized, improving the rate of protein uptake in the localized area and significantly reducing the volume of BMP required and lowering potential risks [103]. While historically a wide range of synthetic and natural polymers have been developed into hydrogels for drug delivery [104], for bone regeneration, mechanical support is a critical factor that limits the choice of hydrogel candidates suitable for this application. As such, advanced delivery systems have also been developed which often also include the use of a support scaffold or suitable rheological properties for injectable applications [105,106]. Krishnan et al. [107] created a hybrid system involving a PCL nanofiber mesh with an alginate hydrogel in which the BMP was suspended. After a 12-week implantation the authors analyzed the difference between clinically used products and their system with the same dose of BMP. While similar bone regeneration was observed, the BMP release in the alginate hydrogel was sustained over a significantly longer period, mitigating the significant complications associated with the initial high-dose burst of BMP in standard, uncontrolled systems [107]. Likewise, TCP scaffolds were used as a mechanical support structure for BMP-loaded gelatin and alginate hydrogels. The release profile was able to be controlled by altering the degree of alginate cross-linking, thereby BMP suspension in the hydrogel and lower doses of BMP were successfully used with reasonable efficacy in combination with the scaffold system for regenerating bone [108]. Shi et al. fabricated zirconia and HA to form

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porous scaffolds combined with a chitosan-based hydrogel to suspend BMP in a complex implant to regenerate bone in using a slow-release BMP system [109]. In addition, in the wider scope of TERM research, hydrogels have primarily been developed for the suspension of cells in high water-content cross-linked polymer networks. Therefore the interaction between the BMP and cells suspended in the hydrogel has been thoroughly investigated to determine the optimal conditions for rapid in vivo tissue development. Recently, Ho et al. reported the enhanced survival of MSCs suspended in an alginate hydrogel with BMP as well as 100% union of rat femoral defects in 12 weeks compared to BMP-free samples that did not bridge the defect site [103] (Fig. 14.8). Human adipose-derived stem cells have also been

Figure 14.8 (A) X-ray, (B) μCT images of scaffolds implanted in rat femurs, and (C) Regenerated bone volume is significantly increased in a rat femoral segmental defect model treated with BMP-2-loaded photo-cross-linked alginate hydrogels incorporated with mesenchymal stem cells compared to those without BMP-2 treatment. BMP, Bone morphogenetic protein. Source: Reprinted by permission from John Wiley & Sons, Inc. from Ho SS, Vollmer NL, Refaat MI, Jeon O, Alsberg E, Lee MA, et al. Bone morphogenetic protein-2 promotes human mesenchymal stem cell survival and resultant bone formation when entrapped in photocrosslinked alginate hydrogels. Adv Healthc Mater 2016;5:2501
9. Available from: https://doi.org/10.1002/adhm.201600461.

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successfully suspended in BMP-laden gelatin hydrogels to promote osteogenesis [110]. These strategies highlight the trend in research efforts toward a more complex biomimicry approach, combining the inorganic mechanical support structures of natural bone with high water-content gels for cell suspension and sustained release of bioactive ingredients including BMP to stimulate rapid tissue remodeling.

14.4.2.3 Matrix bone morphogenetic protein surface coatings Surface coatings are employed to improve cellular attachment on otherwise inert or unfavorable scaffold surface topologies. In particular, surface roughness increases the surface area available for the attachment of biological components and host tissue
implant interactions [11]. For example, metal implants and otherwise inert scaffold surface coatings are pivotal for increasing bioactivity and osseointegration within the defect site [111]. Existing research widely recognizes the critical role of surface coatings for optimizing the biocompatibility of scaffolds to improve integration into the defect areas [112,113]. Moreover, advanced techniques have demonstrated the use of surface coatings to provide spatiotemporally controlled BMP delivery to the defect site. Metallic implants, although nonregenerative, have been widely used particularly in the restoration of joint function and for defects in load-bearing bones [114,115]. Commonly, titanium, titanium alloys, or stainless steel are used for nonregenerative implants. Metals generally have excellent mechanical properties, including high strength and wear resistance, making them particularly suitable for high loadbearing regions. Metallic implants have seen recent worldwide success in a number of treatments including total calcanectomy and sternocostal reconstructions using 3D printed titanium replacements [116,117]. While these treatments have been largely successful, there are still a number of recognized drawbacks with metal implants, from inconveniences during daily life (such as airport security) to significant risks of toxicity and immune reaction due to the release of metal ions into the bloodstream after wear [118]. Also, the lack of tissue adherence has led to the development of biocompatible surface coating for more effective integration into the defect site [119,120]. Therefore surface modifications have been proposed to enhance the host
implant interaction and subsequent osseointegration, including a variety of biomaterials mimicking natural bone [121]. In addition to HA and calcium-based coatings, BMP has been included in surface coatings to instigate rapid bone formation at the implant interface [122,123] (Fig. 14.9). Thorey et al. reported the in vivo characterization of bone formation around titanium implants with BMP covalently bound to the surface as well as absorbed into the structures implanted into rabbit femoral condyles, demonstrating superior bone formation for the BMP-coated scaffolds compared to uncoated controls [124]. Dual growth factor delivery was achieved by Ramazanoglu et al. by coating titanium implants with a calcium phosphate surface onto which BMP-2 and VEGF were loaded. Enhanced density of mineralized bone around the implants was observed after 4 weeks implantation in pig calvaria [125]. Using a different strategy, BMP-loaded TiO2 nanorod films were fabricated with the addition of bioactive

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Figure 14.9 BMP-2-coated titanium screws implanted for 5 weeks, stained McNeil’s Tetrachrome, basic Fuchsine and Toluidine Blue O staining. (A) Bone formation is evident around the entire BMP-2-coated titanium screw, stained in red; and (B) bone is seen in direct contact with the titanium surface. BMP, Bone morphogenetic protein. Source: Reprinted from Liu Y, de Groot K, Hunziker EB. BMP-2 liberated from biomimetic implant coatings induces and sustains direct ossification in an ectopic rat model. Bone 2005;36:745
57. Available from: https://doi.org/10.1016/J.BONE.2005.02.005 with permission from Elsevier.

glass and loaded with BMP. Accelerated osteogenic differentiation was observed in vitro using bone marrow
derived MSCs whereby the improved surface characteristics of the films enhanced cellular attachment [126].

14.4.3 Complex biomaterials and biofabrication systems 14.4.3.1 Bioactive materials Fundamentally, the use of BMP in bone regeneration implants is to induce osteogenesis or improve bioactivity. Given the existing concerns and limitations surrounding BMP use, as discussed early, bioactivity from other sources is being investigated in bone TERM research to move away from the use of BMP altogether [127]. Only a number of systems have achieved this so far, focusing on biomaterials with high intrinsic bioactive properties to induce osteogenic gene expression. Examples include the development of advanced formulations of bone cements that are based on methyl methacrylate/methacrylic acid in a self-curing cement with collagen-loaded polymer microspheres to simulate the delivery of BMP into the defect site [128]. Also, Stevens et al. have recently developed a strontiumsubstituted bioactive [129] which, when incorporated with PCL, can be readily fabricated into 3D-printed structures. A number of publications have demonstrated the efficacy of fabricating bioactive scaffolds using fused deposition modeling 3D printing [130], melt electrospinning writing [131,132], and calcium phosphatecoated scaffolds which exhibited osteoblast-related gene expression in vitro even without the addition of osteogenic media [133]. These strategies, while still in their infancy, point to a post-BMP future where growth factors may be removed entirely

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from bone regeneration devices to reduce costs, side effect profiles and improve patient outcomes.

14.4.3.2 Advanced biomanufacturing Using 3D printing (additive manufacturing) techniques, highly controlled spatial patterning is achievable to create complex structures with optimal biological performance using some of the next-generation strategies listed previously. Biomanufacturing encompasses the use of a range of advanced manufacturing techniques that can arrange biological materials into 3D constructs [134]. Additive processes, including 3D bioprinting, have seen a significant rise in attention over the last decade due to their ability to process a wide range of materials in biological-friendly conditions for biofabrication applications. The combination of the advancements in TERM research to incorporate BMP into materials, controlled release systems, and coatings have been applied to biofabrication strategies to create patient-specific, mechanically supportive implants with high biological functionality [135]. The parallel development of bioactive materials, for use as mechanically robust scaffolding, in combination with highly biologically functional cell-laden hydrogels and controlled-release growth factor systems provides an advanced platform for the fabrication of patient-specific tissues that will allow for rapid tissue regeneration. As an all-in-one biofabrication system, sterility, customization, and rapid fabrication, these systems aim to be readily integrated into a hospital environment (Fig. 14.10). While combination systems have yet to be developed, substantial progress has been made in the field of bioprinting, both in the commercial and research spheres

Figure 14.10 Advanced biofabrication systems aim to combine multiple cell types in precisely arranged complex hydrogel networks with tissue engineered blood vessel infiltration, micron-scale degradable, and bioactive scaffolds for mechanical support and complex controlled-release systems such as microparticles for the safe, cost-effective, and efficient delivery of BMP within the implants. 3D Printer image courtesy of Dr. Sean Powell, Queensland University of Technology (QUT).

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whereby 3D bioprinters have the ability to fabricate complex 3D biological structures using a range of biomaterials and BMP carrying method discussed in this article. For example, using inkjet bioprinting, Phillippi et al. demonstrated a proof-of-concept method to spatially arrange BMP-2 in specific patterns onto a fibrin substrate to control muscle-derived stem cell differentiation into osteogenic and myogenic lineages [136]. It was shown that variation in the patterning of the structure allowed control over the cell differentiation pathways. Recent studies have also reported successful fabrication of porous alginate and polyvinyl alcohol scaffolds cobioprinted with BMP-2 and bovine serum albumin [137] as well as MSC-laden gelatin methacrylamide bioprinted structures with BMPs bound to collagen microfibers [138].

14.5

Conclusion

In conclusion, BMPs have been of significant interest in the development of bone tissue engineering solutions and clinical treatments, particularly in spinal care. However, despite the significant research in the fields of TERM, financial investment, and translational efforts to drive BMP products into the clinic, shortfalls in product efficacy and safety typically associated with its off-label use has seen a shift in the research landscape. Researchers are now seeking to overcome these challenges through advanced biofabrication techniques, using safer, lower doses in more complex delivery systems to improve the spatiotemporal delivery of BMPs within the localized defect and develop techniques for harnessing the regenerative capacity of BMPs. This shift in research focus will ultimately deliver new products that promise high levels of safety, lower costs, and significantly improved patient outcomes.

Acknowledgement This research is supported by the Australian Research Council Industrial Transformation Training Centre in Additive Biomanufacturing and Anatomics Pty Ltd.

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34. Available from: https://doi.org/10.1634/stemcells.2007-0520. [137] Luo Y, Luo G, Gelinsky M, Huang P, Ruan C. 3D bioprinting scaffold using alginate/ polyvinyl alcohol bioinks. Mater Lett 2017;189:295
8. Available from: https://doi. org/10.1016/j.matlet.2016.12.009. [138] Du M, Chen B, Meng Q, Liu S, Zheng X, Zhang C, et al. 3D bioprinting of BMSCladen methacrylamide gelatin scaffolds with CBD-BMP2-collagen microfibers. Biofabrication 2015;7:44104. Available from: https://doi.org/10.1088/1758-5090/7/4/ 044104.

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Adipose tissue engineering

15

Fiona Louis1 and Michiya Matsusaki1,2 1 Joint Research Laboratory (TOPPAN) for Advanced Cell Regulatory Chemistry, Graduate School of Engineering, Osaka University, Suita, Japan, 2Department of Applied Chemistry, Graduate School of Engineering, Osaka University, Suita, Japan

15.1

Introduction

Adipose tissue is the largest tissue in the body, representing 10%
30% of total human body weight [1]. This is a loose connective tissue, which can store energy in the form of lipids in its fat cells called adipocytes. Adipose tissue has also important endocrine and secretory functions, which appear vital for many physiological responses of the organism such as metabolic homeostasis, angiogenesis, inflammatory, and immune regulation [2]. The maintenance of these functions is determinant for the construction of an in vitro artificial adipose tissue model suitable for cosmetic and pharmaceutical assays, or for plastic and reconstructive surgery purposes. Adipose tissue can be easily harvested from patient, but its implantation for filling soft tissue defects remains a reconstructive challenge. Over 5.6 million of these kinds of procedures are performed annually in the United States [3]. The majority of the operations are needed after tumor resection or following trauma, to resorb visible asymmetry or large wounds [4,5]. Soft tissue reconstruction frequently employs the use of autologous grafts, including vasculature; but these procedures are highly invasive and show donor site morbidity. The implantation of adipose tissue from less invasive liposuction, which does not include blood supply, is associated with longterm problems of migration, extrusion, or immune response, like the capsular contracture, where collagen fibers, which appeared around the implant, start to shrink and compress it. There is a special need of a complete regeneration of a healthy adipose implant to ensure long-term successful clinical outcomes after tissue reconstruction.

15.2

The adipose cells for the reconstruction of in vitro models

The adipose tissue appears mostly composed of adipocytes that store lipids in multiple fat vesicles until merging in a unique fat vesicle in mature adipocytes. More in details, the adipose tissue contains in fact four different components important for its homeostasis (Fig. 15.1). Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00008-8 © 2020 Elsevier Ltd. All rights reserved.

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Figure 15.1 The main components of the adipose tissue. Adipose tissue is mainly composed of mature adipocytes, adipose-derived stem cells, and endothelial cells, surrounded by their extracellular matrix. All these components are of importance for the whole tissue homeostasis. Source: Credit: Fiona LOUIS.

The main cells are the mature adipocytes, the adipose-derived stem cells (ADSCs), and the endothelial cells. Both mature adipocytes and ADSC secrete several key growth factors while endothelial cells allow the diffusion of the oxygen and the nutrients. Some studies showed that both ADSC and mature adipocytes in addition to endothelial cells, as well as a support (scaffold), are required to maintain the survival of the adipose graft [6]. To recreate the in vitro adipose tissue, both primary cells and cell lines have been employed. Adipose cell lines and primary adipose stem cells are the most used to study the process of adipogenesis, while primary mature adipocytes are more difficult to maintain in vitro.

15.2.1 Preadipocytes cell lines The most used cell lines for in vitro adipose tissues are the 3T3-L1, the 3T3-F442a, and the Ob17 murine cell lines [7
9]. They are homogeneous, well defined and can be extended in culture. Adipocytes cell lines differentiate much more readily than primary cells and can even spontaneously convert into adipocytes in the presence of serum when growth arrest is maintained through confluence [10,11].

15.2.2 Primary bone marrow mesenchymal stem cells Mesenchymal stem cells (MSCs) can be isolated from embryonic or adult tissues, including the bone marrow MSCs (bMSCs). bMSC can differentiate into mesenchymal

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lineages, including bone, cartilage, tendon, ligament, marrow stroma, adipocytes, dermis, muscle, and connective tissue [12]. These primary cells have been widely used for in vitro adipose model and regenerative strategies [13,14].

15.2.3 Primary adipose
derived stem cells Another type of MSCs, ADSCs, is relatively abundant and can be harvested more easily and less invasively than bMSC. As they present no apparent disadvantages compared to other MSCs, they are the preferred source [15]. Freshly isolated ADSCs have also the capacity to differentiate along the multiple mesenchymal lineages [14], while cells harvested from cultured ADSC appear quite homogeneous with limited possible commitments [16]. ADSCs secrete a variety of growth factors, including anti-inflammatory and proangiogenic cytokines, modulating the local immune response and improving the long-term maintenance of the graft survival and its retention [17]. They can also differentiate themselves into vascular endothelial cells [18] contributing to the graft vasculogenesis [19]. Finally, these cells may serve as a reservoir for de novo adipogenesis on site [20] to generate adipose substitutes in soft tissue defects [21].

15.2.4 Primary mature adipocytes The use of primary mature adipocytes for in vitro adipose tissue still remains a challenge. As they possess a cytoplasm composed of 80%
90% lipids, they are easily traumatized by the mechanical forces of liposuction, resulting in 90% damaged cells. The remaining 10% tend to form localized necrosis after injection. Being already terminally differentiated, they cannot proliferate and the maintenance of their unilocular shape is difficult in vitro, they usually dedifferentiate after 1 week [22,23].

15.2.5 The importance of the vascularization in adipose models Both in vitro and in vivo regenerated adipose tissues require a certain long-term stability to allow formation of tissue structural organization and to become incorporated into the host vasculature, respectively. This stability can be impaired by the reduced availability of oxygen and nutrients in the inner parts of the reconstructed tissue. In vivo, adipose tissue appears highly vascularized, each adipocyte being in contact with at least one blood vessel [24]. This vascularization was built during the postnatal adipose tissue development, preceding the adipocyte differentiation and maturation [25,26]. Blood vessels are crucial for adipose tissue regeneration and should be added in vitro or to be built by de novo angiogenesis on site.

15.2.6 Human cells or other species? To achieve a suitable adipose model reconstruction with physiological functionalities, the differences between murine and human adipose tissue have to be

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considered. Adipose tissue is a multidepot organ in both humans and rodents, but anatomical differences in deposition location exist and adipocytes metabolic heterogeneity between species can be found. This is of importance, knowing that usually murine fat pads have to be pooled before mouse ADSC isolation to obtain sufficient sample amount, which is unnecessary with human tissue samples. Several studies comparing human and murine adipose tissues showed that although many similarities can be observed, it clearly appears that weaned laboratory rodents are not the perfect model for humans in metabolic or developmental studies. The metabolic differences are especially seen when glucose transport sensitive to insulin experiments or catecholamine-stimulated lipolysis (activation of β1, β2, β3 receptors) is performed, murine adipose tissue being much more sensitive than human [27,28]. Also, the existence of scattered brown adipocytes in human adult is still debated while they are easily found in mice [29]. While adipogenesis in rodents and humans has been extensively studied, this is not the case of adipocytes from ruminant species. The reason is the differences in lipid metabolism between ruminant and nonruminant species. The lipids synthesis itself displays differences. Rodents and humans lipids/fatty acids are primarily synthesized in the liver, using glucose for de novo fatty acid synthesis [30], while in ruminant animals, lipids are generated into adipocyte droplets using acetate as the principal precursor [31,32]. Ruminant adipocytes are thus less responsive to circulating insulin because glucose is not their major source of fatty acid synthesis [33]. Therefore the only physiological models available as human models are primates, which make perfect human-like adipose models difficult using animal, further justifying the cessation of animal experimentation in favor of the use of human cells for adipose tissue engineering.

15.3

Current existing in vitro adipose tissues models

Differentiated adipocytes, representing the final goal to reach in adipose tissue regeneration, are difficult to culture in vitro due to their lipids content. In twodimensional (2D) culture, they detach from the bottom of the plastic dish and float in the culture medium. Three-dimensional (3D) structure is thus even more required for these mature adipocytes than all the other cell types. Among the different in vitro 3D adipose models, cell aggregation in spheroids or the use of both synthetic and natural polymers have been extensively studied for in vitro models (Fig. 15.2). Various advantages and disadvantages can be observed in the scaffolds regarding their biocompatibility, mechanical and chemical properties, or degradability. The scaffolds can exist as powders, injectable hydrogels, sheets, beads, or macroporous foams. In all conditions, the scaffold stiffness is of importance and can control the differentiation of ADSC even in the absence of additional growth factors: into myogenic or osteogenic commitments when cells are forced to spread on mild or stiff surfaces (at least 8
17 kPa) [34] or into adipogenic differentiation when cells

Figure 15.2 Current adipose tissue engineering strategies. Cells and structural components are all of importance for the regeneration of a functional adipose tissue presenting physiological adipose characteristics. The addition of other adipose cell types or surrounding tissues should allow the improvement of the model toward a fully relevant engineered adipose tissue suitable for drug screening assays. Source: Credit: Fiona LOUIS.

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present round shape on soft surfaces (at 2 kPa, like adipose tissue native stiffness) [35]. Also, low cell spreading on restricted cell-mediated degradation surfaces can favor the adipogenesis of stem cells [36].

15.3.1 Reconstruction of an in vitro adipose tissue without scaffold Adipose stem cells need high cell
cell contacts as inducer of adipocyte differentiation. High cell density is also important for mature adipocytes to maintain their functionality, a reduced cell concentration leading to their dedifferentiation [37,38]. In 2D condition, only horizontal cell
cell contacts can be reached in confluent cells, while 3D scaffold-free cultures provide cell
cell contacts in all dimensions. Self-spheroid formation happens in cell culture when cell
cell interactions dominate over cell
substrate interaction. The easiest way to achieve this condition is using low-attachment cell culture surfaces. Covalently bound hydrophilic and neutrally charged hydrogel layer on the surface, for instance, a polymer containing the phosphorylcholine [Lipidure-COAT plate (AmsBio)] or the elastin-like polypeptide
polyethyleneimine, can be added on classic plastic cell culture dishes to alter cell morphology and induce adipogenic differentiation of MSCs [39]. To avoid cell attachment, substrate nanopatterning can also be used. The substrate topography affects the cell shape, the differentiation and the genes expression. A nanopattern that allows individual MSCs to remain rounded will induce their adipogenic differentiation, whereas straight grooves and grids, which cause the cells to elongate, increase their osteogenesis [35,40]. This method is also suitable to induce cell aggregates of human ADSC adipogenesis successfully using a specific micropatterned surface [41]. Another way to induce cell aggregation is to isolate the cells from the bottom surface like with the hanging-drop method, which allows the formation of adipose spheroid in a suspended drop before being transferred to low attachment plate for culture. Scalable adipose spheroids (200
500 μm diameters, necrosis can happen above 500 μm) can be produced that are suitable for quantitative classic adipogenesis assessments [25,42,43]. In a similar way, magnetic nanoparticles incorporated inside the cells allow the construction of 3D spheroids by promoting levitating cell to cell interactions. The resulting magnetic levitation adipospheres displayed more efficient lipid droplet accumulation than 2D cultures and neovascularization was even possible by coculturing with endothelial cells [44]. The last method for in vitro adipose tissue reconstruction without scaffold is to perform culture on a chip. Preadipocytes aggregated in the chip and subsequently formed adipose tissue with microfluidic nutriments and adipogenesis growth factors supply. Multilocular adipocytes were induced, showing extracellular matrix (ECM) secretion (collagens) and lipids uptake [45].

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15.3.2 Use of synthetic scaffolds Biodegradable synthetic polymers are among the most widely scaffolds used for adipose tissue engineering due to the possibility to control their chemical and mechanical properties as well as their degradability, and because many are already FDA approved for clinical use (e.g., sutures and mesh). Some examples of the existing synthetic scaffolds, which have shown potential in supporting regenerated adipose tissue, are summarized in Table 15.1. The polylactic-co-glycolic acid scaffold types are the most synthetic scaffolds used for in vitro adipose tissue reconstruction. Their degradability can be controlled by varying the molecular weight, the crystallinity, and the ratio of lactic to glycolic acid subunits [48,53
55]. Few synthetic scaffolds allowed to get a full differentiation of ADSC until mature adipocytes. The ones who succeeded used murine 3T3-L1 that is known to differentiate faster than human ADSC and they required at least 1 month of differentiation. More than half studies performed in vivo implantation experiments but generally of only 1- or 2-month durations. The in vivo results showed more induced vascularization than adipogenesis. In addition, we can have some concerns about the potential impact of synthetic scaffolds degradation products on surrounding cell functions.

15.3.3 Natural components in adipose tissue engineering Natural scaffolds are made from components that can be isolated from living organisms. They are biocompatible and they present in vivo
like mechanical and biological properties. They are thus often applied to replace or restore structure and function of damaged tissues/organs, due to their ability to adequately support cell adhesion, migration, proliferation, and differentiation. Naturally derived biomaterial can be classified into many groups, including polysaccharide-based biomaterials (cellulose, chitin/chitosan, glucose, etc.), protein-based biomaterials (collagen, gelatin, silk, etc.), and decellularized tissue
derived biomaterials, created by eliminating all cells from native tissues/organs. Table 15.2 presents some examples of the published natural scaffolds for in vitro adipose tissue reconstruction. Most of the current natural scaffold models used primary human ADSC to validate their scaffolds and more than half of them performed subsequent in vivo murine implantation experiments. Mature adipocytes differentiated from ADSC were found mostly in physiologically relevant adipose ECM components’ scaffolds like in hyaluronan, gelatin, laminin, collagen, matrigel, and decellularized tissues. Collagens are indeed the principal adipose ECM component [at 25%
35% (by dry weight)] [100,101], among which collagen type I is the most prevalent [102,103]. Collagen type I is thus often used for inducing ADSC adipogenesis up to mature adipocytes phenotype, in different forms such as coating, sheets, hydrogels, or microfibers. Matrigel derives from the basement membrane of mouse sarcoma and contains undefined mixture of ECM protein and growth factors. It has been extensively studied with successful adipogenesis of multiple cell types and induced

Table 15.1 Examples of synthetic scaffolds used in adipose tissue engineering. Synthetic scaffolds

Cell used

Culture time in vitro

Differentiation up to mature adipocytes

In vivo experiments

Remarks

References

Polypropylene

hADSC

1 month

No, still multilocular adipocytes

Yes, up to 9 months

[46]

Polytetrafluoroethylene

hADSC or rADSC

5 days

Cells not shown

No

Polylactic-co-glycolic acid

Murine 3T3-L1 preadipocytes hbMSC

6.5 months

Yes, up to 24 weeks No

Murine 3T3-L1 preadipocytes Murine 3T3-L1 preadipocytes

15 days

Yes, all mature adipocytes from day 35 No, still multilocular adipocytes Yes, some mature adipocytes

Good dimensional stability of the scaffold with adipogenesis and neovascularization induced after in vivo implantation Hydrogels were coated with collagen type 1, albumin, or fibronectin to improve the cell adhesion Neovascularization induced after in vivo implantation


No




[50]

Yes, up to 4 weeks

Neovascularization induced after in vivo implantation

[51]

rADSC

7 days

[52]

10 days

Yes, up to 8 weeks Yes, up to 6 weeks




hADSC

Neovascularization induced after in vivo implantation

[53]

hADSC or rabbit bMSC

1 or 2 weeks

Yes, high amount of mature adipocytes from day 21 and all mature adipocytes from day 35 No, still multilocular adipocytes Yes, but very few mature adipocytes, more multilocular adipocytes Cells not shown

Yes, up to 2 or 8 weeks

[54,55]

rbMSC

4 weeks

No, still multilocular adipocytes

No

FGF-2 release from the scaffold was found to enhance adipogenesis and neovascularization after in vivo implantation Additional bFGF in the scaffold induces the adipogenesis in vitro of encapsulated cells

21 days

39 days

[47]

[48] [49]

[56]

Polyethylene glycol

Murine 3T3-L1 preadipocytes

Up to 42 days

Yes, some mature adipocytes

No

Polyethylene glycol diacrylate

hbMSC

4 weeks

Yes, some mature adipocytes

Yes, up to 4 weeks

hbMSC

1 week

No, still multilocular adipocytes

Yes, up to 4 weeks

Polyethylene terephthalate

3T3-L1 cells

34 days

No

Pluronic F-127

rbMSC

14 days

No, still multilocular adipocytes at day 4 and no images at 34 days No, still multilocular adipocytes

Poly(amidoamine) oligomer macroporous foam

Murine 3T3-L1 preadipocytes

17 days

No, still multilocular adipocytes

Yes, up to 50 days

No

Additional laminin-derived binding protein increased the in vitro preadipocyte cell attachment and their adipogenesis Scaffold maintains predefined shape and dimension after in vivo implantation Scaffolds maintained 100% of their volume and adipogenesis induced after in vivo implantation


[57]

More cell-interactive composite scaffold and in vitro adipogenesis even without adipogenic medium RGD sequences induced the integrinmediated in vitro preadipocyte binding and adipogenesis; neovascularization induced after in vivo implantation

[61]

[58]

[59]

[60]

[62]

Mature adipocytes correspond to the adipogenic differentiated state where adipocytes contain only one lipid vesicle filling all their cytoplasm. hADSC, Human adipose
derived stem cells; hbMSC, human bone mesenchymal stem cells; rbMSC, rat bone mesenchymal stem cells: rADSC, rat adipose
derived stem cells. FGF-2, fibroblast growth factor 2; bFGF, basic fibroblast growth factor; RGD, Arg-Gly-Asp. Source: Credit: Fiona LOUIS.

Table 15.2 Nonexhaustive list of the common natural scaffolds used in adipose tissue engineering. Natural scaffolds

Cell used

Culture time in vitro

Differentiation up to mature adipocytes

In vivo experiments

Remarks

References

Hyaluronan

hADSC

Yes, some mature adipocytes

No




hADSC

Up to 17 or 36 days 14 days

Cells not shown

hADSC







Yes, up to 12 weeks Yes, up to 8 weeks

hADSC, 3T3-L1, 3T3F442A

Up to 30 days

Yes, most of mature adipocytes from day 30

No

Aliginate

hADSC

10 days

Chitosan

rADSC

11 days

No, still multilocular adipocytes Cells not shown

Fibrin

hADSC

14 or 28 days

rADSC

13 days

Yes, 10 weeks Yes, up to 14 days Yes, 6 weeks Yes, up to 1 year

hADSC

14 days

Cells not shown

Yes, up to 9 months

hADSC

7 days

No, still multilocular adipocytes

Yes, 7
8 days

Neovascularization induced after in vivo implantation Noncell-adhesive cross-linked hyaluronan enhances even more the adipogenesis, advantageous in terms of adipogenesis and angiogenesis Hyaluronan functionnalized with collagen I, collagen VI, and the cell-binding domain of fibronectin Adipogenesis and neovascularization induced after in vivo implantation Adipogenesis and neovascularization induced after in vivo implantation Adipogenesis and neovascularization induced after in vivo implantation Adipogenesis and neovascularization induced after in vivo implantation. No evidence of cell necrosis, cystic spaces, nor fibrosis Adipogenesis and neovascularization induced after in vivo implantation. No signs of an inflammatory response or evidence of tissue necrosis in the implants Neovascularization induced after in vivo implantation

[63] [64] [65]

No, still multilocular adipocytes No, still multilocular adipocytes

[66]

[67]

[68] [69] [70] [71] [72] [73]

[74]

[75] [76]

Silk




[77]

No

Induced vascularization

[78]

Yes, 4 weeks Yes, 4 weeks or up to 9 months Yes, 6 weeks

Adipogenesis induced after in vivo implantation Adipogenesis and neovascularization induced after in vivo implantation

[79]

[81]

No

bFGF gelatin beads induce adipogenesis and neovascularization after in vivo implantation Sandwich culture of adipose tissues fragments between hADSC monolayer and gelatin gel. adipogenesis induced after in vivo implantation Methacrylated gelatin

No




[57]

No

Collagen coating on polystyren beads

[83]

No

Collagen sheets

[84]

Yes, 4 weeks No

Collagen gel, adipogenesis induced after in vivo implantation Collagen microfibers

[85] [86] [87]

Up to 3 months

hADSC and HUVEC hbMSC and hADSC

Up to 2 weeks 22 days

hADSC

4 weeks

No, still multilocular adipocytes

hADSC

14 days

Cells not shown

Coculture human adipose tissue fragments and hADSC

Up to 55 days

Maintenance of mature adipocytes at least 14 days, no differentiation of the hADSC

Yes, 10 days

14 days

Laminin

Human mature adipocytes 3T3-L1

Collagen

C3H10T1/2

10 days

hbMSC

51 days

mADSC or hADSC

9 or 14 days

Maintenance of mature adipocytes Yes, most of mature adipocytes at day 42 No, still multilocular adipocytes Yes, most of mature adipocytes at day 51 Yes, some mature adipocytes

Mouse mature adipocytes and hADSC

Up to 21 days

Gelatin

Up to 42 days

Maintenance of mature adipocytes and vascularization No, still multilocular adipocytes Yes, some mature adipocytes

No

Liquefied adipose tissue

Maintenance of mouse mature adipocytes and some human mature adipocytes at day 21

[80] [46]

[82]

[22]

(Continued)

Table 15.2 (Continued) Natural scaffolds

Matrigel

Decellularized human placenta Decellularized adipose tissue

Cell used

Culture time in vitro

Differentiation up to mature adipocytes

In vivo experiments

Remarks

References

Human mature adipocytes Mouse adipose tissue fragments

14 days

Maintenance of mature adipocytes at least 7 days Maintenance of mature adipocytes at least 7 days

No

Collagen gel

[22]

No

Collagen gel

No

Collagen gel

[88] [89] [23] [90]

No

Collagen gel, on-chip culture

[91]

Yes, 2 weeks Yes, up to 6 weeks

Adipogenesis and neovascularization induced after in vivo implantation Matrigel supplemented or not with 1 ng/mL bFGF, adipogenesis and neovascularization induced after in vivo implantation Adipogenesis and neovascularization induced after in vivo implantation

[25]

Up to 4 weeks

Human bone marrow adipose tissue fragments and hbMSC

Up to 3 weeks

Human mature adipocytes mADSC

Up to 36 days Up to 18 days

Maintenance of mature adipocytes at least 7 days, still multilocular adipocytes differentiated from hbMSC after 1 week Maintenance of mature adipocytes at least 6 days Yes, some mature adipocytes

3T3-F442A preadipocytes







hADSC

Up to 14 days

Yes, some mature adipocytes

Yes, 10 days

Human mature adipocytes hADSC

Up to 7
10 days 5 days

Maintenance of mature adipocytes for 7
10 days No, only proliferating cells

Up to 10 days

hADSC and rADSC

Up to 14 days

No, still multilocular adipocytes No, still multilocular adipocytes

maintenance of mature adipocytes after in vivo implantation Adipogenesis and neovascularization induced after in vivo implantation


[94] [95] [96]

hADSC

Yes, up to 3 months Yes, up to 6 weeks No Yes, up to 12 weeks

Adipogenesis and neovascularization induced after in vivo implantation

[98] [99]

[92] [93]

[64] [66]

[97]

Mature adipocytes correspond to the adipogenic differentiated state where adipocytes contain only one lipid vesicle filling all their cytoplasm. hADSC, Human adipose
derived stem cells; mADSC, mouse adipose
derived stem cells; rADSC, rat adipose
derived stem cells; hbMSC, human bone mesenchymal stem cells; HUVEC, Human umbilical vein endothelial cell; bFGF, basic fibroblast growth factor. Source: Credit: Fiona LOUIS.

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graft neovascularization after in vivo implantations, but its clinical application is limited by its murine and tumorigenic origins. Decellularized tissue scaffolds are the perfect imitation of the in vivo cell environment components. It consists of removing the cells from human tissues using usually multiple freeze-thaw cycles and lyophilization. These acellular scaffolds fully mimic the structural and biochemical properties of native in vivo ECM of placenta or adipose tissue. The other types of natural scaffolds such as alginate, chitosan, fibrin, and silk showed only multilocular adipocytes even after 28 days of differentiation. These less physiological scaffolds were found anyway to induce the adipogenesis and neovascularization of ADSC after in vivo implantation. Concerning these in vivo assessments, most of the studies checked the implanted graft after 1
4 months only and few went until 9
12 months. These short durations were generally enough for the adipogenesis and the revascularization of the grafts, but the long-term stability assessment should be improved. The use of natural materials for tissue engineering applications has advantages with respect to biocompatibility, while their mechanical and biological properties tend to match those found in vivo. The recent advances using decellularized matrix appears ideally, decellularization being designed to remove the immunogenic materials from donated biological sources while retaining as many constitutive components as possible. However, clinical applications are still needed to validate their potential for soft tissue replacement applications.

15.4

Medical applications of adipose tissues grafts

Adipose tissue is the key component of soft tissues throughout the body, protecting underlying structures and imparting a normal appearance. Several types of cancers such as breast or facial cancer need tissue resection, leaving patients with soft tissue defects and disfiguration to fill using soft tissue regeneration. The same happens after work-related trauma, as well as after burns, which necessitates the reconstruction of not only the skin but also subcutaneous soft tissue. Engineered adipose tissues can be the key for the reconstruction of all these soft tissues needs. But despite all the publications investigating synthetic and natural materials for adipose tissue engineering, human clinical studies using 3D engineered adipose tissue constructs for soft tissue reconstruction have not yet been initiated. The only in vivo models are summarized in Tables 15.1 and 15.2, using frequently athymic mice to assess the volume retention of human fat grafts as well as the growth/differentiation of ADSC in scaffolds within a living system. This animal model has the advantage to provide physiological conditions, but it is distant from clinical application in terms of size, function, and immune health. In current surgery applications, autologous soft tissue grafts or synthetic materials alone are predominant for reconstruction procedures. Each approach has its own pros and cons and depends to the type of patients and the soft tissue defects based on clinical judgment. In term of patients’ feelings, autologous adipose tissue is

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often used as the preferable filler, providing a more natural shape and suppleness; but the stability of the graft in long-term assessments gives unpredictable results and represents substantial challenges for contemporary medical practice. The following sections present the current applications of adipose tissues implantations, alone or mixed with fibrin sealant, in soft tissue surgery, showing the range of possible applications of future developed engineered adipose tissues.

15.4.1 For cosmetic surgery Autologous fat grafts in cosmetic surgery procedures focus on enhancing and reshaping structures of the body to improve appearance and confidence. The common types of cosmetic surgery operations consist in breast augmentation, fat injection to reduce the wrinkles or increase the cheeks and lips, etc. The facial adipose tissue is separated into the deep and superficial layers. During aging process, the volume of one compartment can shift downward or migrate, leading to an alteration of facial contours [104]. However, the autologous adipose tissues injected showed a highly unpredictable volume loss, which relies on its revascularization. In a longterm follow-up studies on patients treated for facial wrinkles, only 3%
4% maintained long-term correction for more than 14 months, avoiding early cell death and histologic fibrosis [105]. Only small volumes of autologous fat grafting, where diffusion can support cell survival, within a 2-mm distance from an artery, were found with maintained success after transplantation [106,107]. These small grafts improved the facial symmetry and the extremity contour defects [108,109].

15.4.2 For reconstructive surgery Due to its large volume, breast reconstruction using adipose tissue transplantation is one of the most challenging reconstructive surgery operations. Autologous fat grafting for breast reconstruction was abandoned for several decades and has recently been used again due to the recent problems with breast synthetic prosthesis. Adipose tissue implantation allows for more natural shape and suppleness, while allowing movements and aging of the reconstructed breast to remain comparable with the contralateral breast. However, autologous fat transplantation also showed unpredictable resorption rates, ranging from 30% to 70%, requiring corrections or repeated fat injections to achieve satisfactory results [110]. It was shown that patients with a larger volume injected and longer length of follow-up had significant increases in risk of developing necrosis in the graft. One retrospective review on 176 patients between 2011 and 2016 found 10.5% incidence of fat necrosis with 5% chance of biopsy or excision [111]. Also, fat transfer usually needs a volume of tissue extracted from the donor whose size is greater than the potential graft site, leading to a substantial donor site morbidity involved with the procedure. Without actual good in vitro model of implantable fat tissues, one of the best current procedures for breast regeneration is to use deep inferior epigastric perforators flap. Adipose tissue fragments are extracted from the abdomen and transferred

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to the new breast. Then the surgeon performs microsurgery to attach the blood vessels from the abdominal graft to the breast vasculature. This procedure is timeconsuming and is not totally without morbidity, patients being at high risk for subsequent complications such as wound infection or reoperation [112].

15.4.3 For wound healing Adipose tissue implantation is also used for large wound healing. Studies have demonstrated that fat grafts with additional ADSC can promote wound healing [113,114], while improving the neoangiogenesis around the graft [115,116]. ADSCs are also relatively resistant to the wound hypoxic conditions and can also contribute to adipose tissue regeneration, by undergoing adipogenic differentiation [117]. A common example of wound healing is the case of chronic leg ulcers in diabetic patients, which are complicated to treat, and results in high morbidity as well as significantly reduced quality of life. ADSC implantations, alone or into a fibrin sealant, were shown to improve the healing of chronic ulcers and to reduce the pain. Further studies, however, are needed to define their long-term safety and efficacy [118].

15.4.4 For bone healing Despite the substantial medical advances in fractures healing, 5%
10% of annual fractures in the United States end in non-union. Several studies showed that bMSC and ADSC are induced to migrate to the fracture site to promote bone healing [119]. Adipose tissue graft implantation was found to enhance the rate of fracture healing with a significant increase of new bone tissue in rabbits from 14 days [120]. Other studies with rats [121], rabbits [122], and dogs [123] using ADSC alone also showed their potential for osteogenic differentiation. During the healing process, the adipose tissue seemed to be absorbed and replaced by bone tissue, without tissue volume lost.

15.5

Further developments needed

A multiple of reconstructed adipose tissue models exist in vitro, but they still need some improvements to mimic more the in vivo surroundings adipose tissue, of importance to get relevant data for drug screenings applications of for metabolism pathway studies.

15.5.1 Vascularized adipose tissues So far, in vitro engineered adipose tissue models cannot reproduce the in vivo vascular support system. Many attempts were made to get de novo angiogenesis in the in vitro tissues, but finding a functional culture medium allowing both

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differentiation of adipocytes up to mature differentiated state and endothelial cells in coculture still represents a major challenge [124
127]. The importance of ADSC for the formation of endothelial tubular structures was also shown [128], along with ECM stroma interactions for directing the vessel growth [78,129].

15.5.2 Addition of other surrounding cells types 15.5.2.1 Adipose pericytes Pericytes are contractile cells that enclose the blood vessels. Some of them can differentiate into fat cells in response to glycerol, but their implication in fat accumulation is still not known [130]. Adding pericytes in the engineered adipose model can help to better understand their role in adipogenesis and even after fat grafts implantations.

15.5.2.2 Adipose macrophages A variety of immune cells can be found in the adipose tissue, especially macrophages, the percentage of which differs from 10% in lean humans to 40% in obese humans [131]. This percentage correlates with the increased adipose tissue secretion of proinflammatory molecules, maybe contributing to the pathophysiological consequences of obesity (e.g., insulin resistance, type 2 diabetes) [132]. So far, no in vitro 3D model exists to better understand this process. Only 2D coculture systems using monocytic cell lines and preadipocytes for conditioned media or direct cocultures are available. They highlighted the roles of macrophage-secreted factors on increased lipolytic activity of adipocytes, insulin resistance [133,134], and remodeling processes, by upregulating several matrix metalloproteinase [135].

15.5.2.3 Skin cells Adipose tissue is primarily located beneath the skin and is particularly involved in the skin homeostasis and aging, by its role on the cycling of hair follicles, the wound healing, or the skin protection again infections [136
138]. Adipokines secreted by the adipocytes induce skin integrity by enhancing the migration of the keratinocytes, the ECM secretion of the fibroblasts [139], and the maturation of the blood vessel [140]. On the other hand, keratinocytes also influence the subcutaneous adipose tissue homeostasis by stimulating adipocyte differentiation in vivo and in vitro through epidermal Wnt/β-catenin [141]. Moreover, Rac-dependent paracrine pathway from keratinocytes can modulate both white and brown dermal adipogenesis [142].

15.5.2.4 Muscle cells Cross talks also exist between intramuscular adipose tissue and skeletal muscle cells mechanisms. Adiposity can be linked with possible muscular dysfunctions, since mature adipocytes can be generated from myofibroblasts in the case of large

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skin wounds [143]. However, only few in vitro models have already cocultured human myotubes and adipocytes. The direct coculture in 2D showed that the secretion of altered adipokines from obese adipose tissue, notably adiponectin or MCP-1, contributes to increase the insulin resistance of human skeletal muscle cells, as a marker of type 2 prediabetic state. In 3D culture, when collagens gel containing myoblasts or obese subcutaneous mature adipocytes were cultured, an increased secretion of adipocyte cytokines was observed, leading to a decreased expression of skeletal muscle contractility complex genes and myogenesis, consequently inducing atrophy [144
146].

15.5.2.5 Cancer cells It is particularly well-known that adipose tissue can provide an energy source to cancer cells increasing their tumorigenicity by proinflammatory cytokine secretion. In the case of breast cancer, tumor cells secrete factors that stimulate adipocytes secretion and promote their dedifferentiation to increase the stem-like cell content in the stroma [147], which then enhances their proliferation and invasive properties [148]. More specifically, adipocytes are involved in the epithelial
mesenchymal transition in several cancers [149]. Multiple breast cancer cell lines cultured in human decellularized adipose tissue scaffolds had a slower proliferation rate, similar to in vivo, and increased drug resistance when compared to 2D models [150]. hMSC line isolated from breast reduction mammoplasties seeded in collagen I scaffolds showed an enhanced migration of cocultured breast cancer cell [151]. Concerning pancreatic cancer, when seeded on the upper layer of collagen type I embedded ADSCs or minced pieces of mouse visceral fat, their motility became affected by the remodeling of the collagen fibers in the tissue. ADSCs also produced dense collagen matrices at cancer invasion sites to enhance tumor progression [152].

15.5.3 Reconstruction of the different types of adipose tissues Adipocytes have extraordinary plastic properties. Four types of adipose tissues exist with possible conversion between them, but no in vitro models are available for all of them. White adipocytes are known to be the lipid storage cells which provide energy, brown adipocytes use their lipids to heat the body, yellow adipocytes of the bone marrow can act on the bone marrow cells, and pink adipocytes are involved in the production of milk in the breasts.

15.5.3.1 “White” (fat storage adipose tissue) Unique innate adipocyte characteristics can be found with functional distinctions, according to the location of the white adipocytes depots. One of the biggest differences is between the visceral and the subcutaneous fat. Subcutaneous fat depots tend to increase their fat cell number while visceral adipocytes increase their size (lipids content) and exhibit more proinflammatory gene expression and

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obesity-related insulin resistance [153]. Usually in vitro studies do not compare adipocytes from several locations and subcutaneous ADSCs are the most used as they can show at least three times more effective adipogenesis than visceral ADSC in 2D culture [154]. Only one 3D model used these two types of adipocytes to show that collagen hydrogels can help to maintain both depot-specific gene expression and functions [85].

15.5.3.2 “Brown” (heat production adipose tissue) Brown adipose tissue (BAT) uses lipids for thermogenesis in response to cold exposure via uncoupled respiration. Humans and animals have two major forms of BAT: native BAT located at distinct anatomic sites [155] and inducible BAT (also called beige or brite fat) coming from “browning” process in white adipocytes in response to endogenous or exogenous stimuli [156]. The current challenge to cure obesity is the direct in vivo conversion of white adipose tissue (WAT) into beige adipocytes, or the use of in vitro models to generate beige adipose tissue from autologous adipose
derived stromal cells and transplant them back into the patient. Few engineering models of brown or beige adipose tissue have been developed. Spheroids of white adipocytes can be induced to express the brown markers uncoupling protein 1 (UCP1) and Cidea, at a higher level than traditional 2D culture [42]. The use of a synthetic scaffold (polyethylene glycol diacrylate) with varying levels of stiffness increased the metabolic activity and induced the expression of UCP1 and Cidea of rat and human white ADSCs [157]. And 3D spheroids derived from mouse BAT have a maintained expression of BAT markers [158].

15.5.3.3 “Yellow” (bone marrow adipose tissue) Another adipose tissue, which is also a unique adipose depot, is in the bone marrow where bone marrow adipocytes are involved in the signaling with local and distant cells. In obesity, both bone marrow adipose tissue (BMAT) and WAT increase, while in starvation/anorexia condition, WAT decreases and BMAT increases [159]. Only one model exists currently on the in vitro development of BMAT. Human and mouse bMSCs seeded in silk scaffolds can be stably cultured for 3 months in vitro with good viability and adipogenesis, along with less inflammatory phenotype than in 2D culture. When myeloma cells are added in this model, delipidation of the BMAT adipocytes occurred, highlighting the model interest to elucidate this specific signaling and identify novel therapeutic targets [160].

15.5.3.4 “Pink” (breast adipose tissue) During pregnancy and lactation, white adipocytes of mammary glands convert reversibly to epithelial cells to compose the milk-producing glands, being called pink adipocytes. In vitro 2D culture studies and explant experiments using white or brown adipocytes have confirmed their ability to differentiate into alveolar cells or mammary glands’ myoepithelial cells [161]. But no 3D engineered

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model exists so far and there is a special need of this particular adipose tissue for breast reconstruction surgery for instance.

15.6

Conclusion

Several artificial or natural biomaterial choices are possible in adipose tissue engineering approaches for in vitro models or regenerative purposes. However, significant improvements have first to be done on vascularizing the tissues to allow larger functional and maintained constructions for soft tissue regeneration. This scale-up step of the engineered adipose tissues will be challenging to get suitable volume, while ensuring the proper nutrients diffusion. 3D bioprinting of the tissues seemed a solution to overcome the issues, printing large tissues, including vasculature. But despite all the bioprinting progress over the years, many challenges still remain unsolved concerning the long duration of the printing process, the bioink used or the still low viability, and cell differentiation observed. Thus, so far, no bioprinting process was published yet using the very fragile mature adipocytes. Concerning the drug screening application of the engineered adipose tissues, future work adding other adipose cell types or surrounding tissues should then allow to move closer to physiologically relevant models. The current interesting solutions are the organs-on-a-chip that connects multiple organs to create a single integrated system of healthy or disease models [162]. This technology can allow the full communication and growth factors diffusion between all the related tissues found around the adipose tissue (muscles, skin, cancer cells, etc. see Fig. 15.3). The best final model could even add brain-on-chip for autonomic nervous system secretions participation, along with nerve innervation of the tissue [163].

Figure 15.3 Ideal reconstructed adipose tissue for drug screening application. Organs-on-a-chip system allows the reconstruction of an interconnected in vitro adipose tissue also involving the surrounding tissues, here muscles, skin and cancer cells. This model can simulate the human communications between these tissues in healthy or disease conditions. Source: Credit: Fiona LOUIS.

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Blood
brain barrier tissue engineering

16

Agathe Figarol and Michiya Matsusaki Department of Applied Chemistry, Graduate School of Engineering, Osaka University, Suita, Japan

16.1

Introduction, specificities of the blood
brain barrier

16.1.1 A highly selective barrier Evolution has pressured to protect the brain of vertebrates from external aggressions. Thus three different barriers protect the central nervous system (CNS): the arachnoid matter, the blood
cerebrospinal fluid barrier, and the blood
brain barrier (BBB). The first two barriers wrap the whole CNS and have extremely low surface area compared to the BBB. Blood vessels but also the c. 100 billion blood capillaries in an adult brain form indeed a network of around 600 km length for 12
18 m2 surface exchange area [1]. The BBB ensures the protection of the CNS by controlling the brain microenvironment. It notably restricts the passage of psychoactive drugs (from caffeine to cocaine), but also stops the transport of the great majority of pharmaceutical drugs: 100% of large molecules and more than 98% of small molecules [1]. The transendothelial electrical resistance (TEER), which measures the impermeability of biological barriers, can reach around 1900 Ω cm2 in frog brains, which is similar or superior to that of a tight epithelium [2]. Later studies have predicted up to 8000 Ω cm2 in rat brains but measured more moderate values from 30 to 5900 Ω cm2 [3,4]. BBB disruption can occur following strokes and a number of neurological diseases: multiple sclerosis, Alzheimer’s and Parkinson’s for example [5
7]. Impairment in the blood circulation and exchange regulation can lead to further tissue damages and necrosis [8]. In these cases, angiogenesis is necessary for the brain to trigger its limited ability to repair. Engineered BBB vascularized tissues could thus support full brain function recovery.

16.1.2 Endothelial cells, pericytes, and astrocytes Three cell layers overlap to form the BBB and impart its high barrier function: brain endothelial cells, pericytes, and astrocytes feet (see Fig. 16.1). The brain endothelium resembles the endothelium from the rest of the human body, but interstitial spaces are 50
100-fold tighter and form a nonfenestrated cell layer [9]. The expression levels of enzymes and proteins from endothelial cells in BBB differ Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00022-2 © 2020 Elsevier Ltd. All rights reserved.

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Figure 16.1 Schematic section of the blood
brain barrier. Source: Credit: Agathe Figarol.

also, especially for tight junction proteins and transporters [10,11]. In the numerous capillaries, perivascular cells do not express smooth muscle actin like in the arteries and arterioles. They are identified as pericytes (or Rouget cells). The CNS and retinal vasculatures are regarded as the most pericyte covered [12,13]. The coverage in surface ratio (pericytes/endothelial cells) is about 22%
37% depending on the region of the cerebral cortex [14
16]. Pericytes secrete a wide range of vasoactive agonists to support the contractile function [17
19], as well as structural constituents of the basal membrane [13]. Evidences of the interdependence of cells’ secretion on the cellular organization and maturation of the BBB were exposed, the hypothesis being that pericytes decrease endothelial proliferation and favor their differentiation [13,17,18,20]. Astrocyte end feet constitute the third layer of the BBB. Modern microscopy technics allowed to confirm the presence of glial cells with several stem branches, some of their endings (or feet) forming a lacework of fine lamellae covering endothelial cells and pericytes, as foreseen by Virchow and Golgi [21
23]. The surface coverage ratio (astrocyte/endothelial cells) was measured at around 63%, allowing a total coverage of the endothelium surface of 99% by both pericytes and astrocytes [14]. Astrocytes provide biochemical support to the endothelial cells, which in return secrete factors of differentiation for astrocytes [23]. Along other functions, the astrocytes’ long and thin stem branches are assumed to interact closely with neurons for the organization and operation of synaptic connections [24].

16.1.3 Extracellular matrix The extracellular matrix (ECM) plays a crucial role in tissue organization and structural support to cell interactions. In a healthy adult brain, it represents 20% of the total volume [25] and results from two kind of matrices: the (peri-)neuronal interstitial matrix and the endothelium basement membrane [26,27]. The latter engulfed the basal side of the endothelial cells, the pericytes, and astrocyte feet. The brain ECM contains relatively small amounts of fibrous proteins (mostly collagen type

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IV, laminin, and fibronectin) but a dense network of proteoglycans and other glycoproteins (mostly heparin sulfate proteoglycans as perlecan, and hyaluronan in the basement membrane) [26,28,29]. Diffusion, cell migration, and differentiation through the BBB are constrained by its chemical composition, tortuosity, and thickness (20
200 nm) [30]. ECM is taking a more and more important part in BBB engineering switching from two-dimensional (2D) to tridimensional (3D) models.

16.1.4 First attempts in blood
brain barrier models BBB models were primarily focusing on drug delivery assays; hence, the priority was to retrieve an endothelium barrier. In the early 1990s brain endothelial monolayers were simply grown on top of the membrane of cell culture inserts [31]. Models were then enhanced by introducing one or two other cell types in the inserts on the other side of the membrane or on the well bottom. Coculture with pericytes, astrocytes or in different models, glial cells, neurons or neuronal precursors can increase the TEER levels, the expression of tight junction proteins and restrict the barrier permeability [32
35]. This coculture model on insert has been widely documented for drug transport and remains the most commercialized BBB model. Correlation to in vivo rat experiments reached levels of 0.89 or 0.99 in Nakagawa and Josserand’s studies, respectively [36,37]. 2D models reached a step further with the development of microfluidic devices. TEER values in dynamic conditions can be about 100 times higher than in static conditions [38,39]. Studies introducing ECM as coating between cell layers [40] finally paved the way for more complex 3D models that we will detail here.

16.2

Spheroids

Some cell types, such as astrocytes and pericytes, have the capacity to gather in a 3D spherical shape when cultured on low-attachment support or as hanging drops, allowing simple 3D models [41,42]. In ECM compounds enhanced medium, when in coculture with endothelial cells, they tend to arrange into three defined layers: a surface recovered by endothelial cells, a supporting layer of pericytes, and astrocytes underneath constituting the spheroid core (see Fig. 16.2A). Cho et al. and Urich at al. have obtained such results with agarose gel and methylcellulose [42,43]. Protein expressions and response to toxic compounds shifted when compared to 2D culture. Moreover, Cho’s study has shown, for a range of 18 compounds, permeability results that recapitulated in vivo data. Some studies have included neurons or neural stem cells, microglia, and oligodendrocyte progenitor cells to mimic the entire neurovascular unit [44]. Although it opened new perspectives for a more complex BBB modeling, the 3D conformation was less defined than for more traditional three-culture spheroids.

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Figure 16.2 Schematic representations of 3D BBB models: (A) spheroid, (B) templated growth on inorganic materials, (C) templated growth on inorganic material for endothelial cells and in hydrogel for astrocytes and pericytes, (D) templated growth on hydrogel, (E) templated growth on hydrogel with sprouts. In red: endothelial cells, in green: pericytes, in blue: astrocytes, arrow: direction of shear flow. 3D, Tridimensional; BBB, blood
brain barrier. Source: Credit: Agathe Figarol.

16.3

Templated vessels’ growth

16.3.1 Rigid channels The cylindrical geometry of blood vessels and capillaries imposes a curvature constraint on endothelial cells, which influence their differentiation and orientation. Microchips in rigid polymers, such as polydimethylsiloxane or polypropylene, allow to grow endothelium at the internal surface of a carved cylindrical or rectangular channel (see Fig. 16.2B and C). The chip generally contains a second compartment with supporting cells, separated from the endothelial cells by a porous membrane. Shear flow can be applied to both compartments distinctly [45,46]. The support cells compartment can be filled with a hydrogel (most often collagen type I) to better mimic the brain microenvironment. Astrocytes, pericytes, and neurons seeded in such gel were more likely to resume a 3D conformation [47]. Systems can be made further complex with the addition of supplementary gel or medium channels [48]. However, in any case, channel diameters seem to be the limiting criterion. Cucullo’s study has indeed shown TEER and permeability values comparable to in vivo for cylindrical channels of several hundred micrometers, but they were drastically altered for thinner channels with dimensions closer to capillaries [45]. Direct exchanges between cells are furthermore still limited, and rigid channels do not allow vasodilation or constriction of the templated vessels.

16.3.2 Extracellular matrix channels Collagens composes usually 25% of human tissue mass with collagen type I being the main type present in the full human body and second in the brain [28]. Cylindrical channels can be shaped in gelled collagen type I for direct endothelial cells seeding (see Fig. 16.2D). After 2
3 days of static monoculture, tight junctions, evidence of endothelium formation, could be observed [43]. Coating the

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collagen type I channel with fibronectin and collagen type IV stabilized this shaped endothelium under light flow for 1 week [49]. The system permeability was equivalent to 2D endothelial monolayer on top of a similar gel. TEER values however were decreased by at least half. In a latter study by the same team, it was shown that the endothelial cell type was of high influence on these values and that iPS (induced pluripotent stem) cell
derived human brain microvascular endothelial cells gave more satisfying results that usual human umbilical vein endothelial cells (HUVEC) [50]. Several channels can be carved in the same chip and still reach a stable endothelium over 3 weeks [51]. Tourovskaia et al. obtained satisfying results with additional coculture of astrocytes and pericytes embedded into the collagen type I gel [52]. After 5 days the support cells migrated near the endothelial cells, and the average permeability to fluorescent tracers was in the range of isolated mammalian venules. Herland et al. found later that astrocytes adhered partially to the basal side of the endothelial vessel-like, in the vicinity of which, basal membrane proteins were secreted [53]. Their system was however stable for 4
5 days only for channels with high diameters; under 300 μm, the gel collapsed. Partyka et al. published a similar study reaching a smaller 180 μm diameter channel with the collagen type I hydrogel enriched with Matrigels (Section 13.1 will describe further this matrice) and hyaluronic acid that enhanced astrocytes spreading and activation [54]. These diameter ranges make ECM channels resemble blood vessels or venules. Although a step further for toxicity and drug delivery applications, they do not retrieve vasculogenesis or angiogenesis of small capillaries as needed for tissue regeneration.

16.4

Sprouts and guided capillaries growth

Tourovskaia et al. [52] attempted to enhance their ECM channel system with the formation of capillary-like sprouts (see Fig. 16.2E). The addition of angiogenic factors—vascular endothelial growth factor (VEGF), basic fibroblast growth factor, and phorbol-12-myristate-13-acetate—induced indeed small branchings of endothelial cells covered with pericytes and basal membrane proteins after only 1 day of treatment. Unfortunately, this was only achieved with HUVEC and not with brain microvascular endothelial cells. Another study was carried out with HUVEC and mouse immortalized pericytes [55]. The addition of pericytes seemed to increase the length of the newly formed sprouts but lower their diameter, which is coherent with its supposed function of angiogenesis regulation. The network expansion and interconnection seemed however to be limited. In a different approach, brain microvascular endothelial cells growth can be guided around alginate microbeads (B150 μm diameter) in collagen type I gel [56]. It is arguable whether this forced organization can be seen as BBB capillaries, especially as fibroblasts were used as support cells, but the change of perspective is interesting.

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Capillaries self-organization

16.5.1 Capillaries self-organization on top of Matrigels We have already seen that Partyka et al. used Matrigels as additive in their collagen-based gel [54]. Matrigels is the trade name of a mouse sarcoma extract
based matrix. Its composition is complex and batch dependent. It comprises structural proteins (such as collagen type IV, heparin sulfate proteoglycans, and nidogen) and numerous growth factors. When used as a thick coating, it favors morphogenic organization of endothelial cells as tubules, or capillary-like structures (CLS), some of them containing lumens (Fig. 16.3A) [57]. If CLS formation is widely referenced with noncerebral endothelial cells, it is not so common to find equivalent results with cerebral ones. CLS obtained after 48 hours culture on Matrigels were indeed found less defined with immortalized microvascular endothelial cells from human brain than from human foreskins (both in coculture with immortalized human neural stem cell) [20]. The addition of other growth factors, such as those derived from cultured endothelial progenitor cells, can increase the CLS definition for rat brain endothelial cells [58]. The coculture of the three BBB cell types (primary brain endothelial cells, astrocytes, and pericytes) can also enhance the self-organization in CLS on Matrigels [42,59]. In Itoh’s study, if the CLS formation on Matrigels was very fast (CLS of about 200
300 μm length and 5
10 μm diameter in a few minutes to 6
8 hours), its degradation was also equally rapid [59]. After 24 hours of monoculture, endothelial cells started to present some signs of apoptosis and detached. The protective effect of astrocytes and pericytes allowed CLS maintenance up to 18 hours longer. This study also showed a 3D organization close to in vivo, with pericytes migrating close to endothelial cells and the major part of astrocytes situated near the CLS. Very few astrocytes feet were observed, but they formed some pseudopodia that could be the sign of cellular activation. These mitigated results have to be put into perspective as the support cells have been added to the culture system after the organization of endothelial cells as CLS. Goodwin suggested that CLS of noncerebral endothelial cells are more

Figure 16.3 Schematic representations of 3D BBB CLS models: (A) CLS on top of Matrigel, (B) oriented CLS growth in microchips, (C) self-organized CLS in hydrogel. In red: endothelial cells, in green: pericytes, in blue: astrocytes, in brown: fibroblasts, open arrows: direction of CLS growth. 3D, Tridimensional; BBB, blood
brain barrier; CLS, capillary-like structures. Source: Credit: Agathe Figarol.

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uniform and have a better 3D organization when cells are directly plated within a Matrigels matrix [57]. Results for brain endothelial cells are awaited. Pham et al. proposed a kind of iPS-derived endothelial cells spheroids embedded after formation in a Matrigel matrix [60]. Few capillaries were formed. Interestingly, they conducted an in vivo transplantation in a rat brain. It remained difficult to draw a proper conclusion from their results, but it showed the potential to push toward tissue regeneration from BBB modeling.

16.5.2 Capillary self-organization on microchips For 3D CLS formation in microchips, hydrogels with a more controlled composition are favored. Kim et al. [61] showed that HUVEC cocultured with pericytes could form CLS in fibrin gel. HUVEC and pericytes were seeded along a central channel filled with fibrin gel, a third channel on the opposite side containing fibroblasts. After 3
6 days, HUVEC and pericytes migrated into the central gel, HUVEC forming CLS toward fibroblasts, and pericytes forming a support network in its vicinity (see Fig. 16.3B). Similar experiments were carried out with brain microvascular endothelial cells cocultured with human mesenchymal stem cells, with or without neurons in the third channel [62]. Results were satisfying with fibrin and fibrin/Matrigels gels but not in fibrin/hyaluronan gels that indicate a strong impact of the ECM. Endothelial cells can also be seeded directly in the central gel channel. In this way, the orientation of CLS can be controlled: for example, HUVEC CLS in fibrin gel were seen to orientate first toward a fibroblasts seeded channel, then back on the opposite direction after seeding of neurons and astrocytes in a third channel [63]. In collagen type I gel, iPS-derived endothelial cells were also seen to organize in CLS interacting with spheroids of neural stem cells and without any additional support cells [64]. It is only very recently that such 3D CLS in microchips were developed using the three BBB cell types [65]. Human iPSinduced brain endothelial cells, primary human pericytes, and primary human astrocytes were embedded into fibrin gel and put into the central channel of a microchip. In 4
5 days brain endothelial cells formed CLS. In monoculture the capillary network started to degrade after 7 days, while it remained stable and more defined with the addition of pericytes and even more with pericytes and astrocytes. Similarly, lumens were found more circular in coculture, tight junction expression was increased, and permeability to fluorescein isothiocyanate linked (FTIC)-dextran was reduced (after closing up the system with monolayer of endothelial cells seeded on the side channels). One advantage of the microchip is the possibility of continuous medium perfusion, which would certainly be the next step for these systems. For regenerative applications, however, the use of devices such as microchips may be a limitation when thinking about in vivo implantation.

16.5.3 Capillaries self-organization in device-free hydrogels Let us go back to 1997 to find the first published study on device-free 3D selforganized CLS BBB model. Ment et al. embedded primary beagle brain

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Figure 16.4 Schematic representations of 3D BBB CLS in a device-free technique. Confocal image obtained after 7 days of culture of a hydrogel plug on a 24 well plate, and CD31 immunofluorescence of immortalized brain microvascular endothelial cells. 3D, Tridimensional; BBB, blood
brain barrier; CLS, capillary-like structures. Source: Credit: Agathe Figarol.

microvascular endothelial cells with primary neonatal rat astrocytes in a simple collagen type I gel [66]. About 200 μL drops of the mix were deposited on a petri dish. After polymerization for 20 minutes at 37 C, the drops formed gel plugs. Medium could then be added without damaging the plugs. By day 6, endothelial cells organized as weak but observable CLS, astrocytes migrated nearby, and crucial extracellular proteins (laminin and fibronectin) were expressed. Ahmad et al. used a similar technique to form collagen type I plugs containing rat brain endothelial cells, primary rat astrocytes, and primary rat pericytes [67]. Cyst-like structures were first formed, cells then migrated out of the plugs, connected to the nearby plugs, and finally started to arrange into CLS after 4
6 days. Angiogenic factors (such as VEGF) were thought to be responsible for the CLS formation as their levels gradually increased up to 3 days then remained stable during the CLS organization. Osaki et al. conducted a device-free macroscale assay in parallel to their assay on microchips but with HUVEC and neuronal spheroids [64]. Recent work in our laboratory (Figarol et al., unpublished, Figs. 16.3C and 16.4) derived from a previous vascular model [68] achieved the formation of device-free 3D selforganized CLS in a ECM
collagen microfiber mix with human immortalized or iPS-derived brain microvascular endothelial cells, human immortalized astrocytes and pericytes [69
72]. Drops of the cell
hydrogel mix were gelled on culture well plate (30 minutes
1 hour at 37 C), medium was added and the system was cultured for 1 week. The ECM composition was optimized and different sizes of drops with well-defined and opened CLS could be obtained, as well as thicker gels in culture inserts. Further optimizations, function validation, and long-term culture are however still needed for translational research.

16.6

Current challenges in translational research

BBB engineered models aims to be a step “from bench to bedside” for the development of therapies, diagnostics tools, or medical procedures supports. Their first application is the in vitro drug transport screening. The BBB poses indeed a

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tremendous hurdle for systemic delivery of drugs for the treatments of neurological disorders. In vitro cellular models represent a more effective, cost-reductive, and ethical mean to investigate drug delivery through the BBB. They are especially more relevant for fast screening than the alternative animal models or postmortem human brain tissues. They ease the research for drug transport improvements through innovative use of receptor-mediated nanocarriers, extracellular vesicles and physic method such as ultrasonic treatment [73]. Pharmaceutical companies prioritized the robust 2D on inserts models so far, but the demand for 3D engineered model intensifies in order to reduce the gap between in vitro and clinical studies. The biggest challenge for this application is to calibrate the models with validated BBB standards (caffeine, Rhodamine 1-2-3, Lucifer yellow, etc.) and part with fluorescence analysis for permeability assay (as it was done in Campisi’s study for example [65]). BBB models’ second translational application is the development of new drugs for diagnostic and treatment purpose. Pathological models are an uprising challenge. The BBB can be dysregulated by age, genetic, environmental factors; and its role is still not fully grasped in the cases of brain cancers, autism, diabetes, depression, and other brain diseases [74]. The efforts already made in understanding BBB angiogenesis for the sprouting or self-organized capillary models could bear fruit, for example, for studying recanalization after strokes. In the case of the BBB the use of primary cells is problematical as it requires extremely invasive procedures. iPS cells circumvent this issue and could allow personalized therapy development. Genetic related conditions could be retrieved in BBB patient
specific models. However if a few BBB models were engineered with iPS-derived brain microvascular endothelial cells, obtaining long-term stable culture of iPS-derived pericytes and astrocytes is still challenging. BBB models should furthermore balance between microfluidic systems, and scaffold-free flexible systems. Flow has indeed been seen to affect cell differentiation, barrier function, hence drug transport and efficiency. However, microchips and other flow system restrain the shape of the BBB model and limit the perspectives for large tissues and implantations.

16.7

Implantation prospects

Angiogenesis favors neuron development and improves recovering in stroke patient. However, simple intakes of angiogenic growth factor such as VEGF were linked with BBB leakage, brain edema, and neurological troubles [40]. Injection of an acellular hydrogel in the damaged area was proposed as an alternative. Nih et al. [75] developed a gel composed of hyaluronic acid enriched with VEGF-loaded heparin nanoparticles. It showed promising results on mice with mature capillary network formation and partial behavioral recovery. Therapeutic effect was indeed significant for the highest dose of bound VEGF. Unfortunately, even if somehow modulated by the heparin nanoparticles, this high VEGF dose still led to inflammatory effects. This approach sounds exciting as it does not require upstream tissue

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engineering. However, the latter may be essential to get rid of the inflammatory outcomes and to promote faster neurogenesis in the vascularized tissue. Implantation assisted by hydrogels of some brain cells starts to be carried out in vivo [76], but very rare results on implantation of engineered BBB tissue in vivo have been published yet. We have cited Pham’s study that implanted a kind of organoid with human iPS-derived endothelial cells into a rat [60]. They showed indeed better results with prevascularized organoid than nonvascularized control one (n 5 1). As this research is still at an infancy stage, both acellular and tissue implants should be studied in parallel to achieve full comprehension of the angiogenesis in the very specific context of the BBB.

16.8

Conclusion

Vascularization is an essential step to regain a fully functional tissue. Specificities of the BBB make the modeling of this tissue very challenging. However, along the evolution of biomaterials, BBB models are becoming more and more relevant. Lately, the optimization of hydrogels as ECM allowed the development in vitro of 3D self-organized BBB capillary networks. The increasing speed of the research in this area could lead to more realistic models in the next years. Implantation trials for regenerative brain tissue seem like a reachable dream. Reliable BBB engineered models must be prepared to face a much more complex environment. Gathering expertise from immunology, neurology and biomaterials fields could be the key. A full “brain on a chip” would be a tremendous asset for translational medicine, and prediction of the migration and organization of neurons in the BBB engineered tissue after implantation.

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37. Available from: https://doi. org/10.1016/j.biomaterials.2018.10.023. [51] Kim JA, Kim HN, Im S-K, Chung S, Kang JY, Choi N. Collagen-based brain microvasculature model in vitro using three-dimensional printed template. Biomicrofluidics 2015;9. Available from: https://doi.org/10.1063/1.4917508. [52] Tourovskaia A, Fauver M, Kramer G, Simonson S, Neumann T. Brief communication: tissue-engineered microenvironment systems for modeling human vasculature. Exp Biol Med (Maywood) 2014;239:1264
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[53] Herland A, Meer AD, van der, FitzGerald EA, Park T-E, Sleeboom JJF, Ingber DE. Distinct contributions of astrocytes and pericytes to neuroinflammation identified in a 3D human blood-brain barrier on a chip. PLoS One 2016;11:e0150360. Available from: https://doi.org/10.1371/journal.pone.0150360. [54] Partyka PP, Godsey GA, Galie JR, Kosciuk MC, Acharya NK, Nagele RG, et al. Mechanical stress regulates transport in a compliant 3D model of the blood-brain barrier. Biomaterials 2017;115:30
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83. Available from: https://doi.org/ 10.1016/j.mvr.2007.05.006. [58] Di Santo S, Seiler S, Fuchs A-L, Staudigl J, Widmer HR. The secretome of endothelial progenitor cells promotes brain endothelial cell activity through PI3-kinase and MAPkinase. PLoS One 2014;9:e95731. Available from: https://doi.org/10.1371/journal. pone.0095731 info:. [59] Itoh Y, Toriumi H, Yamada S, Hoshino H, Suzuki N. Astrocytes and pericytes cooperatively maintain a capillary-like structure composed of endothelial cells on gel matrix. Brain Res 2011;1406:74
83. Available from: https://doi.org/10.1016/j. brainres.2011.06.039. [60] Pham MT, Pollock KM, Rose MD, Cary WA, Stewart HR, Zhou P, et al. Generation of human vascularized brain organoids. NeuroReport 2018;29:588. Available from: https://doi.org/10.1097/WNR.0000000000001014. [61] Kim J, Chung M, Kim S, Jo DH, Kim JH, Jeon NL. Engineering of a biomimetic pericyte-covered 3D microvascular network. PLoS One 2015;10:e0133880. Available from: https://doi.org/10.1371/journal.pone.0133880. [62] Uwamori H, Higuchi T, Arai K, Sudo R. Integration of neurogenesis and angiogenesis models for constructing a neurovascular tissue. Sci Rep 2017;7:17349. Available from: https://doi.org/10.1038/s41598-017-17411-0. [63] Bang S, Lee S-R, Ko J, Son K, Tahk D, Ahn J, et al. A low permeability microfluidic blood-brain barrier platform with direct contact between perfusable vascular network and astrocytes. Sci Rep 2017;7:8083. Available from: https://doi.org/10.1038/s41598017-07416-0. [64] Osaki T, Sivathanu V, Kamm RD. Engineered 3D vascular and neuronal networks in a microfluidic platform. Sci Rep 2018;8:5168. Available from: https://doi.org/10.1038/ s41598-018-23512-1. [65] Campisi M, Shin Y, Osaki T, Hajal C, Chiono V, Kamm RD. 3D self-organized microvascular model of the human blood-brain barrier with endothelial cells, pericytes and astrocytes. Biomaterials 2018;180:117
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6006. [72] Yamashita M, Aoki H, Hashita T, Iwao T, Matsunaga T. Effect of compounds Y on the barrier function of human iPSCs derived brain microvascular endothelial cells. Toxicol Lett 2018;295:S124. Available from: https://doi.org/10.1016/j.toxlet.2018.06.677. [73] Ochocinska MJ, Zlokovic BV, Searson PC, Crowder AT, Kraig RP, Ljubimova JY, et al. NIH workshop report on the trans-agency blood
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Tissue engineering in urology

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Elif Vardar De´partement Femme-Me`re-Enfant (DFME), Centre Hospitalier Universitaire Vaudois (CHUV), Lausanne University Hospital, Lausanne, Switzerland

17.1

Introduction

Tissue engineering (TE) is a rapidly evolving field, offering to develop tissue substitutes as an alternative therapy for the treatment of damaged tissues and organs. TE applications for human urinary system mainly aim to repair, restore, and regenerate three main organs of the system: kidney, bladder, and urethra. The urinary system is responsible for filtering out the metabolic side/end products from the blood. These products are converted into urine, produced by the kidneys, collected in the bladder, and excreted through the urethra. Renal tissue shows high level of complexity, consisted of more than 20 specialized cell types, Fig. 17.1, [1]. The smallest functional unit of the kidney is called the nephron, which is responsible for the blood filtration and reabsorption of molecules. Two kidneys consist of approximately 1
3 million nephrons in total. Each nephron has a structural part, called glomerulus, surrounded by a Bowman’s capsule, a proximal tubule, a loop of Henle, and a distal tubule. All nephrons drain into the collecting ducts within the medullary pyramids. The glomerulus is the key filtering unit of the kidney. The glomerulus has two main cell types, endothelial cells and podocytes. These cell layers are separated by a double-layered basement membrane, made of extracellular matrix (ECM) proteins, for example, collagen type IV and laminin, and heparin sulfate proteoglycans, which functions as an actual filtra˚ Stokes radius. tion barrier. Glomerulus filters plasma components below 36 A Almost 180 L of filtrate composed mainly of creatinine, urea, glucose, and amino acids, is generated by the glomerulus every 24 hours. In addition to the filtration function, kidney is an endocrine/metabolic organ, which secretes erythropoietin and prostaglandins [2]. An injured renal tissue undergoes severe complications. If the acute stage of renal injury is not repaired, this acute kidney disease (AKD) may be accompanied by other persistent complications, called chronic kidney disease (CKD) symptoms. CKD has five stages based on the severity of renal impairment, one of which is permanent kidney failure, known as the end-stage of renal disease. Both the AKD and CKD can be related to several conditions, such as malfunctioning kidneys, other organs, or systemic diseases. CKD is one of the major health problems, which is responsible for 1.5% of deaths worldwide [3]. There are millions of patients in Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00013-1 © 2020 Elsevier Ltd. All rights reserved.

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Figure 17.1 Diagram of the human kidney. The nephron, accommodating glomerulus and renal tubules, is the smallest functional unit of kidney. The unfiltered blood that is carried by afferent arterioles is filtrated in the glomerulus. The filtered blood is carried away from the kidney to the renal vein. The glomerulus filtrate, mainly consisted of water, glucose, small proteins, ions, and salts, is formed within Bowman’s capsule. The glomerulus filtrate further moves into the renal tubules. Most of these small molecules are reabsorbed back into the blood when the filtrate flows through the renal tubules. When the glomerular filtrate exists the collecting ducts, it is called urine that is stored in the bladder before it is eliminated through the urethra.

critical condition waiting for renal transplantation on the organ waiting list. Renal transplantation is a highly effective restorative treatment for patients with kidney failure [3]. However, several significant drawbacks are associated with this option, such as donor shortage, surgical morbidity, and the need for lifelong immunosuppression. TE and regenerative medicine strive to design new biomaterials to address these limitations. Urine is a slightly acidic fluid, stored in bladder, Fig. 17.2, until it is ready to be excreted from the body. The bladder is a hollow organ, located in the lower abdomen. The bladder tissue is composed of three distinct tissue layers: mucosa, submucosa, and muscularis layer. Mucosa layer contains transitional epithelium that is watertight and protects the underlying tissue layers from acidic or alkaline urine. Submucosal layer connective tissue consists of blood vessels and nervous tissue, supporting the surrounding tissue layers. Muscularis layer is composed of smooth muscle and fibroelastic connective tissues, responsible for bladder contraction during urination to expel urine from the body. The muscularis layer is also present in internal urethral sphincter, surrounding the urethral opening and holding urine in the urinary bladder. During urination the sphincter muscle complex controls the flow of urine in the urinary bladder through the urethra. Bladder is a compliant tissue that adjusts the pressure generated by the urine filling, as well as preventing urinary reflux to the kidney. If bladder tissue loses its compliance, which is defined as the increase in pressure per unit of volume (V/P), it can further result in kidney failure [4]. Therefore the bladder tissue should be able to adapt its capacity to the volume of the

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Figure 17.2 Diagram of human urinary bladder (female) and its general architecture. The urine is formed in kidney, passed through the ureters, and then transported to the urinary bladder to be stored until urination. Highly specialized urothelial cells form the urothelium layer that is a passive barrier protecting the surrounding tissues from the toxic effects of urine. The submucosa is composed of connective tissue that contains nerves and blood vessels (green cross-linked lines 5 collagens, red circles 5 blood vessels, gray lines 5 nerves). The muscularis layer is the smooth muscle layer that is responsible for structural support to the bladder. The smooth muscle bundles are oriented in longitudinal and circumferential direction. The orientation of the smooth muscle bundles allows the contraction and expansion of the bladder, facilitating physiological functions of filling and emptying. The internal sphincter muscle complex is an involuntary smooth muscle that controls the urine flow in the urinary bladder through the urethra. The serosa is a thin membrane and forms the outermost layer of the bladder.

accumulated urine. When bladder tissue is damaged due to a variety of clinical disorders, including end-stage bladder diseases, pelvic trauma, genitourinary malignancy, or congenital abnormalities, the most commonly applied treatment modality is to reconstruct the damaged bladder with gastrointestinal tissue segments. However, the use of these segments is associated with numerous complications such as metabolic disturbance, increased mucus production, infections, and even malignancy. The success rate of traditional bladder repair surgeries is limited due to postoperational risks, and thus new therapeutic approaches are needed for bladder reconstruction. TE may contribute to bladder reconstruction by developing new offthe-shelf solutions, helping to preserve normal bladder function. As yet, TE strategies can be broadly classified into two categories: cell-based therapies and

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engineered biomaterials. Cell-based therapies use the particular group of cells for implantation into a defect site to promote regeneration. These cells can be delivered into the surgery site within a carrier, often described as biomaterial, to promote appropriate tissue regeneration in the surgery site. To go beyond cell-carrier function, biomaterials should directly regulate cell differentiation and metabolism. Therefore future research on urological TE will essentially aim at designing smart biomaterials for functional tissue regeneration while overcoming on the regulatory challenges to be faced during the transition from research to the patient. Among the number of different TE approaches, 3D-printing technology may be a promising field as it allows for rapid testing and processing of biomaterials, mainly for renal transplantations that are more often needed than any other urological organs. 3Dprinting technology in combination with biological approaches holds great potential in designing functional biomaterials for urological tissue replacements. Although there are number of encouraging results, several challenges must be still overcome before their transitions to the clinics, including vascularization, functionality, unwanted immune response, and ideal animal models for testing. This chapter discusses recent emerging technologies and future prospects about TE of two urinary tract organs: kidney and bladder.

17.2

Biomaterials for urological tissues

17.2.1 Kidney tissue engineering CKD is one of the leading causes of mortality worldwide. There is no effective cure for end-stage renal failure. This stage can be manageable with dialysis for a period of time. The full renal function recovery can be retrieved through kidney transplantation. However, kidney transplantation has its own serious drawbacks, one of which is the need of immunosuppressive medications for life to prevent kidney rejection [5]. Herein, TE for new organs holds a great potential to replace diseased tissues providing alternative biomaterials for patients. Renal tissue ECM is composed of glycoproteins, growth factors, and structural proteins. Decellularized ECM scaffolds for artificial kidney engineering attract attention due to similarities to the native kidney tissue. Mostly, human or animal kidneys are used as template ECM scaffolds for kidney tissue regeneration. In particular, decellularized ECM scaffolds are obtained through the elimination of respective cells from native kidney tissue. However, removing all existing cellular material from the tissues remains a significant challenge. On the other hand, the resulting acellular structure should have inherent functionality that promotes cell adhesion, migration, proliferation, and differentiation. Decellularized ECM scaffolds have a significant advantage due to their three-dimensional structures composed of collagen and other biopolymers, which facilitate reconstruction. Their mechanical properties are yet weak due to chemical processing. Multiple research groups have explored that structural stability can be enhanced by cross-linking of decellularized scaffolds. The most commonly used cross-linkers for decellularized

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scaffolds are glutaraldehyde, 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide, and genipin. Although the scaffold degradation is delayed by cross-linking, the biggest concern for cross-linked biomaterials is the potential local toxicity of cross-linkers [6,7]. Yet the overall process of decellularization could also destroy the composition of ECM, which contributes to cell-mediated functions of the kidney, such as macromolecular sieving, solute transport, and vascularization [8]. This challenge may be overcome by scaffold recellularization, using primary endothelial, epithelial, or stem cells that are also needed to induce a functional endothelium formation and renal vessel regeneration in the decellularized ECM scaffolds. In the study of Paulo et al., decellularized porcine kidneys were recellularized with renal epithelial cells. The porcine kidneys were seeded using MS1 endothelial cells followed by decellularization [9]. In vitro blood perfusion experiment results demonstrated that reendothelialized porcine kidney improved vascular function. In another study, human endothelial cells and renal stem cells were delivered to the decellularized mouse kidneys through the cannulated renal artery [10]. The recellularized construct was then implanted into the dorsal skin of mouse. Three months after the surgeries, the implants were harvested and immunostained for specific proteins that play significant roles in glomerulus formation. The results showed that the cellular implant groups promoted functional host cell repopulation, leading to an improved function of kidney filtration. In conclusion, engineered decellularized constructs may potentially trigger functional kidney regeneration when it is implanted to surgery site. Therefore decellularization procedures indeed need to be standardized in a tissue-specific manner. Consequently, tissue-specific decellularization methods should be promptly developed, aiming to restore biofunctional properties of kidney. One favorable approach that attracts intensive attention for building kidney substitutes is the combination of kidney progenitor cells with 3D-bioprinting technology. The first patent on 3D-printing technology was reported in 1986 [11]. The technology was based on stereolithography, where 3D scaffolds could be fabricated layer-by-layer and simultaneously cross-linked with UV radiation. This layer-bylayer method allowed direct construction of complex scaffolds, which had a significant easy production advantage over traditional biomaterial fabrication techniques, for example, freeze drying, electrospinning, self-assembly, phase separation, and solvent casting [12]. More recently, extrusion, inkjet, and laser-assisted 3D printing methods have been commonly used for multilayered biomaterial fabrication [13,14]. These techniques have been further perfected to facilitate cellular and bioactive 3D biomaterial fabrication, which could be applied for number of different tissues such as the heart, blood vessels, kidney, and liver [14
16]. 3D bioprinting allows tailored design of complex 3D structures using several natural biopolymers, such as gelatin, collagen, fibrin, alginate, and hyaluronic acid. Although some of these materials are highly effective in supporting cell viability and growth compared to synthetic ones, a limitation yet is that the mechanical properties of these natural biopolymers are not always compatible with tissue regeneration process [17]. Therefore researchers tend to cross-link or blend natural biomaterials, for example, silk, alginate, chitosan, collagen, fibrinogen, gelatin, hyaluronic acid with synthetic ones, for example, poly(ethylene glycol), poly(ε-caprolactone), poly

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(glycolic acid) (PGA), poly(lactic acid) (PLA), and poly(lactic-co-glycolic acid) (PLGA) to reinforce their mechanical properties for soft tissue applications as well as enhance their printability [18]. Currently, bioprinting of a large-scale functional kidney tissue is not possible. Thus the latest 3D kidney tissue printing efforts have been directed to a bottom-up approach based on assembly of engineered miniature kidney units to obtain a full kidney substitute. These miniaturized units can be also used for modeling purposes. The organson-a-chip technology has been an emerging trend in imitating the complex function of native tissues in vitro and ex vivo [19]. Promising advances have been made in implementing this technology to TE of the kidney in the past a few years. Briefly, miniature components of nephron, which is the structural and functional unit of kidney, can be easily grown in plates using organs-on-a-chip technology. This technology can be successfully applied in the field of drug toxicity testing and renal replacement therapies [20]. To this end an interesting study reported by Homan et al. in 2016, where a fibrin
gelatin-based hybrid proximal tubule was printed on perfusable chips to recapitulate in vivo kidney function [21]. Unlike 2D culture results that fail to elucidate nephrotoxicity and drug-induced pathologies in kidney, 3D-printed artificial tubules showed significantly improved in vitro coculture results. A functional tissue-like epithelium formation was observed when the cells were cultured on 3D-printed proximal tubules. Obtained results look promising because kidney injuries are often associated with damage in renal proximal tubules [22,23]. Therefore creative 3D-bioprinted organ designs can be marked as an emerging key-enabling technology for the translation of new functional biomaterials to the clinics in the near future. In addition, organoid technology that is defined as unique self-organization process of miniaturized organs in specific 3D cell culture systems enables appropriate 3D models mimicking complex in vivo function, which is not achievable with the currently available preclinical animal models [24]. Numerous studies have shown that the organoids have become one of the most interesting research topics in the tissue-engineering field in the past decade [25
27]. Organoid systems provide realistic disease models, which can reveal important biological mechanisms and facilitate the high-throughput screening of new drugs. This is due to the fact that organoid cell culture methods include different stem cell types, such as embryonic stem cells and induced pluripotent stem cells (iPSCs), which have the ability of self-organization and recapitulating normal tissue function, due to their inherited developmental cues. Thus in the future the animal testing may be replaced with organ-on-a-chip platforms along with organoid technology. For example, a commercially available 3D-bioprinted kidney, ExVive, showed an ideal environment to study nephrotoxicity tubular transport of xenobiotics, proteins, and ions. ExVive artificial kidney contains layers of primary renal proximal tubule epithelial cells, primary renal fibroblasts, and endothelial cells. It was demonstrated that ExVive tissues remained viable and functional for 28 days [28]. Followed by these striking advancements in 3D-printing technology for the last 30 years, a compelling idea based on 4D bioprinting of biomaterials has been further attracted attention due to its great potential to fabricate stimuli-responsive 3D structures [29].

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The use of stem cells or kidney progenitor cells in biomaterial design is still considered as a viable option for functional kidney regeneration [30
32]. This is due to their potency to repopulate different cell types within the kidney. However, a major challenge of using harvested stem cells is their impaired differentiation ability into the appropriate phenotype in vitro [32]. Recent trends in kidney TE show that the ongoing organoids studies in combination with 3D-printing technologies could someday lead to the generation of functional ex vivo artificial kidneys that can be applied to the patients as gold standard treatment. These advancements are stepping-stone achievements and bring artificial scaffolds a step closer to the clinical use. Bringing structural complexity in scaffold design, rather than decorating them with biological signals, can facilitate the fasttrack transition of these scaffolds to the bedside, bypassing regulatory challenges. Finally, although several preliminary attempts were somewhat encouraging, there is still a long way to go before using off-the-shelf biomaterials for full kidney regeneration.

17.2.2 Bladder tissue engineering Reconstructive bladder surgery is commonly performed on patients with congenital disorders, trauma, bladder dysfunction, and cancer. Implantation of bowel segments, derived from gastric, small or large bowel tissue, is recognized as the gold standard available treatment when bladder augmentation or replacement is needed. However, serious complications, including metabolic anomalies, mucus production, urinary tract infection, and stone formation, can be observed when these patients receive bowel segments as replacement [33]. Therefore TE proposes viable alternative treatment modalities for bladder reconstruction. Several synthetic polymers, such as PLA, PGA, PLGA, teflon, and silicone, have been used in bladder repairs due to their versatile fabrication techniques with controllable mechanical properties, degradation rate, and morphology [34]. However, when they are implanted into patients, they are not very well tolerated in the body, largely due to their limited biological functionality. Very first synthetic bladder scaffold based on a molding technique was published in 1999 [35]. The synthetic bladder scaffold was made of a biodegradable PGA coated with PLGA 50:50 polymers, reconditioned using native bladder tissue urothelial and smooth muscle cells. The cellular scaffolds showed adequate functional and anatomical properties in a canine model, which had been shown for the first time in tissue-engineering field. In the follow-up study a composite scaffold, composed of PGA and decellularized collagen matrix, was tested in patients. In the study, all patients had undergone cystoplasty due to some problems resulted from their poor bladder compliance. Urodynamic studies demonstrated that the patients’ bladder capacity, which was measured for 5 years at different time points, was greatly increased [36]. This study was full of promising results in clinics that showed the great potential of using hybrid biomaterials in new implantable systems for bladder TE. Hybrid biomaterials that are obtained by blending synthetic polymers with natural biomaterials have been the point of interest of many research groups [37,38]. Besides the structural complexity intended to

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have in biomaterial design, biomaterial functionality is another important design attribute for contemporary scaffolds. Biomaterials that are considered as optimum for bladder tissue-engineering applications aim at mimicking the elastic properties and urothelial permeability functions of native bladder tissue. For example, decellularized ECM animal tissue matrices that are obtained from bladder submucosa, small intestine submucosa, and gallbladder are commonly used scaffolds in the repair of bladder defects due to their mechanical and biological similarities to the native bladder [36]. Numerous studies showed that ECM scaffolds could accelerate neovascularization and thus promoting regeneration [39,40]. Kikuno et al. demonstrated that the addition of nerve growth factor and vascular endothelial growth factor (VEGF) to decellularized bladder matrix enhanced angiogenesis and neurogenesis in a neurogenic rat bladder model [41]. Nevertheless, one of the unavoidable problems largely due to repeated surgeries in the implant area is to obtain inappropriate local reorganization of host cells and chronic inflammation in the implant site, leading to scar tissue formation. Scar tissue is a nonelastic and fibrotic tissue, whose dimension and mechanical properties are mismatching with the natural tissue in the implant site. Therefore biotechnological tools are required to tackle this problem, facilitating complex scaffold designs for tissue-engineering applications. Scaffolds can be functionalized to promote regeneration. The use of growth factors in scaffolds appears to be a promising area of research in bladder regeneration research [42]. The growth factors are bioactive molecules, triggering a set of cellular activities for complete tissue regeneration when they are incorporated to scaffolds [43]. They may further accelerate appropriate regeneration during reconstruction in the implant area and may hinder the host tissue fibrosis [44]. Consequently, the current approaches in the design of biomaterials have been directed on developing biologically active biomaterials. Another important point is to manage growth factor release from the scaffolds in a controlled manner, which is highly crucial to avoid initial burst effect. The burst release of a bioactive molecule is mostly not well tolerated in organisms [45]. One approach to prevent the burst release of these molecules is to engineer microor nanoparticles, loaded with different bioactive molecules [46,47]. Several combinational approaches have already demonstrated promising results for micro- or nanoparticles, acting as a reservoir for controlled delivery of different molecules. For example, Jiang et al. proposed a drug-delivery system consisted of decellularized pig bladder acellular matrix (BAM) and PLGA nanoparticles dispersed in a thermosensitive gel [48]. In the study, PLGA nanoparticles were loaded with VEGF and basic fibroblast growth factor to promote tissue regeneration in a rabbit bladder model. A new highly vascularized tissue formation akin to native bladder was observed at the implant site when the bladder was reconstructed with growth factor
loaded scaffolds, whereas a poor regeneration was noticed at the implant site that was reconstructed with pristine BAM scaffolds. Several exogenous biomolecules such as cell-recognizing peptides, structural proteins, and growth factors, can be firmly conjugated to different biomaterials using an advanced bioengineering technique, which is known as recombinant protein engineering. It is a strong biotechnology tool, using DNA cloning techniques. Recombinant protein engineering methods allow specific modifications in the

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biomimetic scaffolds, which make them highly functional biomaterials. There are two FDA-approved recombinant growth factor-based bone substitutes, INFUSE (recombinant human bone morphogenetic protein-2, Medtronic, Minneapolis, MN) and OP-1 (recombinant human bone morphogenetic protein-7, Stryker Biotech, Hopkinton, MA) in the market. They have both shown to induce new healthy bone tissue when they were implanted to patients [49,50]. Yet, specific regulatory challenges in relation to the transition of biological biomaterials from bench to bedside remain. Briefly, decellularized biomaterials can also cause several problems due to their harvest origin, such as disease transmission, allergies, or intense immune stimulation. Perhaps, these biomaterials can be more favorable when the fabrication costs are lowered. This can be realized through embracing advanced innovative decellularization methods over conventional methods. Most recently, an extensive research has been conducted in the field of additive manufacturing and bioinks for TE applications. As discussed in the previous section, 3D-printing platform allows the controlled deposition of different biomaterials layer by layer, enabling the fabrication of unique scaffolds with varying compositions, bioactive factors, and cells that can promote functional regeneration in the implant area. A novel digital light processing
based 3D bioprinter, allowing patient-specific decellularized bioink production, was successfully described in a recent study [51]. In the study, both decellularized heart ventricle scaffolds and liver scaffolds were printed as fibers with a thickness of 250 μm using a customized high-resolution 3D printer. The printed scaffolds were then recellularized using human-iPSC (hiPSC)
derived cardiomyocytes and hiPSC-hepatocytes. It was shown that microstructural patterning promoted stem cell differentiation toward the relevant phenotypes. Therefore kidney-specific bioinks hold a great promise in enabling production of artificial bladder tissues, maintaining both living cells and native bladder ECM components. The imminent bladder TE advancements will benefit from these two evolving advanced technologies in the foreseeable future. Besides, 3D-printing platform can combine with computational medicine providing researchers with mechanistic models to predict many essential experimental parameters, such as drug release rate and the simultaneous action of several drugs [52]. Furthermore, these variables for scaffold fabrication can be programmable depending on the patient needs. Referring to the rapid advancement of technology worldwide, these model assumptions applied new biomaterial designs may reduce the number of experiments in the initial phase. Therefore this promising tool seems to gain more attention in the design of future biomaterials. Bioreactors guide the development of tissue-specific biomaterials. They can provide physiologically relevant environmental factors, such as pH, temperature, and gas exchange, during scaffold culture [53]. Particularly, a tissue bioreactor can imitate normal physiological conditions of bladder, for example, filling and emptying, which leads to a functional tissue development after the implantation. It is evident that preconditioning of scaffolds has great advantages, one of which is that it provides wellorchestrated microenvironment, which promotes biomaterial maturation before implantation. A series of bioreactor types have been developed for TE applications (cf. Chapter 19: Platelet-rich plasma in tissue engineering), such as compression,

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perfusion, stress-strain, flow, and electrical stimulated bioreactors [54
58]. Tissue bioreactor systems have been used for the reconstruction of different types of 3D engineered tissues, including tissues, cardiac, liver, and bone, cartilage [59
62]. Number of preclinical studies showed that bioreactor-aided cell seeding and culturing of scaffolds prior to implantation promoted neovascularization in the scaffold [62
64]. Urothelial and smooth muscle cells are two main cell types in the bladder tissue. Urothelial cells are unique group of cells that form the urothelium layer of the bladder, creating a barrier to the toxic substances in urine. Urine storage and voiding function of bladder is dependent to the relaxation and contraction of bladder smooth muscle cells. Particularly, it has been known that both the urothelial cells and bladder smooth muscle cells under dynamic stimulation were more likely to fully differentiate. In the study of Cattan et al, when human urothelial cells are subjected to shear flow that is provided by a flow bioreactor, they are able to fully differentiate, leading to functional bladder epithelium formation [65]. On the other hand, in vitro bioreactor environment is not fully capable of reproducing the microvasculature of native tissues. More recently, the use of in vivo bioreactors for tissue-engineering applications appears to gain attention because they can provide the most physiologically relevant environment to the biomaterials prior to implantation. This approach emerged from the need of vascularized large-scale tissue substitutes. Briefly, this biomaterial fabrication technique is based on the creation of a vascular bed in the peritoneal cavity, dorsal skin-fold chamber, kidney capsule, or omentum space of the organism [66,67]. This bed acts as a natural bioreactor, enabling vascularization of the biomaterials before implantation. Such a biomaterial with adequate blood supply network that is implanted to the surgical site can boost functional tissue formation, preventing fibrosis. A clinical study, based on showing the safety of an in vivo bioreactor approach, was conducted on patients with breast cancer [68]. In the study, 3D-printed acrylic tissue chambers were implanted in the chest wall of the patients with breast cancer. This was performed to enable vascularization of the acrylic tissue chamber before its implantation. Notably, they had shown that their in vivo chest bed system was capable of creating a well-vascularized large-scale new tissue 6 months after the implantation. Therefore self-vascularized biomaterials grown in the patient’s own body can be a realistic approach for the future tissue-engineering applications. Over the last few decades the challenges faced with TE are still relevant today. To be able to produce customized and functional tissue substitutes, the TE techniques should deviate from conventional tissue-engineering approaches to shift toward more biologically based approaches that combine engineered smart biomaterials with advanced cell therapy.

17.3

Conclusion and future perspectives

Despite the difficulties encountered in urological TE during the last 30 years, the encouraging effort of researchers will eventually find a way to tackle all kinds of challenges, which will help thoroughly TE evolve toward higher levels of success.

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The novel technologies will grow rapidly and will eventually adapt themselves to the regulatory requirements in the foreseeable future. Particularly, 3D-printing technology has a great potential to open new avenues for creating smart biomaterials, which may allow further advances in TE and regenerative medicine. Biomaterials with programmable simple architecture will help people avoid regulatory challenges during clinical transition of new biomaterials. Considering overall TE advances so far, the future of the clinically relevant biomaterials will be determined by innovative research conducted with interdisciplinary groups of scientists, clinicians, and industry.

Conflict of interest We hereby declare that there are no commercial or financial relationships that could be construed as a potential conflict of interest.

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[63] Petrenko YA, Petrenko AY, Martin I, Wendt D. Perfusion bioreactor-based cryopreservation of 3D human mesenchymal stromal cell tissue grafts. Cryobiology 2017;76:150
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45.

Further reading Robb KP, Shridhar A, Flynn LE. Decellularized matrices as cell-instructive scaffolds to guide tissue-specific regeneration. ACS Biomater Sci Eng 2018;4(11):3627
43.

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Respiratory tissue replacement and regeneration: from larynx to bronchi

18

Lea Fath1,2,3, Esteban Brenet 2,3,4, Dana M. Radu5, Emmanuel Martinod5,6 and Christian Debry1 1 Department of Otorhino-Laryngology, ENT Surgery, University Hospital of Strasbourg, Strasbourg, France, 2Inserm UMR 1121, Biomate´riaux et Bioinge´nierie, Strasbourg, France, 3Universite´ de Strasbourg, Strasbourg, France, 4Department of OtorhinoLaryngology, ENT Surgery, University Hospital of Reims, Reims, France, 5Assistance ˆ ˆ ˆ Publique - Hopitaux de Paris, Hopitaux Universitaires Paris Seine-Saint-Denis, Hopital Avicenne, Chirurgie Thoracique et Vasculaire, Universite´ Paris 13, France, 6Laboratory Hypoxia and the Lung INSERM UMR 1272, Universite´ Paris 13, France

18.1

Introduction

The airways extend from the nasal fossae to the ends of the pulmonary alveoli and comprise an aerodigestive junction, the larynx. These airways can be the site of various pathologies requiring the development of reconstruction and replacement techniques of the larynx and the trachea. The reconstruction of the larynx, a trifunctional complex organ ensuring swallowing, phonation, and respiration, requires either a complete laryngeal graft or the creation of an artificial larynx comprising the combination of two structures: (1) restoring laryngeal sphincter function and (2) another prolonging the remaining trachea after total laryngectomy. The research pathways for tracheal reconstruction are therefore indispensable preambles for the reconstruction of the artificial larynx, and in particular for the structure prolonging the trachea. The four main lines of research explored in tracheal and laryngeal reconstructions are laryngeal transplantation, allo- and autografts to compensate for tracheal loss, biomaterials, and tissue engineering.

18.2

Normal respiratory tissue

18.2.1 Embryology During the fourth week of life in utero the laryngotracheal cleft appears on the ventral side of the primitive digestive tract, below the fourth pair of pharyngeal arches. The laryngotracheal cleft is a longitudinal sulcus lined by the endoderm that will Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00012-X © 2020 Elsevier Ltd. All rights reserved.

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further given rise to the respiratory epithelium of the larynx, trachea, and bronchi. The connective tissue comes from the adjacent mesoderm [1]. At the end of the fourth week the groove expands at the craniocaudal junction of the anterior intestine and forms a pocket: the laryngotracheal respiratory diverticulum. The diverticulum separates from the anterior intestine through the formation of bilateral longitudinal ridges. They will merge to form the tracheoesophageal septum, separating the trachea in the ventral position, from the esophagus to the dorsal position. This is the reason for the common vascularization and innervation elements between trachea and esophagus. At the beginning of the fifth week the tracheal bifurcation is in place. The cartilaginous rings develop from the tenth week due to the condensation of the mesoblast around the laryngotracheal diverticulum [2,3].

18.2.2 Larynx The larynx, located at the intersection between the digestive and aerial ways, is responsible for three essential functions: breathing, swallowing, and phonation.

18.2.2.1 Descriptive anatomy A medial organ located in the visceral sheath at the medial and anterior parts of the neck, the larynx (Fig. 18.1), has a triangular pyramid shape whose posterosuperior base inclines forward to the hyoid bone and pharynx, and whose lower vertex meets at the upper end of the trachea [4]. Its dimensions vary according to age and sex: the volume of the larynx is greater in men and lower in the cervical region than in

Figure 18.1 Anatomy of the larynx: (1) hyoid bone, (2) hyothyroid membrane, (3) thyroid cartilage, (4) cricothyroid membrane, (5) cricoid cartilage, and (6) trachea.

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women. Thus, in adulthood, vocal codes measure 14
15 mm in men and 10
15 mm in women [5]. The larynx is made up of 11 cartilaginous structures: G

G

G

thyroid, cricoid, and epiglottic cartilages, which are odd and median: arytenoid cartilage, Santorini cornice, Wrisberg wedge-shaped, and anterior sesamoid, which are even and lateral; and interarytenoid cartilage and posterior sesamoid cartilage, inconstant.

The hyoid bone, U-shaped with posterior concavity, does not belong to the larynx but is attached to it. The larynx is a mobile organ, which elevates during swallowing and the emission of high-pitched sounds, and it is lowered during the production of bass sounds. This mobility is ensured by the cricothyroid joints (allowing the movement of the vocal cords) and the cricothyroid joints (playing a role in the modulation of the voice). There are two kinds of muscles: the extrinsic muscles, which go from the larynx to the neighboring organs and are elevators or depressors of the larynx, and the intrinsic muscles, which are divided into three groups according to their action on the vocal cords: G

G

G

tensor muscles: cricothyroid muscles dilator or abductor muscles: posterior cricoarytenoidal muscles constrictors or adductors: lateral cricoarytenoidal muscles, lower and upper thyroarytenoid, and interarytenoid.

Arterial vascularization is provided by three pedicles: G

G

G

the superior laryngeal artery the cricothyroid artery or middle laryngeal artery the lower laryngeal artery

The upper and lower laryngeal veins drain into the superior thyroid veins. The posterior laryngeal veins flow into the lower thyroid veins. Lymphatic drainage of the larynx can be divided into three territories: G

G

G

the first: supraglottic, the largest and most dense, which is drained in areas IIa and III; the second: subglottic, less dense than supraglottis, which drains to areas VI (prelaryngeal lymph nodes) and areas III and IV; and the third very poor: glottic.

The innervation of the larynx is ensured by the upper and lower laryngeal nerves, branches of the vagus or pneumogastric nerve, tenth pair of cranial nerves.

18.2.2.2 Endoscopic anatomy The endoscopic anatomy of the larynx leads to the division of the larynx into three functional units: G

the supraglottic stage containing the epiglottis, the aryepiglottic fold, the arytenoids, the ventricular bands, and the laryngeal ventricle;

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G

G

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the glottal floor containing the vocal cords and the anterior and posterior commissures; and the subglottic stage that is limited by the vocal cords to the lower part of the cricoid cartilage.

18.2.2.3 Physiology The larynx acts as a real sphincter and provides a role of protection of the airway during the passage of the food bolus and avoids false routes (aspiration) through the closure of the glottis. During swallowing, closure and ascension of the larynx release the cricoid, allowing to direct the food bolus toward the esophageal opening [6]. The epiglottis tilts and protects the glottal plane. The second role of larynx is the generation of sound [7]. The vibration of the mucosa of the vocal cords under the effect of the air passing through the laryngeal column allows the production of sounds. This vibration will vary depending on the tension and therefore the length of the vocal cords which is under the control of two muscles, the cricothyroid, and thyroarytenoid muscles. The third function of the larynx, breathing (or connecting the upper and lower respiratory tracts), is possible thanks to the opening of the vocal cords during the inspiration. During exhalation the vocal cords get closer under the action of adductor muscles of the larynx. The larynx also intervenes during closed glottis efforts (coughing, defecation, childbirth, etc.), to maintain a significant glottal pressure.

18.2.3 Lower airways: trachea, carina, bronchi, bronchioles 18.2.3.1 Anatomy The trachea (Fig. 18.2) is a membranocartilaginous duct in the form of a cylinder flattened at its posterior surface. This posterior surface represents a quarter to 1/5 of the circumference. Its anterior cylindrical portion has superimposed transverse projections corresponding to the tracheal rings. The trachea goes obliquely down and back. Thus in its cervical portion, it is only 15 mm from the skin, while at the sternal fork, it is 3 cm from the skin. In the thorax, due to this oblique orientation and the reverse inclination of the anterior chest wall, the trachea becomes even more posterior and difficult to access anteriorly. In addition, the trachea is deviated slightly to the right by the arch of the aorta which may be responsible for pathological tracheal compression. The average length of the trachea is 12 cm in men and 11 cm in women. The cervical and thoracic segments are almost the same size: 6
7 cm. However, this length varies according to age, sex, and subjects. Moreover, in the same subject, it can also vary according to whether the larynx is at rest or in motion, depending on whether the head is in extension or in flexion. This amplitude of variation is of the order of 3
4 cm. This flexibility allows for surgical resection-anastomosis of trachea that can carry up to six tracheal rings [8]. Tracheal size also depends on sex (greater in men than on women) and age, which explains the different sizes of tracheostomy tubes and intubation tubes.

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Figure 18.2 Tracheal anatomy: (A) cervical trachea: (1) cricoid cartilage and (2) upper limit of the sternal manubrium; (B) thoracic trachea: (3) Aortic impression of Nicaise and Lejars and (4) interbronchial ligament.

The trachea lies in front of the esophagus and has two segments: a cervical and a thoracic. The cervical trachea extends from the inferior border of the cricoid cartilage to the jugular incision of the sternum, opposite the second thoracic vertebra. In cervical hyperextension the trachea becomes more anterior, which may facilitate certain surgical procedures such as tracheotomies. The thoracic trachea begins at the height of the jugular incision and ends with a bifurcation (carina), at the level of the fifth thoracic vertebra, giving rise to the main right and left bronchi. From their origin the bronchi diverge at an angle of 70 degrees to each other. The right bronchus is short (20
25 mm), large (15
16 mm) and vertical at an angle of 25 degrees to the tracheal axis, which explains the frequency of the right bronchial foreign bodies. In contrast, the left bronchus is longer (40
45 mm), smaller in size (10
11 mm) and horizontally at a mean angle to the tracheal axis of 45 degrees. Throughout its height the trachea is surrounded by a loose celluloid tissue, playing the role of serous and promoting tracheal movements. At the thoracic level, this tissue extends throughout the mediastinum, explaining the thoracic spread of cervical infections or emphysema, during tracheal wounds.

18.2.3.2 Histology The trachea (Fig. 18.3) consists of two tunics: an external fibromusculocartilaginous and an internal mucosa. Throughout its height the trachea is surrounded by a celluloadipose tissue, playing the role of serous and promoting its movements.

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Figure 18.3 Microscopic section of trachea.

The outer tunic is a fibroelastic sheath that allows for tracheal dilatation and relaxation during respiratory movements. In the thickness of this connective tissue are the hyaline cartilaginous rings, in the form of “horseshoes,” open behind, preventing the tracheal lumen from collapsing with the inspiration. Posteriorly, transverse smooth muscle fibers unite the ends of the cartilaginous rings, forming the tracheal muscle. The contraction of this muscle results in the reduction of the tracheal size. The tracheal mucosa consists of a ciliated, goblet pseudostratified respiratory epithelium with short apical villi responsible for mucociliary activity and drainage of glandular secretions to the pharynx. This epithelium rests on a thick basement membrane that separates it from the underlying chorion. The chorion is composed of a loose connective tissue, highly vascularized, denser in depth, forming a band of fibroelastic tissue. The deeper submucosa is rich in seromucous mixed glands, the number of which decreases in the distal part of the trachea. The irritation of the tracheal mucosa by tobacco smoke causes a metaplasia of the respiratory epithelium, evolving into a stratified squamous epithelium with disappearance of ciliary activity.

18.2.3.3 Physiology The trachea, by its cervicothoracic situation and its fibroelastic structure, allows the passage of air to the pulmonary alveoli. But the trachea is not just an air duct. Through its mucociliated epithelium, it allows the discharge of secretions upstream to the larynx, spontaneously and during coughing efforts. In addition, the presence of lymphoid clusters in its wall gives it an immune defense function.

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18.3

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Airways diseases

18.3.1 Laryngeal diseases Any decrease in the caliber of the larynx causes breathing difficulties with inspiration: which is called laryngeal dyspnea. A stenosis corresponds to a decrease in the size of the laryngeal or tracheal lumen compared to the normal caliber. The causes of laryngeal stenosis are diverse but are most often acquired in adults and generally congenital in children. Laryngeal stenosis could be G

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congenital: lack of resorption of the epithelial plate of the primitive bowel and acquired supraglottic [9]: ingestion of caustic, gastroesophageal reflux, inhalation burns, iatrogenic (laser surgery) and glotto-subglottic [10,11] (Fig. 18.4): traumatic or prolonged intubation, damage due to unsuitable intubation probe, (due to mucosal damage that is the starting point because of pressure point ischemia, resulting in mucosal necrosis and subjacent perichondrium exposure and stenosis), sequelae of benign or malignant laryngeal tumors, and also surgical treatment of these tumors by partial laryngectomies, radiotherapy, and external laryngeal trauma. G

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Laryngeal dyspnea can cause respiratory distress, which can be fatal in the absence of appropriate management. In addition, depending on the etiology of dyspnea, the mobility of the larynx can be impaired, thus preventing it from acting as a sphincter protecting the lower respiratory tract. Moreover, in this area, patients can also suffer from swallowing disorders that can cause the passage of the food bolus directly in the airways which will cause inhalation pneumopathies. If not treated correctly, these inhalation pneumopathies can be fatal. Thus central degenerative neurological pathologies (stroke, amyotrophic lateral sclerosis, multiple sclerosis, and Parkinson’s syndrome) or peripheral neuropathy (Guillain-Barre´ syndrome and

Figure 18.4 Glottic stenosis.

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myopathies) can cause serious disorders of swallowing, resulting in pulmonary pneumonitis. Laryngeal cancers also impair glottic mobility and reduce the diameter of the laryngeal spinneret. It is therefore quickly necessary to perform a tracheotomy (establishment of another airway by introducing an incision in the trachea secured with a cannula) to bypass the obstacle and allow normal breathing and protection of the lungs.

18.3.2 Tracheobronchial diseases 18.3.2.1 Tracheal stenosis Tracheal stenosis may be congenital or acquired [12].

18.3.2.1.1 Congenital stenosis Congenital stenosis represent less than 1% of cases; they can be isolated or associated with cardiovascular abnormalities. There are segmental stenoses (mainly due to complete cartilaginous rings and not in the physiological horseshoe shape) or generalized tracheobronchial hypoplasias (disorganization of cartilaginous rings, fibrous dystrophy, etc.) [13].

18.3.2.1.2 Acquired stenosis G

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Non tumor lesions The main causes are the sequelae of intubation [14] or tracheostomy, posttraumatic, or idiopathic lesions due to inflammatory diseases such as Wegener’s disease (causing circumferential damage to the mucosa with respect for the integrity of costal cartilages) or atrophic polychondritis. Tumor lesions Intrinsic Malignant tumors, of which the most common histological types are cystic adenoid carcinomas, squamous cell carcinomas, lymphomas, sarcomas, are at the origin of narrowing of the tracheal size. Benign tumoral etiologies such as papillomatosis or inflammatory granulomatous type can also result in the development of stenosis. Extrinsic G

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Extrinsic compression is the origin of tracheal stenosis with a healthy mucosa but may eventually damage the strength of the cartilaginous rings of the trachea and give tracheomalacia, resulting in the collapse of the trachea during respiratory movements.

18.3.2.2 Oeso-tracheal fistulas These fistulas are related to a lack of estracheal septum formation that normally separates the primary and tracheal buds between the fourth to eighth week of embryogenesis. They are most often associated with atresia of the esophagus but can also be isolated [15]. These fistulas are associated with other malformations in 93% of cases. The most common syndrome found is VACTERL syndrome associating vertebral, anorectal, cardiac, tracheal, esophageal, renal, and limb malformations. They are also found in Goldenhar syndrome or CHARGE syndrome.

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Replacement and regeneration strategies

18.4.1 Laryngeal transplantation The risk of organ transplantation and immunosuppression imposed by posttransplant treatment is justified in the terminal phases of a disease whose outcome would be fatal because of the vital nature of the organ in question. Laryngeal transplantation has long been considered unjustified because of possible survival after total laryngectomy. Nevertheless, the first laryngeal transplantation in humans was performed in a 40-year-old patient who required total laryngectomy after laryngeal trauma, in 1998 by Strome. The graft made it possible to obtain a good quality voice [16]. However, tracheotomy could never be weaned due to incomplete reinnervation of the graft. This patient presented several episodes of rejection in the years following implantation, leading to a progressive deterioration of the graft and finally to its explantation 14 years after its implantation [17]. The problem of immunosuppression, essential in the case of laryngeal transplantation, is a major problem that is difficult to accept ethically in a carcinological context because of the risk of recurrence or secondary neoplastic localization. However, progress in this area offers hope for more and more selective immunosuppressive drugs that may be prescribed in oncological settings someday. The number of compatible donors is also a limiting factor that is difficult to overcome. This intervention of great difficulty also requires a trained surgical team. Laryngeal transplantation therefore appears to be a remarkable surgical feat which is difficult to achieve at the present time, addressing a limited number of patients and currently not suitable for oncological management.

18.4.2 Biomaterials Since the 2000s, Debry et al. [18] have begun designing, developing, and testing an artificial larynx to try to remove the tracheostomy port, a problem that has never been solved since the first laryngectomy in 1873. This laryngeal prosthesis combines an immovable rigid structure of porous titanium extending the trachea and a removable mechanically functional structure recreating the laryngeal sphincter [19] (Fig. 18.5). Several studies have been carried out to optimize the tissue integration of the prosthesis: a first work made it possible to highlight that the addition of an endoprosthetic silicone tube improved survival in the large animals [20,21]. A biodegradable polymeric material, poly(L-lactic acid), was then positioned on the endoluminal side of the prosthesis allowing better tissue integration than bare porous titanium prostheses and control over the relative positioning of the epithelial layer and the incoming connective tissue component [22]. The next study then focused on the variation in titanium bead size showing that the colonization rate of prostheses was accelerated when the size of the beads was decreased [23,24].

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Figure 18.5 Videofluoroscopy in a patient implanted with an artificial larynx: passage of the food bolus in the esophagus without false route.

These studies led to the first successful worldwide implantation of artificial larynx in humans in 2012 [18,25]. Long-term fatigue-free breathing with closed tracheostomy, both diurnal and nocturnal, in the last implanted patients was proven, with an understandable whispering speech, but persistently with swallowing with false residual routes (secondary aspiration), but no patient presented pneumonitis during the implantation period. Other limitations to this artificial larynx were the use of titanium as a biomaterial for the laryngeal prosthesis causes a compression of the mouth of the esophagus during swallowing, as well as tissue integration defects were highlighted, with the need to coat the prosthesis with a pedicled muscle flap of pectoralis major. An extrusion of the immovable part of the artificial larynx prosthesis into titanium is also a long-term risk, especially after radiotherapy, even if the technique of flap coverage has prevented such occurrences for periods up to at 16 months (death by uncomplicated secondary location in relation to the implant). However, it appeared that the improvement potential of this technique remained considerable (1) by using a more flexible biomaterial for the body of the prosthesis, (2) by allowing a complete cellular integration, and (3) by developing the concept not of artificial larynx but of artificial pharyngo-larynx, in order to minimize the false routes by optimizing the design of the pieces by gutters mimicking the functionality of the piriform sinuses.

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18.4.3 Tissue engineering The goal of tissue engineering is to recreate a functional organ similar to that intended to be replaced. The analysis of the results obtained by tracheal engineering also makes it possible to progress in laryngeal rehabilitation techniques. In fact, outside the vocal cords and the laryngeal muscles, the trachea and the larynx have a similar histological structure. Tracheal reconstructions thus constitute a preamble to laryngeal reconstructions, which are more complex because they also require restoration of the laryngeal sphincter. Tissue engineering combines a matrix that provides the shape and support necessary for the architecture of reconstructed tissue and cells, which colonize the matrix when placed in a favorable environment. The constituted tissues can be developed in vitro before being implanted in vivo. In tracheal engineering, it is necessary to associate the restoration of (1) rigid cartilaginous tracheal rings connected by a flexible connective tissue plate and (2) a ciliated respiratory epithelium whose primary role is to protect the airways from the external environment by mucociliary purification of secretions and microorganisms. Bronchial reconstruction methods follow the same specifications as tracheal engineering but require to apply these concepts to small dimensions, which corresponds to an additional challenge.

18.4.3.1 Scaffold The ideal scaffold in laryngeal tissue engineering must mimic the properties of the extracellular matrix of the biological larynx and must have different properties: (1) rigidity to maintain the airway permeability without collapse during inspiration to ensure a supportive role to surrounding tissues; (2) porosity to promote the most complete integration possible to the surrounding tissues of the seeded cells; (3) three-dimensional architecture [26], to allow adhesion and cell proliferation, then differentiation into chondroblasts; (4) promote neoangiogenesis; (5) to resist the aggressive medium existing in the airways, namely, a very variable pH, the presence of multiple bacteria [27], molds, and moisture; (6) present an absence of toxicity and to be biocompatible; and (7) ensure the absence of immunogenicity within the implanted organism.

18.4.3.1.1 Manufactured scaffolds Manufactured scaffolds were made of natural biodegradable compounds such as collagen (or denaturated collagen, i.e., gelatin [28]), alginate or hyaluronic acid or synthetic polymers such as polyglycolic acid, polylactic acid, polypropylene fumarate, and metals. These systems can also be supplemented with growth factors [29,30]. These scaffolds have produced good outcomes in tracheal tissue engineering both in vitro [31] and in vivo: promising outcomes have been obtained in rats [32] and rabbits [33] but were less conclusive in sheep [20] and dogs [34]. Their main disadvantage in vivo is defective integration with neighboring tissues, causing extrusion of the implanted scaffolds, migration, surgical site infection, and excessive inflammatory reactions due to intraprosthetic or anastomotic stenosis.

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Synthetic 3D printed scaffolds based on CT scan reconstruction of the trachea in pigs have been produced [35,36] using polycaprolactone-based biomaterials. This reconstruction involved anterior tracheal defects or the whole trachea circumferentially several centimeters high (as many as 4 tracheal rings). Postoperative followup was provided for up to a maximum of 3 months, and endoscopic follow-up in the final analysis revealed substantial formation of granulation tissue that obstructed the prosthetic device.

18.4.3.1.2 Biological matrix in tracheal engineering 1. The decellularized tracheal matrix The tracheal extracellular matrix is also widely used in tracheal tissue engineering. A process of decellularization is carried out in order to eliminate all the cells present within the tissue while preserving the extracellular matrix. The goal is to make it less immunogenic and thus avoid immunosuppressive therapy. The methods of decellularization are multiple but the most frequently used ones are chemical and enzymatic [37]. The tracheal matrix is seeded with the cells of interest after decellularization: stem cells, chondrocytes, and epithelial cells. Good results were obtained in vitro after the seeding of chondrocytes on tracheal matrices previously treated with CO2 laser to make the matrix more porous and then decellularized by 18 cycles of enzymatic debridement. The complexity and the length of the treatments necessary for this matrix decellularization seem not very transferable to the current clinical practice, as much at the tracheal level as at the laryngeal one. Controversial results in tracheal reconstruction in humans using this type of matrix have put this aspect of research currently at standby [38]. 2. The aortic matrix

The use of an aortic matrix in tracheal engineering in vivo has already been proven, as well as in digestive surgery [39]. Widely used in vascular engineering [40], it has already been used in tracheal engineering, never as a single matrix but always associated with biomaterials such as polycaprolactone [31].

18.4.3.2 Cells 18.4.3.2.1 Cartilage The tracheal cartilaginous rings and the thyroid and cricoid cartilages are composed of hyaline structures. This hyaline cartilage contains few cells (10% of the mass) of mesenchymal origin called chondroblasts (young cell) and chondrocytes (mature cells) located in stalls, chondroplasts. These cells will secrete the extracellular matrix that is composed of 40% of substance containing glycosaminoglycans and aggrecans binding hyaluronic acid and thus forming a negatively charged macromolecular complex attracting molecules of H2O, and 50% of collagen fibers (mainly type II collagen). 1. Cartilage cells Chondrocytes can be removed in vivo from the nasal septum, conch, and costal cartilages [41]. Their use is limited because of the invasive and painful removal, exposing the patient to the risk of infection, and the small number of cells available. In addition, their

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cell culture is difficult to control [42] because of the very rapid dedifferentiation of chondrocytes into fibroblasts if they are grown on a two-dimensional matrix [43] and cell multiplication is much slower than stem cells. 2. Mesenchymal stromal cells

The mesenchymal stromal cells of fetal origin [44,45], which are very abundant in the umbilical cord, are undifferentiated mesenchymal stem cells possessing a high capacity for self-renewal. They have the major advantages of inducing low immunogenicity, secreting antiinflammatory factors and being provided with a high capacity for differentiation into chondroblasts and chondrocytes. The use of mesenchymal cells from Wharton’s jelly in cartilage tissue engineering has been validated by multiple studies [46
48]. The chondrogenic pathway can thus be obtained in the presence of dexamethasone, TGF-β1 (transforming growth factor-β1), L-proline, and ascorbic acid. This pathway is favored by hypoxia and requires a three-dimensional matrix [49,50].

18.4.3.2.2 Respiratory epithelium The proximal airway surface epithelium provides protection to the respiratory mucosa through a variety of mechanisms, including mucociliary clearance, regulation of ion and water fluxes, and secretion of defense molecules. The second line of protection is provided by intercellular junctional complexes to preserve the barrier function of the epithelium. The surface respiratory epithelium is pseudostratified: all cells are anchored to the basal lamina, but only some of them extend to the bronchial lumen. This epithelium consists of basal, ciliated, muco-secretory, neuroendocrine, and intermediate cells, with a network of submucosal glands containing mucous and serous secretory cells. The vibratory cilia of the hair cells have an essential role of purification by mucociliary clearance. In the same mode as cartilage tissue, tissue engineering may allow in vitro reconstitution of respiratory epithelium to the endoluminal face of the trachea. This delicate step is divided into two parts: (1) obtaining a confluent epithelium and (2) the differentiation of the latter into ciliary muco-secreting respiratory epithelium [51]. Several teams obtained differentiated epithelium from human epithelial primary tracheal cells [52]. Once seeded on the surface of the matrix, the goal is to first obtain a confluent epithelium with strong intercellular junctions. The second differentiation step requires culture under air/liquid conditions [53]. This is based on a culture system in two compartments separated by a porous membrane on which the seeded matrix rests. Once the confluence is obtained, the liquid medium is removed from the upper compartment and the surface of the epithelium is exposed to the air. After a maturation period of 2 weeks the aerial exposure allows the differentiation into ciliary cells presenting a motility, also the apico-basal differentiation. The difficulty then lies in creating a gradient of growth factors, allowing the cilia to beat in a coordinated and unidirectional way as in in vivo conditions [51].

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18.4.3.2.3 Cocultures The use of a bioreactor with double seeding chamber allows the simultaneous inoculation mesenchymal stem cells and tracheal epithelial primary cells, proliferation and differentiation of mesenchymal stem cells to chondrocytes, and the formation of the respiratory epithelium at the same time on the endoluminal face [54]. Studies have examined the use of decellularized trachea as a matrix for this double seeding bioreactor system with in vivo rodent tests giving interesting results, but with a survival of animals limited to two week, or 1 year in a heterotopic position (paravertebral muscles) [55].

18.5

Transplant

In a similar way to tissue engineering, tissue grafts used in tracheal rehabilitation are a preamble to laryngeal rehabilitation work. The ideal laryngeal substitute after total laryngectomy is composed of two parts: one recreating the laryngeal sphincter and the other extending the trachea. The procedures developed in tracheal reconstruction thus make it possible to extrapolate these results in order to answer the problem of the “tube” prolonging the trachea and supporting the laryngeal sphincter.

18.5.1 Nonliving tissue transplants Bioprostheses consisting of freeze-dried [56], frozen or chemically fixed tissues, glutaraldehyde-treated tracheal allografts, freeze-dried aortic grafts, and formalinfixed tracheae were implanted in animals and then in humans, but they have been up to now unsuccessful because of the absence of vascularization of the tissues which eventually result in necrosis [57].

18.5.2 Autografts Many tissues have been used as autografts to compensate for tracheal defects, such as fascia [58], periosteum, aorta, dermis [59], pericardium [60], the association of atrial cartilage with oral mucosa [61], and jejunum [62]. These grafts may be surrounded by flaps (omentum and pericardium) to provide vascular supply and to make the construction impermeable. Encouraging results have been obtained, but these highly invasive techniques are not very applicable in humans. Tracheal reconstruction was performed by Delaere and Hardillo [60] from a latero-thoracic fascia reinforced by atrial cartilage in the rabbit. The set was tubulated and transferred with his pedicle to replace 2 cm of tracheal defect. An absence of endoluminal epithelial colonization associated with partial necrosis and granulation buds obstructing the neo-trachea which required rapid animal sacrifice. Olias et al. [63] successfully performed a tracheal reconstruction of 8 cm long using an antebrachial flap in which had been implanted upstream bands of costal

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cartilages previously sutured in a circle. Ten weeks later, he grafted oral mucosa to the surface of the vascularized cartilage by the antebrachial flap. During a last operative period the free antebrachial flap enriched with costal cartilage and mucosa was removed, tabulated, and anastomosed instead of the trachea, ensuring the permeability of the respiratory tract. This technique is unfortunately not usable in the cases of restoration after cancer resection, because of the too long delay between the first and last interventions. In addition, this technique is invasive and mutilating for the donor areas of the grafts. Recently, Kolb et al. [64] published the restoration of almost all of a trachea (from the first tracheal ring to the carina) in a 12-year-old boy with laryngeal stenosis, using a myoduncal flap of latissimus dorsi costal cartilages spaced 2 cm apart, tubular flap held by a Y silicone stent. The restoration of the trachea was performed in a single operation. A postoperative follow-up at 4 years shows no major complications and the tracheotomy could be weaned definitively 2 years after the intervention.

18.5.3 Allografts Because of the invasive nature of the autograft, other teams have focused on the use of allografts stented by a prosthesis, often silicone, to overcome tracheal deficiencies. Unprocessed fresh tissue before implantation is responsible for immunogenicity, requiring long-term immunosuppressive treatment to prevent rejection of the graft, which is not possible in a carcinological context of laryngeal rehabilitation. Allografts must undergo treatment before implantation. The goal is to decrease their immunogenecity, while maintaining the structure of the extracellular matrix. Many complex protocols have been proposed, hardly feasible in current clinical practice. Aortic allografts were removed to be implanted to replace 7 cm of circular tracheal defect in ewes by Martinod et al. [65]. Three preimplantation graft treatment techniques were evaluated: (1) cryotherapy at 280 C, (2) glutaraldehyde fixation, and (3) decellularization protocol (performed with sodium dodecyl sulfate, hypertonic rinse followed by 20 C). Cryotherapy seemed to be the most promising technique: the histological examination after cryotherapy showed a preservation of normal aortic architecture. At 1 month an important inflammatory reaction is found, at 3 months a nonkeratinized epithelium appears, at 6 months a continuous mucociliary epithelium is observed, and at 12 months in the graft. Thus even after extensive inflammation of the tissue in the first month, the vessel is recolonized by a well-differentiated ciliated epithelium. Additional studies conducted by Martinod et al. on the ewe [65
67], then on the man [68,69] confirmed these results. Cryotherapy seems to be an effective solution to decrease immunogenicity while preserving the tissues of the grafts to implant. No immunosuppression is necessary because of the disappearance of the vascular tissue. This has been demonstrated by the study of the SRY gene on cartilage neocells: [69] no SRY gene of male origin was found in the newly formed cartilage, meaning that the newly formed cells are indeed cells from the female host organism and not the male donor.

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Recently, a prospective publication [70] of 20 patients, 13 of whom had undergone a tracheal, bronchial or carinal reconstruction by aortic allograft, from 2009 to 2017, reported a 95% postoperative survival rate (90 days) without direct adverse effect of the surgical technique. The endoluminal stent was removed on average 18.2 months after implantation, with a median follow-up at 3 years and 11 months and a survival rate of 76.9% (10 out of 13 patients). These amazing results, still imperfectly explained, certainly open a new route of tracheobronchial reconstruction, and by extension, laryngeal replacement. By not requiring any immunosuppressive treatment, this technique could therefore be applied to all patients with laryngectomy for malignancy (the vast majority of cases).

18.6

Conclusion and outlook

The development of an ideal laryngo-tracheobronchial substitute has not yet been achieved. The combination of different research approaches can hope to create a hybrid substitute combining tissue engineering, allograft, and biomaterials. The aortic allograft technique, which does not require immunosuppression and has been proven in tracheal reconstruction, is a potential candidated to be tested for the body of artificial laryngeal prosthesis. If the results obtained at the tracheal level can be extrapolated to the larynx, it would then be possible to use this “neo-laryngeal tube” to serve as a support for the laryngeal sphincter while allowing integration into the surrounding tissues of the recipient organism. These processes of chondrogenesis and reepithelialization of aortic allograft in vivo remain long in clinical practice and need to be accelerated. In vitro optimization of grafts by treatment and seeding of the aortic allograft could help to address this problem. The main problem of laryngeal reconstruction is still the realization of an effective sphincter allowing the weaning of the tracheotomy without risk of false routes during swallowing. Innovative 3D printing techniques could provide a flexible biomaterial recreating this laryngeal sphincter while restoring the flexibility of the initial tissues, which is necessary to maintain a good quality of life of the rehabilitated patients.

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9. [43] Francioli SE, Candrian C, Martin K, Heberer M, Martin I, Barbero A. Effect of threedimensional expansion and cell seeding density on the cartilage-forming capacity of human articular chondrocytes in type II collagen sponges. J Biomed Mater Res A 2010;95(3):924
31. [44] Lim J, Razi ZRM, Law J, Nawi AM, Idrus RBH, Ng MH. MSCs can be differentially isolated from maternal, middle and fetal segments of the human umbilical cord. Cytotherapy 2016;18(12):1493
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7. [46] Kim D-W, Staples M, Shinozuka K, Pantcheva P, Kang S-D, Borlongan CV. Wharton’s jelly-derived mesenchymal stem cells: phenotypic characterization and optimizing their therapeutic potential for clinical applications. Int J Mol Sci 2013;14 (6):11692
712. [47] Aleksander-Konert E, Paduszy´nski P, Zajdel A, Dzier˙zewicz Z, Wilczok A. In vitro chondrogenesis of Wharton’s jelly mesenchymal stem cells in hyaluronic acid-based hydrogels. Cell Mol Biol Lett 2016;21:11. [48] Reppel L, Schiavi J, Charif N, Leger L, Yu H, Pinzano A, et al. Chondrogenic induction of mesenchymal stromal/stem cells from Wharton’s jelly embedded in alginate hydrogel and without added growth factor: an alternative stem cell source for cartilage tissue engineering. Stem Cell Res Ther 2015;6:260. [49] Bhattacharjee M, Coburn J, Centola M, Murab S, Barbero A, Kaplan DL, et al. Tissue engineering strategies to study cartilage development, degeneration and regeneration. Adv Drug Deliv Rev 2015;84(Suppl. C):107
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40. Gebauer DW. Plastic reconstruction of tuberculous bronchostenosis with dermal grafts. J Thorac Surg 1950;19:604 20628. Delaere PR, Hardillo J. Tubes of vascularized cartilage used for replacement of rabbit cervical trachea. Ann Otol Rhinol Laryngol 2003;112(9 Pt 1):807
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22.

Platelet-rich plasma in tissue engineering

19

Anne Lehn1,2 1 Department of Pediatric Surgery, University Hospital of Strasbourg, Strasbourg, France, 2 UMR DIATHEC, EA 7294, Translational Medicine Federation of Strasbourg (FMTS), University of Strasbourg, Strasbourg, France

19.1

Introduction

Platelet products have been extensively used to improve wound healing in clinical and surgical applications. Many platelet derivatives can be obtained using different procedures of platelet preparation and leading to variable effects on growth factor (GF) release and cell growth.

19.1.1 Blood composition 19.1.1.1 Plasma Plasma is the liquid component of blood, in which the red blood cells, white blood cells, and platelets are suspended. It constitutes more than half of the blood’s volume and consists mostly of water that contains dissolved salts (electrolytes) and proteins (Fig. 19.1). The major protein in plasma is albumin. Albumin helps keep fluid from leaking out of blood vessels and into tissues, and albumin binds to and carries substances such as hormones and certain drugs. Other proteins in plasma include antibodies (immunoglobulins), which actively defend the body against viruses, bacteria, fungi, and cancer cells, and clotting factors, which control bleeding. Plasma has various functions. It acts as a reservoir that can either replenish insufficient water or absorb excess water from tissues. When body tissues need additional liquid, water from plasma is the first resource to meet that need. Plasma also prevents blood vessels from collapsing and clogging and helps maintain blood pressure and circulation throughout the body simply by filling blood vessels and flowing through them continuously. Plasma circulation also plays a role in regulating body temperature by carrying heat generated in core body tissues through areas that lose heat more readily, such as the arms, legs, and head.

19.1.1.2 Red blood cells or erythrocytes Red cells are the most abundant cells in the blood, accounting for about 40%
45% of its volume. The shape of a red blood cell is a biconcave disk with a flattened center. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00027-1 © 2020 Elsevier Ltd. All rights reserved.

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Figure 19.1 Schematic picture of a fractionated whole-blood sample. Source: https://microbeonline.com/buffy-coat-definition-composition-preparation-uses/.

Production of red blood cells is controlled by erythropoietin, a hormone produced primarily by the kidneys. Red blood cells start as immature cells in the bone marrow and after approximately 7 days of maturation are released into the bloodstream. Unlike many other cells, they have no nucleus and can easily change shape, helping them fit through the various blood vessels. However, while the lack of a nucleus makes a red blood cell more flexible, it also limits the life of the cell as it travels through the smallest blood vessels, damaging the cell’s membranes and depleting its energy supplies. The red blood cell survives on average only 120 days. Red cells contain a special protein called hemoglobin that helps carry oxygen from the lungs to the rest of the body and then returns carbon dioxide from the body to the lungs so that it can be exhaled. Blood appears red because of the large number of red blood cells, which get their color from the hemoglobin. The percentage of whole-blood volume that is made up of red blood cells is called the hematocrit and is a common measure of red blood cell levels.

19.1.1.3 White blood cells or leukocytes White blood cells protect the body from infection. They are much fewer in number than red blood cells, accounting for about 1% of the blood. The most abundant type of white blood cell is the neutrophil, which is the “immediate response” cell and accounts for 55%
70% of the total white blood cell

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count. Each neutrophil is alive for less than a day, so the bone marrow must constantly make new neutrophils to maintain protection against infection. The other major type of white blood cell is a lymphocyte. There are two main populations of these cells. T lymphocytes help regulate the function of other immune cells and directly attack various infected cells and tumors. B lymphocytes make antibodies, which are proteins that specifically target bacteria, viruses, and other foreign materials. The other types of white blood cells include monocytes, eosinophils, and basophils.

19.1.1.4 Platelets Platelets are small discoid anuclear cell fragments that are key players in hemostasis (a frontline physiological response to acute tissue injury). Under basal conditions, platelets circulate at 150
400 3 109/L and sequester a diverse array of bioactive molecules within their intracellular granules. They are smaller than red or white blood cells. Platelets are fewer in number than red blood cells, with a ratio of about 1 platelet to every 20 red blood cells. Megakaryocytes give rise to circulating platelets through commitment of the multipotent stem cell located in the bone marrow to the megakaryocyte lineage, proliferation of the progenitors, and terminal differentiation of megakaryocytes [1]. During megakaryocyte maturation, internal membrane systems, granules, and organelles are assembled in bulk. Platelet production by cytoplasmic fragmentation requires highly structured intricate changes in the megakaryocyte cytoskeleton and concomitant assembly of anucleate platelets (Fig. 19.2). Platelets contain various elements, including the α-granules that can release cytokines and GFs. α-Granules, with mitochondria and dense bodies, collectively constitute the organelle zone (Fig. 19.3).

19.1.2 How does platelet-rich plasma work? PRP is an autologous blood-derived concentrate with a high number of platelets contained in a small volume of plasma.

19.1.2.1 History The concept and description of PRP started in the field of hematology [2]. The term PRP was created by hematologists in the 1970s in order to describe the plasma with a platelet count above that of peripheral blood, which was initially used as a transfusion product to treat patients with thrombocytopenia. In the 1980s PRP started to be used in maxillofacial surgery as platelet-rich fibrin. Fibrin had adherence and homeostatic properties, and PRP with its antiinflammatory characteristics stimulated cell proliferation. In the early 2000s the use of PRP extended into orthopedics to boost healing in bone grafts and fractures and later in sports medicine for connective tissue repair.

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Figure 19.2 Platelet formation from myeloid stem cells in the bone marrow. The entire megakaryocyte cytoplasm is fragmented into anucleate platelets. Microtubules Dense tubules Surface-connecting tubule Coat Glycogen Mitochondria Alpha granule Dense granule

Figure 19.3 Ultrastructural features observed in thin sections of discoid platelets cut in the equatorial plane. Components of the peripheral zone include the exterior coat, trilaminar unit membrane, and submembrane area containing specialized filaments that form the wall of the platelet and line channels of the surface-connected canalicular system. The matrix of the platelet interior contains actin microfilaments, structural filaments, the circumferential band of microtubules, and glycogen. Formed elements embedded in this matrix include mitochondria, α-granules, and dense bodies. The membrane systems include the surfaceconnected canalicular system and the dense tubular system, which serve as the platelet sarcoplasmic reticulum. Source: From wikimedia commons. https://commons.wikimedia.org/wiki/File:Platelet_ structure.png.

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19.1.2.2 Platelet action Platelets play a key role in hemostasis and clot formation, but they also contain multiple proteins, cytokines, and GFs involved in wound healing and tissue repair. The two principal components of hemostasis are (1) platelets, specialized cells that adhere to the damaged tissue and form a primary plug reducing blood loss (Fig. 19.4A), and (2) blood coagulation, a complex reaction network that turns fluid plasma into a solid fibrin gel to completely seal the wound (Fig. 19.4B). Platelet aggregation and blood coagulation are extremely complex processes. The attachment of platelets and their accumulation into a thrombus is regulated by mechanical interactions with erythrocytes and the vessel wall, by numerous chemical agents such as thrombin, or adenosine diphosphoric acid , or prostaglandins, or collagen, as well as by an enormous network of intracellular signaling. Blood coagulation includes some 50 proteins that interact with each other and with blood or vascular cells in approximately 200 reactions in the presence of flow and diffusion.

Figure 19.4 Two components of hemostasis: (A) electron microphotograph of blood platelets and (B) main reactions of blood coagulation, a reaction cascade that is initiated by tissue factor exposure at the site of damage and produces fibrin, which polymerizes to create a gelatinous clot. Source: From wikimedia commons. https://commons.wikimedia.org/wiki/File:Classical_blood_ coagulation_pathway.png.

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PRP works using the degranulation of the α-granules in platelets, which contain the GFs (Table 19.1). The active secretion of these GFs begins once the platelets are activated. Platelet activation occurs when a blood vessel wall is damaged. It involves different pathways: those involved in adherence and activation in contact with the damaged vascular wall, and those leading to a signal amplification and recruitment of other platelets. They then recruit other platelets to form the hemostatic clot composed not only mainly of platelets but also of fibrin produced by activated platelets and locally generated thrombin [5]. More than 95% of the presynthesized GFs are secreted within 1 hour after the clotting process begins [6]. The optimal concentration of platelets remains a highly debated subject. Haynesworth et al. showed that the proliferation of adult mesenchymal stem cells (MSCs) and their differentiation were directly related to the platelet concentration. They showed a dose-response curve, which proved that a sufficient cellular response to platelet concentrations first began when a four- to fivefold increase over baseline platelet numbers was achieved [7].

Table 19.1 Growth factors present in platelet-rich plasma [3,4]. Name

Acronym

Function

Platelet-derived growth factor a-b

PDGF a-b

Transforming growth factor-β1

TGF-β1

Basic fibroblastic growth factor Vascular endothelial growth factor Epidermal growth factor

bFGF

Stimulates fibroblast production, chemotaxis and mitosis of fibroblast, smooth muscle cells, stimulates TGF-β1, collagen production, and upregulation of proteoglycan synthesis Regulates the mitogenic effect of other growth factors; stimulates the proliferation of undifferentiated mesenchymal cells, fibroblast and osteoblast mitogen, endothelial regulator and regulator of the collagen synthesis and secretion of collagenase; stimulates angiogenesis and endothelial chemotaxis; and inhibits the proliferation of macrophages and lymphocytes Induces the production of collagen, stimulates angiogenesis and proliferation of myoblasts Increases angiogenesis and vessel permeability, stimulates mitogenesis for endothelial cells

Connective tissue growth factor Insulin-like growth factor-1

CTGF

VEGF

EGF

IGF-1

Stimulates endothelial chemotaxis and angiogenesis, regulates collagenase secretion, and stimulates epithelial and mesenchymal mitogenesis Promotes angiogenesis, cartilage regeneration, fibrosis, and platelet adhesion Chemotaxis for fibroblasts and stimulates protein synthesis. Enhances bone formation by proliferation and differentiation of osteoblasts

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19.1.3 Preparation of platelet-rich plasma
based biomaterials Reporting of PRP preparation in the literature is highly inconsistent and it is therefore very difficult to compare PRP products and their efficiency. A metaanalysis, including 105 studies about the use of PRP in musculoskeletal/orthopedic conditions, showed that only 11 (10%) included comprehensive and reproducible reporting of the PRP preparation [8]. Only 17 studies (16%) provided quantitative metrics on the composition of the final PRP product. In this study, 24 different PRP processing systems were used, highlighting the tremendous variety of protocols to obtain PRP products.

19.1.3.1 Use of anticoagulant An anticoagulant is always needed to prevent platelet activation and degranulation before PRP is obtained. Acid citrate dextrose solution A is the most frequently used but some studies also report the use of sodium citrate, citric acid, and citrate
phosphate
dextrose or calcium citrate. A study on rabbits showed that the type of anticoagulant affected the cell count in PRP but not the concentration of GFs [9]. PRP can be manufactured in two basic formats: plasma-based and buffy-coat preparations. Both begin with whole blood but differ in the centrifugation process, which isolates and concentrates different blood-cell components [10]. Plasma-based methods isolate only plasma and platelet components and remove white blood cells. With these protocols, some platelets are left behind but the focus is to intentionally exclude leukocytes, which are thought to be detrimental to the healing process [11
13]. The main goal of this method is to capture only platelets during the centrifugation, to obtain “pure” PRP. Buffy coat
based methods isolate a platelet-poor plasma (PPP) layer and a buffy-coat layer containing both leukocytes and erythrocytes. These protocols seek to capture all available platelets during the centrifugation. Leukocytes and red blood cells are harvested in order to obtain the highest platelet concentration levels possible.

19.1.3.2 Centrifugation Centrifugation processes vary widely among studies, also depending on the processing machine utilized. Two successive centrifugations are usually performed with a first soft spin to obtain two different phases: the first clouded phase containing platelets, PPP, and a buffy coat (Fig. 19.5) and the second phase containing red blood cells. The second spin (hard spin) is performed after collecting the first clouded phase to further concentrate PRP. Both the spins should be moderated because high-spin centrifugation could be deleterious for platelet function by inducing platelet aggregation and activation. Bausset et al. showed that the platelet concentration factor was significantly higher with a 250 3 g speed centrifugation for the second spin compared to 130 3 g. However, a centrifugation performed at 400 or 1000 3 g did not further

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Figure 19.5 Schematic representation of PRP isolation and platelet lysate preparation. PRP is obtained after differential centrifugation of whole blood. Platelets can be activated by thrombin, CaCl2, collagen, or temperature cycles. They then release molecules contained in their granules-like cytokines or GFs. GFs, Growth factors; PRP, platelet-rich plasma.

increase the platelet concentration factor and is not appropriate for PRP preparation [14]. Likewise, in the metaanalysis concerning PRP use in orthopedics, none of the 105 studies utilized a centrifugation speed higher than 4500 rpm [8].

19.1.3.3 Buffy-coat method The debate remains in whether or not to pick the buffy coat for the second centrifugation. It contains a high concentration of platelets; therefore it is useful to preserve it in order to further concentrate PRP. But the risk when collecting the buffy coat entirely is to include red blood cells from the pellet in the PRP product. Their presence could be deleterious in clinical use, like in an intraarticular injection of the PRP product for instance, leading to hemarthrosis associated with cartilage damage and progressive degenerative arthritis.

19.1.3.4 Activation method to induce platelet degranulation and release of growth factors There are different possible ways to activate PRP that entails inducing the release of bioactive molecules from platelets and the cleavage of fibrinogen. Most commonly, thrombin and/or calcium chloride are used for this purpose. Once activated, PRP will eventually form a hydrogel, which has been reported to secrete GFs [15].

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Thrombin is the most powerful platelet activator. It induces platelet degranulation and release of GFs and cytokines. Autologous thrombin can be found in human serum [16] and therefore used to induce platelet activation at the right moment by adding serum to the platelet-rich preparation. Although bovine thrombin has been used by some authors [17], this is to be avoided through the risk of antibody formation against thrombin, prothrombin, and factor V leading to an induced factor V deficiency [18]. This deficiency can lead to increased bleeding due to the impossibility to form a stable blood clot (factor V is involved in the formation of thrombin). Calcium chloride is a citrate inhibitor and allows the plasma to coagulate. Thrombin causes fibrin to polymerize into an insoluble gel, platelets then degranulate and release GFs [19]. When calcium chloride and thrombin are combined with PRP, a gel or scaffold matrix is produced. Endogenous type I collagen has been found to be equally effective as thrombin in activating platelets and stimulating the release of GFs [20].

19.1.3.5 Storage Stability experiments demonstrated that a storage period of 6 hours at 20 C may be compatible with preservation of platelet functionality and GF release capacity, whereas longer storage duration causes the decrease of GFs concentration within platelet granules [14]. It is known that a decrease in pH due to accelerated production of lactic acid in hypoxic conditions is a major cause for loss of platelet viability [21]. Agitation of the platelet concentrates ensures that the platelets are continuously oxygenated that sufficient oxygen can enter the storage container and that excess carbon dioxide can be expelled [22]. A temperature between 21 C and 24 C is recommended for the storage of platelet concentrates in order not to activate the platelets [23]. Optimal storage conditions should provide a sufficient amount of O2 (21%) thanks to constant agitation, a temperature between 21 C and 24 C, and not exceed 6 hours at rest. The objective of this chapter is not to report an exhaustive list of these protocols but to describe the general principles used to obtain PRP products from whole-blood samples.

19.2

Tissue engineering

19.2.1 An autologous cell culture supplement During cell expansion, animal-derived serum additives, above all fetal calf serum, are used as a medium supplement in the majority of laboratory-standard operating protocols. This common use leads not only to ethical issues but also considerable immunological and contamination risks of using animal-derived products for human cell culture [24]. Therefore alternative products have been studied, including human autologous platelet lysates (PL) [25]. Its safety has been studied, and

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Fernandez-Rebollo et al. were able to show considerable expansion of bone marrow
derived MSCs without significant changes in the DNA methylation profiles, cytoskeletal organization, or focal adhesion [26]. But despite the encouraging results of PL as a cell culture supplement, there is no standardization yet in the composition (amount of plasma, range of GFs) or in the manufacture, which makes it difficult to evaluate. Autologous PRP is another alternative for cell culture supplementation and might be even more effective than conventional medium additives and PL. Bieback et al. [27,28] demonstrated the use of thrombin-activated PRP (tPRP) for isolating and expanding human MSCs. They demonstrated that pooled human serum and tPRP provide a significantly higher proliferative effect on adipose tissue
derived MSCs than fetal bovine serum. The optimal concentration of PRP as a cell culture medium needs to be determined. Atashi et al. incubated adipose tissue
derived MSCs (ASCs) in autologous 1%, 5%, 10%, 20%, 40%, and 60% PRP for 10 days and assessed the proliferation [29]. They showed that 20% nonactivated PRP exceeds the proliferation-promoting effects of thrombin-activated preparations and the effects of 40% and 60% PRP. Furthermore, they demonstrated that autologous nonactivated PRP is able to promote ASCs proliferation while maintaining their phenotype, differentiation potential, and chromosome stability, which show the safety of PRP as a cell culture supplement. Like in PL applications, standardization and exact description of the composition of the PRP product is required to gain consistency in study results. The safety of PRP and its efficiency to promote cell expansion were demonstrated. According to several studies, the optimal PRP concentration as media supplement is between 10% and 20% with media changes every 48
105 hours [30,31]. Although its efficiency is undeniable, cellular mechanisms induced by PRP still need to be fully understood.

19.2.2 Platelet-rich plasma in tissue-engineered constructs PRP is used for various tissue-engineering applications in the context of bone, cartilage, skin, and soft tissue repair. Although PRP is an inexpensive and immunologically safe source of GFs, the tissue regeneration efficacy of PRP alone has been controversial [32,33] because of the rapid inactivation and initial burst of GFs contained in PRP. During physiological wound healing an important role of extracellular matrix is to preserve GFs in an active state and release them when necessary [34]. The objective of PRP constructs is to preserve GFs activity and secrete them at a chosen site and time. It has been combined with different biopolymers not only to improve the mechanical properties of PRP-based materials but also to render the biopolymers bioactive and mimic native extracellular matrix or for the release of GFs. This may represent important biological and economic advantages, as recombinant GFs are very expensive and may promote immunogenic reactions [35]. However, knowing that all platelet-derived products are subject of large donor variation, it is difficult to draw conclusions on the efficiency of PRP in these approaches. A constant fact is the strong correlation between the concentration of released GFs and the platelet concentration in these products [21].

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PRP has been widely used in the field of bone tissue engineering as a cell carrier and/or as a source of osteogenic and angiogenic factors. In this purpose, PRP has been combined with various biomaterials, including biodegradable gelatin hydrogels [36], composite of porous bone mineral and bio-guide membrane [37], or ceramic materials such as hydroxyapatite-β-tricalcium phosphate scaffold (Skelite) [38]. Hokugo et al. [36] reported that the combination of PRP with these gelatin hydrogels had the potential to enhance bone repair in vivo. Rabbit PRP with a mean platelet concentration of 1200 3 103 platelets/μL (sixfold enrichment) was dropped onto freeze-dried gelatin hydrogels; this gelatin hydrogel incorporating PRP was applied to a bone defect of rabbit ulna. The osteoinductive properties of the composite material were compared to gelatin or PRP alone or a combination of PRP with fibrin, with most efficient new bone formation in the PRP
gelatin group. The main hypothesis was that the gelatin captures PRP-derived GFs, which are released from the hydrogel in concert with hydrogel degradation. PRP has been used as a carrier of cocultures of MSCs and endothelial progenitor cells in order to promote neovascularization in bone tissue
engineered constructs [39]. El Backly et al. demonstrated the proangiogenic properties of PRP in a PRPderived periosteal substitute [40]. In this study a membrane consisting of PRP (1000
2000 3 103 platelets/μL) and MSCs was prepared and optimized in terms of vascular endothelial GF and platelet-derived GF-BB release, aiming for the attraction of endothelial cells and osteoinductivity. This membrane was tested in vivo in both ectopic mouse and in a rabbit segmental bone defect model providing evidence of its capacity to biomimic a periosteal response enhancing bone regeneration [40]. The effects of PRP were also investigated in osteochondral and cartilage repair where PRP was mostly utilized as a carrier for chondrocytes or progenitor cells (Fig. 19.6). For example, Xie et al. [41] published a comparative experiment, testing PRP-delivered bone marrow MSCs and adipose stem cells (ASCs) to evaluate their regenerative potential for osteochondral repair. While PRP (mean platelet concentration 1670 3 103 platelets/μL) alone did not possess a regenerative capacity, PRP-delivered MSCs of both origins were shown to mediate cartilage matrix formation in rabbits at 9 weeks postimplantation. After activation, PRP forms a soft hydrogel [21] and therefore represents a suitable material for fat and skin repair and as an adhesive for wound closure. PRP has been immobilized on gelatin microspheres to improve the engraftment and the vascularization of transplanted fat upon subcutaneous implantation [42]. A mixture of PRP with heparin-conjugated fibrin guaranteed a sustained release of GFs for the treatment of full thickness wounds in mice [43]. In this study, it was shown that heparin-conjugated fibrin captured PRP-released GFs to guarantee a sustained release of GFs. Spano` et al. [44] reported that a PRP-based bioactive membrane generated by the combination of PRP, cryoprecipitate prepared from PPP and thrombin, can serve as a bioactive adhesive to induce wound healing by increasing the thickness of the regenerated epidermis and the vessel number in a diabetic mouse chronic ulcer model. In conclusion, PRP has been used for various tissue-engineering applications with different outcomes. In these studies, PRP is utilized either as an autologous

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Figure 19.6 Components and structure of PRP scaffold. (A) Gross appearance of PRP scaffold. (B) Hematoxylin
eosin staining of cryosection of PRP scaffold showed a mesh-like microstructure with platelets and mononuclear blood cells entrapped within the fibrin skeleton (left: low magnification and right: high magnification). (C and D) Scanning electron microscopy of PRP constructs: (C) PRP scaffold and (D) PRP scaffold seeded with MSCs. Left: low magnification and right: high magnification [41]. MSCs, Mesenchymal stem cells; PRP, platelet-rich plasma.

hydrogel cell carrier or as a source of GFs supporting cell attraction, angiogenesis, or direct tissue regeneration. Several strategies are described to sustain the release of GFs over a longer time by capturing the PRP-released GFs in gelatin- or heparin-conjugated fibrin hydrogels. Once again, PRP is often applied as part of a composite biomaterial, which makes it difficult to analyze the single effect of the PRP. Its efficiency in tissue engineering is promising but still needs to be specified by a standardization of the manufacturing protocols and components analyses.

19.3

Platelet-rich plasma in regenerative medicine

The clinical use of PRP is based on the increase in the concentration of GFs and in the secretion of proteins that are able to improve the healing process at a cellular level. Since PRP is an autologous biologic material, it involves a minimum risk of immune reactions and transmission of infectious diseases. Therefore it has been widely used for the recovery of several lesions. First in orthopedics and sports medicine where PRP found many applications: in osteoarthritis (concerning either knee [45] or hip [46]), which involves cartilage damage related to an inadequate response in the inflammatory environment. This pathology leads to a progressive destruction of the cartilage in the articulation leading to chronic pain and lameness. Tendinopathy or tendon injuries can develop in

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any tendon of the body and can occur in three areas: tendon insertion, mid-tendon, and musculotendinous junction. The common symptoms are pain and stiffness of the closest joint and eventually a local inflammation. PRP has been evaluated in the treatment of patellar [47] (knee), rotator cuff [48] (shoulder), and Achilles [49] (ankle) tendinopathy. The epicondylitis is also a form of tendinopathy concerning the insertion of the muscles on the lateral epicondyle of the elbow. This condition is most frequently referred to as “tennis elbow” [50]. Other conditions have led to the use of PRP in orthopedics such as plantar fasciitis [51] which is a degenerative tissue condition of the plantar fascia leading to heel pain, or muscle injuries [52], the most common concerning hamstrings. PRP has also been widely used in oral and maxillofacial surgery, mostly in order to improve bone grafts, which are performed in order to fill a bone defect following surgery or trauma [53]. It has also been evaluated in the repair of total cleft lip and palate which results from a lack of fusion of the medial nasal prominences during fetal development. PRP has been applied between the oral and nasal mucosa during palatoplasty [54] in order to improve postsurgery healing in this challenging location. The indications in plastic surgery are multiple and essentially cover the improvement of wound healing for chronic or challenging wounds. The healing can be difficult because of the poor local vascularization like in diabetic foot ulcers [55] or because the treatment requires a large tissue excision in an area exposed to important moves and tension, not favorable to proper healing. This is what happens in pilonidal sinus disease [56] which is a sinus or an abscess diagnosed in the midline natal cleft in the sacrococcygeal area with a distance of 5
8 cm away from anus. This disease comes with comorbid disorders associated with pain and discomfort in these patients, usually leading to interruption of both personal and social life of the patients. The objective of the treatment is to stop the pain and heal the wound as quickly as possible in order for the patients to be able to recover an acceptable quality of life. PRP has also been evaluated in more esthetic indications like the improvement of hair growth and density in androgenetic alopecia [57]. Other applications of PRP have been reported in ophthalmology for the treatment of severe dry eye [58], also in gynecological disorders such as cervical ectopy or cervical erosion due to hormonal changes. Unexpected bleeding or vaginal infections may occur, therefore a healing treatment can be necessary [59,60]. Veterinarians often use PRP as well, as reported by DeRossi et al. in the improvement of wound healing in the surgical wounds in equine. Table 19.2 sums up the results of these several studies, covering most of the main applications of PRP in regenerative medicine without being exhaustive because of the tremendous amount of studies on this subject. PRP is a promising treatment for some diseases; however, evidence of its efficacy has been highly variable depending on the specific indication. Additional high-quality clinical trials with longer follow-ups will be critical in shaping our perspective of this treatment option.

Table 19.2 Examples of clinical applications of platelet-rich plasma (PRP) in randomized controlled trials. Medical field

Type of PRP

Delivery route

Number of patients

Control group

Outcome

Favors PRP?

References

Knee osteoarthritis

LP-PRP

Intraarticular injection

1423

[45]

LR-PRP

Intraarticular injection

230

1

[50]

Patellar tendinopathy Plantar fasciitis

LR-PRP

Intraarticular injection Local injection

23

1

[47]

430

Ultrasound-guided dry needling alone Corticosteroids

PRP injections are more efficacious in terms of pain relief and self-reported function improvement compared with other injections Significant improvement in pain at 24 weeks and lower percentage of patients reporting elbow tenderness Earlier improvement of symptoms

1

Lateral epicondylitis

Placebo, hyaluronic acid, ozone, corticosteroids (metaanalysis) Bupivacaine

1

[61]

Rotator cuff tendinopathy Osteoarthritis of the hip Achilles tendinopathy Muscle injuries

LR-PRP

Subacromial injection Intraarticular injection Intra-tendinous injection Local injections

40

Placebo

PRP exhibited better efficacy than the steroid treatment after 24 weeks but no statistical difference after 4 or 12 weeks of treatment Improvement in pain

1

[48]

100

Hyaluronic acid (HA)

1/ 2

[46]

54

Placebo

2

[49]

80

Placebo saline solution

Functional improvement and pain reduction but not superior to HA No clinical or ultrasonographic superiority of PRP No significant difference in return to play time or with re-injury rate at 6 months

2

[52]

Orthopedics

NR

NR NR NR

Oral and maxillofacial surgery Bone graft Repair of complete cleft palate

NR NR

Topical Local application

11 44

Autogenous bone alone No application

Favors bone formation in grafted bone PRP can decrease the incidence of oronasal fistula and improves the grade of nasality and velopharyngeal closure

1 1

[53] [54]

NR

Local injection

25

Placebo: saline solution

Increases hair density

1

[57]

NR

Local application Local application

448

/(Metaanalysis)

1

[55]

110

Classic wound dressing

Superior healing rates and lower complication rates Faster healing process, less pain, shorter antibiotic consumption duration

1

[56]

PRP 1 HA

Intraocular injection

30

HA

Reduction in corneal staining and improvement in tear parameters

1

[58]

NR

Local application

120

Laser treatment

Shorter mean time to reepithelialization in PRP group and lower rate of adverse treatment effects

1

[60]

PRP gel

Local application

6

No treatment

More rapid epithelial differentiation and enhanced organization of dermal collagen

1

[62]

Plastic surgery Androgenetic alopecia Diabetic foot ulcers Pilonidal sinus disease

PRP gel

Ophthalmology Severe dry eye

Gynecology Cervical ectopy

Veterinary Wound healing in surgical wounds in equine

Green, Positive effect of PRP, Red, no superior effect of PRP; HA, hyaluronic acid; LR-PRP, leukocyte rich-PRP; LP-PRP, leukocyte poor PRP; NR, not reported.

492

19.4

Biomaterials for Organ and Tissue Regeneration

Conclusion

In the past 20 years PRP application has proved to be safe and effective in experimental animal models and in the clinical treatment of human patients. Its therapeutic use in regenerative medicine has reported favorable effects in connection with safety and a simple methodology of preparation. Other important advantages relate to the low cost of obtaining the preparations of PRP and the lack of immunological reaction risk, PRP being an autologous product. PRP is a promising alternative to animal-derived supplements based on the efficacy of autologous GFs to accelerate tissue healing. The use of autologous PRP is now an effective alternative in regenerative medicine and tissue engineering. Concerning biomaterials based on PRP or combined with PRP and its derivatives, they possess attractive bioactivity but weak mechanical properties, which is why new strategies to improve in vitro and in vivo stability of these materials are being explored. Nonetheless, depending on the preparation technique, PRP composition may be affected. This means that more research is required for the development of standardized protocols as well as an effective categorization of the protein and GF content. New research is required, with greater accuracy and experimental refinement, with a larger number of patients, randomization of samples, standardization of methodological procedures, as well as a consensus to establish more precisely the conditions for which the PRP should be properly and precisely employed.

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81.

Section 3 Emerging and enabling technologies for biomaterials in tissue regeneration

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Nanocomposite hydrogels for tissue engineering applications

20

Azadeh Mostafavi, Jacob Quint, Carina Russell and Ali Tamayol Department of Mechanical and Materials Engineering, University of Nebraska
Lincoln, Lincoln, NE, United States

20.1

Introduction

Current clinical therapies for the treatment of organ failure, tissue loss, and traumatic injuries are based on the use of auto- and allografts [1]. Such therapies have shown to be insufficient, and researchers have been on the search for developing new strategies to solve this major medical challenge [2]. Tissue engineering is a strategy in which biologically functional constructs are fabricated to replace the diseased tissue or to enhance its function [3]. Tissue engineering approaches traditionally rely on the use of scaffolding materials that can be seeded with cells prior to implantation or to recruit host cells and direct them to form a new functional tissue within the patient’s body [4]. Thus engineering proper scaffolds have been at the center of research efforts in the field. A proper scaffold should mimic the mechanical, physical, and chemical properties of the native tissue [5]. It should also offer an environment rich with biological factors essential for recruiting cells and directing their growth and function during the maturation time [6]. The scaffold should facilitate the transfer of the nutrients and oxygen to cells and should recapitulate the architectural features of the native tissues [5]. Obviously, the selection of suitable materials is important in engineering proper scaffolds. Various biomaterials and fabrication technologies have been used for creating tissue engineering scaffolds. The extracellular matrix (ECM) of native tissues is a network of various proteins, polysaccharides, and glycosaminoglycans (GAGs) with high water content capable of modulating the cellular functions and chemical transport around the cells [7,8]. One class of materials that offers similar properties to native ECM is hydrogels, which are formed from a cross-linked network of hydrophilic polymers with high water content [7]. However, to create scaffolds suitable for tissue engineering applications, usually the physical, chemical, and biological properties of the hydrogels should be modified [9]. The composition, polymeric concentration, and cross-linking density of hydrogels can be changed to tailor their mechanical properties, pore size distribution, and degradation profile [10]. To further improve their biological environments, an array of technologies has been utilized to create hydrogel systems that offer an environment in which cells can properly interact with the network and biological factors that are available to direct cellular functions [11]. These strategies include the functionalization of hydrogels Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00023-4 © 2020 Elsevier Ltd. All rights reserved.

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with peptides, proteins, creating interpenetrating network systems, and the addition of micro/nanofeatures [12
14]. The emergence of nanotechnologies has led to the realization of simple or complex nanomaterials that can interact with polymeric chains, control release drugs/ biological factors, penetrate into cells, and modulate their biological responses [15
18]. In this chapter, we will focus on the strategies for fabrication of nanocomposite of hydrogels with a focus on their physicochemical and biological properties.

20.2

Conventional hydrogels and their limitations

Hydrogels can be formed from a vast array of source materials to mimic various features of native ECM: correct structural support, topography, spatial and temporal modulation of chemical and physical cues, proper oxygenation, and nutrition/waste transport. Hydrogels are generally classified based on their source material origin (natural, synthetic, or hybrid), polymeric composition (homopolymeric, copolymeric, multipolymer, or interpenetrating polymeric), type of cross-linking (chemical or physical), physical appearance (matrix, film, microsphere), network electrical charge (nonionic, ionic amphoteric, zwitterionic), etc. [19
21]. Hydrogels generated from natural sources generally have properties differing from synthetically derived hydrogels. Each class offers an array of advantages and disadvantages. Natural hydrogels formed from metabolically active environments (plants, animals, algae, and microorganisms) usually induce relatively limited toxicity. They are usually biodegradable through enzymatic mechanisms, namely, hydrolysis and oxidation [19,22]. Generally, bioactive, most natural hydrogels can promote cell adhesion, cell growth, proliferation, and differentiation [23,24]. However, some natural hydrogels such as alginate require further functionalization to allow for cell binding and cell interaction. Unfortunately, most natural hydrogels suffer from batch-to-batch variation and can lead to transmission of zoonotic pathogens. Recently, using recombinant DNA technologies, purity, and elimination of such transmission have been improved upon. Further, some of the proteins in natural systems can be recognized by the immune system and induce inflammation. Synthetic hydrogels have been used for their long shelf life, strong mechanical properties, ease of tuning, availability of raw materials, and ease of manufacturability and processing. Because they are not native, synthetics usually do not have associated biodegrading enzymes leading to slower degradation. Depending on the material, they can elicit immune and inflammatory responses. However, at the same time, some of the synthetic hydrogels such as poly(ethylene glycol) (PEG)-based hydrogels are considered biocompatible [7,25]. Each base material has intrinsic properties that characterize its utilization in tissue constructs. To improve upon these fundamental properties, hybrid hydrogels have been investigated to merge the benefits of two or more materials while limiting the disadvantages of each [24]. These hybrids can be a combination of natural polymers, synthetic polymers, or a combination of both. Furthermore, the properties of each material can be altered through their synthesis, cross-linking density, polymer concentration, and microstructure.

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20.2.1 Natural polymers The two major categories of natural polymers used in hydrogels for tissue engineering are polysaccharide and protein based.

20.2.1.1 Polysaccharides-based hydrogels Polysaccharides are molecules consisted of sugars chained together to form various link types, shapes, patterns, and molecular weight which impact their physical properties: solubility, viscosity, etc. [26,27]. As a major component of native ECM, polysaccharides are found in numerous renewable sources in a multitude of forms and have several physiological functions. Therefore they have been widely used in many tissue engineering, regenerative medicine, and drug delivery applications. The most common of these are alginate, chitosan, hyaluronic acid, and cellulose [28]. Other polysaccharides that have been explored but to a lesser extent include dextran, chitin, gellan gum, etc. Alginate (algin or alginic acid) is a linear block copolymer originally derived from sea algae [26,27]. It is most commonly ionically cross-linked by replacing Na1 ions with “2 1 ” ions (Ca21 or Br21) [29]. Ionic cross-linking can occur quickly and has led to its use in fiber-based techniques. Through the use of chelators, the cross-linking can be reversed making alginate desired for cell cultures where recovery of the cells is needed or to act as a sacrificial support structure for other hydrogel constructs (Fig. 20.1) [10,27,29]. While ion leaching from an alginate construct occurs when placed in physiological buffers, cross-linking with PEG diamines prevents such leaching and allows for tunability of mechanical properties [30,31]. Alginate has been widely used because of its stability, fast cross-linking, low immunogenicity, ease of use, and availability. It also has been used for its ability to encapsulate and control release macromolecules when placed in variable temperature and pH environments [27,32]. However, it lacks cell-binding moieties, has a small pore size distribution, and cannot undergo enzymatic degradation [26]. Alginate has been used in a variety of applications such as bone, cartilage, intervertebral disk, adipose, cardiac, and neural regeneration [33]. Notably, because of its unique encapsulation abilities, alginate has found a role in the protection of neural stem cells and pancreatic islets [33]. Another popular polysaccharide, chitosan, is derived from chitin, a biopolymer found in the shells of marine crustaceans and the cell walls of fungi [26,27]. It is semicrystalline with its degree of crystallinity based on its level of deacetylation. Therefore the mechanical properties can be tuned by adjusting deacetylation. Chitosan can be used to create porous gels through an internal bubbling process [26]. Its mechanical properties can be further manipulated through pore size, degree of cross-linking, alteration of its hydroxyl and amine backbone, and molecular weight [27,34]. Chitosan is used for its biodegradability, availability, biocompatibility, tunable mechanical properties, ease of processability, antibacterial nature (derived from its positive charge), and electrostatic interactions with anionic negatively charged molecules such as GAGs [26,27,35]. Chitosan has been utilized in a

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Figure 20.1 Alginate and GelMA hydrogel networks. (A) Fabrication schematic of alginate as a sacrificial network where alginate is initially ionically cross-linked to provide structure, then the secondary polymer is cross-linked, and finally the alginate is ionically leached. (B) Wet spinning fiber-based constructs by (I) immersion of extruded fibers into a CaCl2 bath and (II) secondary polymer cross-linking via UV photoinitated or chemical crosslinking. (C) Bioprinting constructs by (I) spatial printing and alginate cross-linking by CaCl2 mist and (II) secondary polymer cross-linking via UV photo-cross-linking. (D) Mechanical characterization of Young’s modulus for varying concentrations (2, 1, 0.5% w/v) of alginate with 10% (w/v) GelMA ( P , .05,  P , .01,  P , .001). (E) Micrographs of microbeads encapsulated in cross-linked alginate:GelMA (2:10% w/v) fibers (I) before and (II) after removal of alginate. (F) SEM images of cross-linked alginate:GelMA (2:10% w/v) fibers (I) before and (II) after removal of alginate. (G) FTIR spectra analysis of alginate removal by EDTA treatment comparing untreated and treated alginate:GelMA (1:10% w/v) fibers, alginate (1% wv), and GelMA (10%). GelMA, Gelatin methacrylate. SEM, scanning electron microscope; FTIR, fourier-transform infrared spectroscopy; EDTA, ethylenediaminetetraacetic acid. Source: Credit: Adapted from Tamayol A, et al. Hydrogel templates for rapid manufacturing of bioactive fibers and 3D constructs. Adv Healthc Mater 2015;4(14):2146
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variety of applications including bone, cartilage, skin, liver, trachea, and neural tissue regeneration [27]. Hyaluronic acid (hyaluronan) is a major component of native ECM and is classified as a GAG [26,27]. It plays physiological roles in synovial fluid, the umbilical cord, and the vitreous humor. Further, it is involved in the regulation of cellular behavior during development, angiogenesis, cell migration, ECM remodeling, wound healing, and inflammation response. It is highly used in tissue engineering for its hydrodynamic characteristics (namely, viscosity and water retention ability), porosity, malleability, biocompatibility, and cell-binding moieties (CD44, RHAMM). However, its use in pure form is limited by high degradation rates and solubility in water and aqueous solutions at room temperatures. Despite these limitations, hyaluronic acid has found use in skin, blood vessel, valvular, cartilage, and bone tissue engineering [27,36]. Cellulose, the most abundant biopolymer, is the main component of plant cell walls [26,27]. Also found and generated in bacteria, its long polymer chains accumulate to form microfibrils. The arrangement of these filaments determines its macroscopic chemical and mechanical properties. Cellulose is highly reactive because of hydroxyl groups on its polymer chain. This enables chemical alteration through substitution. Due to its availability, biocompatibility, and chemical manipulability, it is used for tissue engineering hydrogels [37]. Its disadvantages include its poor solubility in common solvents, poor dimensional stability, biodegradability dependent on cellulases (not natively found in mammals), and lack of antimicrobial properties [26,27,37]. The advantages of cellulose have led to its use in bone, cartilage, vascular, cardiac, and pancreatic tissue engineering [27,38,39].

20.2.1.2 Protein-based hydrogels Proteins are the most abundant organic molecules of the intercellular components and ECM [27]. They are responsible for structural support and many cellular functions. Generally, more complex than polysaccharides in structure, proteins can be described by their primary (amino acid sequence), secondary (polymer backbone formation), tertiary (spatial organization with inclusion of side chains), and quaternary structure (subunits if made up of multiple polypeptide chains) [27]. The folding configuration of the protein determines its functional properties because of interaction with complementary protein structures. The shape and functionality of a protein can be lost due to denaturation caused by excessive heating, extreme pH, sensitivity to sonication, chemicals, etc. [27]. The most common proteins used to form hydrogels for tissue engineering applications are collagen, fibrin, and elastin. Other less commonly used proteins are soy and silk-based materials. Collagen is the most abundant protein in mammalian tissue and is the primary component of ECM [27,40]. It primarily maintains the structural integrity of the ECM but also regulates storage and release of growth factors. Over 20 distinct types of collagen have been discovered [27,40]. It is composed of three polypeptide chains wrapped around one another to form a triple helix. These then bind together to form fibrils that can be further assembled and cross-linked into more complicated

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geometries and networks. Generally derived from animal tissues, they can be a source of zoonic pathogenesis, but discovery of marine sources and creation from human recombinant technologies have overcome this disadvantage. Beyond its wide availability, collagen is known to have low inflammatory and immunological responses. It can be chemically, physically, or enzymatically cross-linked. It is readily degraded by matrix metalloproteinases (MMPs) (primarily collagenase), and without further modification, it is rapidly disintegrated. While natively strong, processing and formation of collagen into hydrogels are generally noted with poor mechanical characteristics and low thermal stability. Due to the advantages of collagen hydrogels, they have been used in a wide variety of tissue engineering applications: cardiovascular, vascular, valvular, cartilage, dermal, corneal, nerve, etc. [27,40]. Gelatin is a protein that is formed from the hydrolysis of collagen and has improved solubility and is less antigenic than collagen [41]. Gelatin is ideal to form hydrogels because of its ease to cross-link and wide number of functional groups [41]. While not as strong as collagen, the mechanical properties of gelatin can be finely tuned by controlling the cross-linking density, gelatin concentration, and chemical modifications [2,41]. A common way to make gelatin light polymerizable is through the addition of methacrylate groups, yielding gelatin methacrylate (GelMA) [42]. Photo-cross-linkable gelatin-based materials are easy-to-pattern and allow creating the spatial organization of different cell types including fibroblasts, skeletal myoblasts, cardiac side population cells, and endothelial cells [2]. Combination of photo-cross-linkable gelatin and other materials allows for the creation of complex structures and to enable unique manufacturing techniques. For example, GelMA-alginate composites have been used to enable immediate ionic cross-linking to retrain structural form, secondary photo-cross-linking, and chelation of alginate to leave a GelMA only construct (Fig. 20.1) [10]. Gelatin hydrogels have found particular use in soft tissue such as cardiac, skin, and skeletal muscular tissue engineering [2,41]. Fibrin is the fibrous protein monomer whose main role is as a scaffold material surrounding platelets in blood clots [2,40]. Found in blood plasma and generated from the cleavage of fibrinopeptide via thrombin, it has been heavily used as a biomedical sealant. It is used in tissue engineering because of its porosity, biodegradability, elasticity, unique polymerization mechanism, storage capabilities, manufacturability, adhesiveness, mechanical support, and role in cell attachment, proliferation, and angiogenesis. It is mostly limited by its rapid degradation in vivo [40,43]. Fibrin hydrogels are further limited by immune response and infectious disease transmission (if heterologous sourced) and the difficulty to synthesize large quantities of thrombin and fibrin. Fibrin has been used in neural, cartilage, bone, cardiac, and vascular tissue engineering [40]. Elastin, the second most common structural component of the ECM, provides elastic properties, the ability to elongate and recoil after stretching [27,40]. It is natively found in many tissues: cardiovascular, lung, intestinal, dermal, and muscle tissues. Its elasticity is determined by the number of tropoelastin monomers and the number of coils that are in each fiber [44]. It plays a role in the regulation of

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arterial development, cell proliferation, migration, focal adhesion formation, and differentiation. It is plagued by its difficulty to process and purify, tending to calcify upon implantation. It is also generally insoluble when in its natural form. Tropoelastin, elastin’s soluble precursor, has been produced using human recombinant technologies [27,40]. Tropoelastin possesses hydrophobic regions—responsible for its elasticity and cross-linking domains. It can be formed into matrices by chemical, enzymatic, pH induced, and physical cross-linking. Because of its unique elasticity, elastin has been used in cartilage and vascular tissue regeneration [27]. Silk fibroin, created by the silkworm Bombyx mori and numerous spider species, is a fibrous protein that has been used in a wide range of biomedical applications in many differing forms: membranes, films, fibers, sponges, textiles, hydrogels, etc. [40,45]. Originally used as a suture material, pure silk fibroin has been used for its strength, elasticity, minimal immune response, and biodegradability [40]. Silk fibroin also enables cell adhesion through nonspecific surface interactions but can be further functionalized through coating with specific binding moieties [40]. Hydrogel silk fibroin has been used as an injectable matrix to induce programmed tumor cell death [45], enhance bone regeneration in critical-sized defects [46], and ocular drug delivery [47].

20.2.2 Synthetic polymers Hydrogels generated from synthetic sources offer high water content and microstructures similar to naturally sourced hydrogels while allowing several of their properties to be tuned during synthesis. Common, synthetic hydrogels include elastin-like polypeptides (ELPs), polyurethanes (PUs), PEG, poly(vinyl alcohol) (PVA), and poly(2-hydroxy ethyl methacrylate) (pHEMA) [2]. ELPs are a class of polymers that possess chains of repeating amino acids [2]. This class of polymer hydrogels mimics native ECM. They have been widely used in the regeneration of vascular, cartilage, ocular, and liver tissues [2]. They are desirable for tissue engineering applications because of the possibility for tuning their degradation rates, offering biologically relevant elasticity, and an overall low toxicity. They, however, do lack cell-binding ligands that should be incorporated through chemical processes. Most lack mechanical stiffness, but lysine-containing ELPs have been shown to yield higher mechanical strength, support cellular penetration, and have aided in ECM remodeling. However, these materials have shown faster degradation rates in comparison to other types of ELPs. Because ELPs are a class of wide-ranging materials, not all of them have been classified and analyzed for biocompatibility or immune response. PUs are an established synthetic material that has been used to generate hydrogels. They have been previously used in cardiovascular applications and are known to be stable in vivo for long periods of time [2]. PUs are used for their processability, toughness, durability, ease of surface functionalization, flexibility, and biocompatibility [2,48]. While usually known for its stability, PU tissue-engineered constructs generally demand some biodegradation. Therefore a new class of PUs has been generated to allow for tunable biodegradability while maintaining

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mechanical durability. This has been achieved through hybridization with other polymers (caprolactone, lactides, amino acids, PEG, etc.) [2]. However, some of these hybrid biodegradable PU-based materials have resulted in mechanical failures and cytotoxic by-products. PU hydrogels have been used for vascular, neural, bone, and cardiac tissue engineering [2,48]. PEG is a synthetic polymer known for its biocompatibility, immunocompatibility, ability to encapsulate cells, and for allowing ECM deposition as it degrades [7]. PEG can be cross-linked using many different methodologies (chemically, covalently, photopolymerization) to which the extent determines its properties (elasticity, degradability, diffusivity, etc.) [25]. Because of its ease of modification, high biocompatibility, and tunability of properties, it has been used for tissue engineering applications. It does not yield integrin binding, not biodegradable without modifications and is antifouling [7,25]. This can be designed to direct cellular functions during its synthesis, but without such alterations, cell interaction with the scaffold is limited. PEG hydrogels have been applied in stem cell differentiation, encapsulation of pancreatic cells, and the regeneration of cartilage, bone, and vascular tissues [25,49]. In one particular example, degradable MMP-sensitive PEG hydrogels were made through polymerization of PEG and adhesive ligands to support neovascularization [50]. A 10% polymer weight of the MMP-sensitive PEG hydrogels showed enhanced endothelial tube formation (Fig. 20.2C) and angiogenesis in transgenic mouse cornea (Fig. 20.2E) [50]. PVA is also used to form hydrogels. Known for its biocompatibility, it is a transparent synthetic polymer that has been used in a wide variety of tissue engineering applications: corneal tissue, cartilage, cardiac valves [51]. Its mechanical properties can be mechanically tuned to that of native ECM. As with many of the synthetic polymers, PVA also lacks cell adhesive moieties. It is limited in application by a tedious and complicated manufacturing process that generally requires freeze
thaw cycles. pHEMA is a synthetic polymer that forms hydrogels [52]. Known for its use in contact lenses, keratoprostheses, and ocular implants, pHEMA is used for its tunable mechanical properties through formulation chemistry and its ability to mimic native tissues [53]. It is antifouling and does not have cell-binding moieties unless further functionalized. Also, it has limited degradability without hybridization with other polymers. pHEMA hydrogels have been employed in vascular, optical, cartilage, and cartilage tissue engineering [52,54]. Although generally hydrogels offer high porosity, favorable transport properties, and tunable mechanical properties, they do have various limitations affecting the possibility of creating a tissue biomimetic environment. The mechanical properties such as tensile and compressive moduli of hydrogel systems are usually different from many load-bearing tissues [2]. Hydrogels, as a bulk, do not offer topographical cues to direct cellular organization. Although this problem can be alleviated by microfabricating hydrogel constructs into scaffolds with highly organized architectures, still the organization cannot easily be preserved throughout the culture time. Most of the hydrogel systems lack biological factors that can initiate or facilitate physiological and biological processes necessary for tissue formation and

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Figure 20.2 Proteolytically degradable PEG hydrogels for enhanced angiogenesis. (A) Schematic illustrating the fabrication and components of MMP-sensitive PEG hydrogels via polymerization of PEG and cell adhesive ligands. (B) MMP-sensitive PEG hydrogel degradation profiles for varying polymer weight percentages in collagenase (10 μg/mL). (C) Endothelial formation in MMP-sensitive PEG hydrogels analyzed through investigation of (i) tubule length and (ii) number of branching points over 6 days of culturing. (D) Confocal (D3 and D6) and bright-field (D6) images of HUVECS (green) and 10T1/2 (red) cells in MMPsensitive PEG hydrogels of varying polymer weight percentages. (E) Confocal images of 10% polymer weight MMP-sensitive PEG hydrogel implanted in a transgenic mouse corneal promoting neovascularization as shown by (i) mCherry tracking of vascularization and (ii) depth profiles from z position relative to hydrogels. The arrow in (ii) vascular infiltration into the hydrogel. MMPs, Matrix metalloproteinases; PEG, poly(ethylene glycol). Source: Credit: Adapted from Moon JJ, et al. Biomimetic hydrogels with pro-angiogenic properties. Biomaterials 2010;31(14):3840
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maturation. Hydrogel systems are typically nonconductive to both ions and electrons and thus do not offer an environment suitable for the function of electrophysiologically active tissues [55]. Engineering nanocomposite hydrogels with the use of suitable nanomaterials have shown to be effective in addressing these challenges by altering the microstructure of the hydrogels and imparting characteristics of the nanomaterials into the bulk properties of the hydrogels. In the next sections, we will discuss various nanomaterials used in tissue engineering applications and then will discuss how nanocomposite hydrogels with tailored physical and biological properties have been developed for tissue engineering applications.

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Nanomaterials for engineering composite hydrogel systems

As mentioned earlier, despite many favorable characteristics of hydrogels, they have limitations in mimicking the tissue properties. Thus there have been numerous research efforts in modulating various properties of hydrogel systems such as mechanical moduli, electrical conductivity, and the presence of biological factors. Incorporating nanoparticles within hydrogels has been a popular approach for tailoring their physical and biological properties because of their high surface-area-tovolume ratio, tunable release profile, and the ability to target cellular components [45]. Nanocomposites can be fabricated in two ways: either prefabricated nanofeatures can be mixed with the hydrogel prepolymer and then cross-linked or the features can be built in situ within an already formed hydrogel network [45]. Nanocomposite hydrogel systems can be tailored based on the type of nanoparticle, shape, concentration, and the method of functionalization.

20.3.1 Methods for creating nanocomposite hydrogel systems 20.3.1.1 Incorporation of prefabricated nanomaterials An elementary approach to create nanocomposite hydrogel systems is to cross-link polymeric network in a nanoparticle suspension. In this approach, prefabricated nanoparticles are suspended in a hydrogel prepolymer; the solution then goes through a gelation process to create the final composite material [56]. This technique has been used with various particles and hydrogel systems. For example, Sershen et al. showed that by varying the optical resonance the nanocomposite hydrogel systems can photothermally stimulate the hydrogel matrix to undergo a driven volume collapse [57]. The biggest setback with this approach is the prevalent leaching of nanoparticles out of the hydrogel matrix. For this reason, other techniques have been developed to address this issue. Nanocomposite hydrogel systems can also be formed with the use of crosslinking groups on the surface of nanoparticles. The benefit of using this method is that multiple bonds can be formed in the hydrogel matrix in comparison to the hydrogel generation using two covalent bonds [56]. For example, Moreno et al. took carboxylic functionalized gold nanoparticles acted as cross-linkers with PVA hydroxyl groups creating an esterification reaction [58]. Similarly, Eguchi et al. successfully cross-linked enzymatically synthesized DNA with gold nanoparticles using the Au
S bond to form a gel network [59]. The enzymatically synthesized DNA acted as a linear polymer, while the gold nanoparticles served as crosslinking sites to create a 3D hydrogel matrix [59]. Prefabricated nanoparticles can also be added postgelation of the hydrogel matrix. Normally gels swell when introduced to aqueous solutions and shrink when exposed to an aprotic solvent [56]. The “breathing” method is a three-step method that first involves shrinking the hydrogel and then introducing it to an aqueous medium dispersed with nanoparticles. The nanoparticles are then taken up by the

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swollen hydrogel; this process can be repeated until a desired density of the nanoparticles is reached. This adaptation demonstrated the ability to enhance uniform distribution of nanoparticles. Nanoparticles can also be added to hydrogels postgelation through a repeated cycle of heating, centrifuging, and redispersion of a preformed colloidal mixture of nanoparticles within a micro-gel; this cycle ends with a final step of annealing [56]. The major challenges associated with the use of prefabricated nanoparticles include the lack of control over the particle distribution within the hydrogel network, resulting in variations in the local physical and biological properties of the composite system.

20.3.1.2 In situ hydrogel conversion There are two different in situ nanofunctionalization techniques: (1) synthesis of the nanoparticles during hydrogel formation and (2) the synthesis of nanoparticles within a cross-linked hydrogel matrix. During the cross-linking process of a hydrogel, chemical reductants can act as a cross-linker for specific hydrogels as well as a synthesizer for nanoparticle formation [60]. A crucial benefit of this first approach is that it is highly cost and time efficient, it also eliminates the use of toxic chemical reductants and cross-linkers [60]. The second in situ technique involves a twostep process where a precursor solution is loaded into the hydrogel after which metal ions undergo a reduction process to form nanoparticles within the hydrogel [60]. For both of these in situ techniques, nanocomposites containing silver [61,62], platinum [63], and gold [64] nanoparticles have successfully been synthesized. In one study, Xiang and Chen used carboxylic acid to anchor silver nanoparticles onto the hydrogel backbone and create pH-responsive systems [61]. The ionic reduction of the ions in the silver nanoparticles inhibited the ability of the nanoparticles to aggregate [61].

20.3.2 Nanoparticles in tissue engineering There are a wide variety of nanoparticles such as ceramic-, metallic-, carbon-, lipid-, and polymeric-based nanoparticles that have been used in tissue engineering applications. These particles fall into two broad categories of soft and hard particles, each with significant benefits and limitations. The following nanomaterials have been widely used in tissue engineering applications. Silicon-based nanoparticles: Silicon-based materials can take numerous forms. One of which, mesoporous silica nanoparticles (MSNPs) have ordered hexagonal pores with a narrow pore size distribution. Consequently, these particles offer a high surface-area-to-volume ratio in comparison to other nonporous nanoparticles [71]. MSNPs are positively charged and as a result negatively charged molecules can be loaded into them, as seen in Fig. 20.3A [72]. They offer a closer to linear release profile in comparison to other polymeric systems. MSNPs can have tunable pore size and excellent biocompatibility. Applications for MSNPs in tissue engineering have been primarily focused around scaffolds for bone tissue engineering.

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Figure 20.3 Structure of nanoparticles commonly utilized for the functionalization of hydrogel systems. (A) MSNPs can be utilized as carriers for a number of other polymers, drug delivery molecules, and florescent imaging molecules [65]. (B) Illustration of gold nanoparticles multiple applications for biomedical therapies [66]. (C and D) Carbon-based nanoparticles used for hydrogel functionalization: graphene, GO, rGO (C) [67], and carbon nanotubes (D) [68]. (E and F) soft nanoparticles used for functionalization of hydrogels: dendrimer (E) [69] and liposomes (F) [70]. MSNPs, mesoporous silica nanoparticles. Source: Credit: Adapted from Moreira AF, Dias DR, Correia IJ. Stimuli-responsive mesoporous silica nanoparticles for cancer therapy: a review. Microporous Mesoporous Mater 2016;236:141
57; Ghosh P, et al. Gold nanoparticles in delivery applications. Adv Drug Delivery Rev 2008;60(11):1307
15; McCoy TM, et al. Graphene oxide: a surfactant or particle? Curr Opin Colloid Interface Sci 2019;39:98
109; Sajid MI, et al. Carbon nanotubes from synthesis to in vivo biomedical applications. Int J Pharm 2016;501(1):278
99; Huang D, Wu D. Biodegradable dendrimers for drug delivery. Mater Sci Eng: C. 2018;90:713
27; Khorasani S, Danaei M, Mozafari MR. Nanoliposome technology for the food and nutraceutical industries. Trends Food Sci Technol 2018;79:106
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Gold-based nanoparticles: Nanofunctionalized hydrogels with gold nanoparticles have been studied for bone and cardiac tissue engineering. Gold nanoparticles or colloidal gold can be prepared in different shapes (most commonly spheres and also nanorods [73,74], nanocubes [75], and nanostars [76,77]) and sizes ranging from 3 to 200 nm [60]. Gold nanoparticles continue to be of particular interest because they exhibit excellent antiinflammatory and biocompatibility characteristics [66]. Gold nanorods (GNRs) have been used in applications for wound adhesives, tissue welding, as well as cardiac tissue regeneration [73,74]. Gold nanoparticles also can be functionalized as nanocarriers, illustrated in Fig. 20.3B, for biomolecules or drugs for tissue repair and regeneration. Silver-based particles: Silver nanoparticles exhibit excellent electrical conductivity and antimicrobial capabilities, mainly in its oxidized form, and their incorporation in hydrogels prevents biofilm formation [78]. These characteristics make the incorporation of silver nanoparticles ideal for tissue engineering applications where resistance to a multitude of bacterial strains is needed to ameliorate the tissue healing process. Recently, many researchers reported cytotoxicity or inhibitory effects of silver particles, and these inhibitory effects are typically more pronounced in smaller particles due to their higher surface-area-to-volume ratio. Silver nanoparticle formulation has been extensively studied using all nanofunctionalization techniques, mentioned previously in this section, in a variety of hydrogels. However, further research must be conducted for a better understanding of cell interaction and toxicity effects. Quantum dots (QDs): QDs are nanoscale semiconducting crystals. These nanoparticles are of interest because of their similar size to biological macromolecules [79]. These nanoparticles can be manufactured in sizes ranging from a few nanometer to a few micrometers, and their exact size can be easily controlled with proper annealing temperatures and growth techniques [79]. QDs have notable photophysical and semiconducting properties that are ideal for bioresponsive and encoded hydrogels [80]. One major disadvantage that comes with the use of these particles is their cytotoxicity; therefore tailoring QDs to the precise size and regulating their dosage in hydrogels is critical. Carbon nanotube (CNT) particles: CNTs, shown in Fig. 20.3D, have been actively researched because of their high aspect ratio, electrically conducting capabilities, mechanical properties, and low density [81]. A notable advantage of CNTs is their ability to be manipulated to structurally align within a hydrogel. Many of their mechanical characteristics directly affect porosity, electrical conductivity, surface roughness, and cellular activity within hydrogels. These attributes help create biomimetic scaffolds. The major disadvantage that accompanies the use of CNTs is that this material can induce chronic toxicity and are not bioresorbable [82] (cf. Chapter 21: Functional carbon-based nanomaterials for engineered tissues towards organ regeneration). Graphene-based particles: Graphene-based particles are another class of carbonbased particles possessing a planar structure, seen in Fig. 20.3C. Nanomaterials in

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the graphene family differ based on the number of layers or chemical modification; for biomedical applications, graphene oxides (GOs) or multilayered graphene is chosen [83]. GO is a chemically modified graphene that has the ability to be functionalized or to be used as a drug carrier [83]. Due to its favorable mechanical and physiological properties, hydrophilicity, and high dispersity, it has become a highly researched nanoparticle as an alternative to CNTs [84,85]. Graphene-based particles, such as CNTs, are not biodegradable and have the inherent risk of toxicity [86]. One of the many benefits of the reduced versions of GO is that the noncovalent interactions increase the absorption of ECM proteins [87]. Minimally layered reduced GO (rGO) can enhance cell proliferation and adhesion, and has been found to be a viable option for tissue repair and drug delivery applications [83]. To decrease cytotoxicity effects, GO and rGO can be functionalized with biocompatible polymers. Nanoliposomes: Liposomes are naturally soft, round shaped, remarkably organized, continuous vesicles formed from one or several phospholipid bilayers dispersed in an aqueous medium, as illustrated in Fig. 20.3F [88,89]. Nanoliposomes are nanometric liposomes of the submicron size, which inherently provide more surface area [89,90]. Nanoliposomes contain other molecules that aid in the stability of their lipid bilayer [91]. These particles can be fabricated from organic sources, resulting in increased bioavailability and biocompatibility. Nanoliposomes are amphiphilic; therefore they are able to increase the in vivo and in vitro stability for a number of active molecules through encapsulating them in the lipid bilayer or a central aqueous cavity [90]. One issue with nanoliposomes is that there is a possibility that the encapsulated molecules leak due to passive diffusion [92]. Dendrimers: Dendrimers are small polymeric molecules that are symmetrically branched and homogeneous with a linear polymer core, shown in Fig. 20.3E [93]. Dendrimers are constructed of dendrons, the size of the dendrimer is dependent on the number dendrons; this branched architecture makes it characteristically and behaviorally different from linear polymers [94]. Dendrimers are widely used in nanomedicine as a result of them being devoid of toxicity and their ability to cross cellular barriers [95,96]. Dendrimers have a wide variety of applications, gaining notable attention as a drug delivery vehicle because of their compact molecular structure and number of functional groups [93]. Moreover, it is possible to control their functionalization, polyvalence properties, bioactivity, and biocompatibility. Polypeptide and polyester dendrimers have been used as scaffolds for corneal tissue laceration repair [97,98]. Functionalized dendrimers have also shown the ability to modulate the behavior of stem cells and osteogenic differentiation [99]. Hard nanoparticles have exceptional mechanical properties, electroconductivity, and prolonged stability while in storage. Nevertheless, many hard nanoparticles are attributed with cytotoxicity and are nonbiodegradable. Soft nanoparticles are highly tunable, environmentally responsive and are generally biodegradable [100]. The combination of hard and soft nanoparticles has led to novel developments in the tissue engineering.

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Properties of nanocomposite hydrogels

In this section, we will discuss how various physical properties (mechanical, structural, and electrical) and biological properties (bioavailability of drugs, factors, gene delivery, and cellular reprogramming) can be achieved by incorporation of nanofeatures.

20.4.1 Tailored mechanical and structural properties Mechanical properties of hydrogels at both cell and scaffold levels affect the cellular response and successful integration of the scaffold with the tissue. It is widely accepted that the changes in the substrate stiffness can affect the differentiation of stromal cells or affect cellular organization. On another level, the bulk properties of the material such as its strength and tensile or compressive moduli can be translated into the properties of the fabricated scaffold and can affect the survival of scaffold postimplantation. For example, skeletal muscles typically go through a significant contraction ( . 20% strain) and have a mechanical strength of several MPa. Thus if a transplanted scaffold does not meet these criteria, the scaffold can rupture, dislocate, or detach after its implantation. Biologically relevant hydrogels typically are not strong enough to meet such criteria, and significant efforts have been made to find ways to improve their mechanical properties. One popular approach has been the incorporation of hard particles. Upon the incorporation of hard polymers into a soft hydrogel, its tensile modulus and stiffness increase, while usually its elastic strain is reduced. Examples of hard nanoparticles that have been incorporated into hydrogel systems with the goal of changing their mechanical properties are CNTs, GOs, metal oxides, and gold and silver nanorods as described before. In one example, CNTs were mixed with GelMA and the tensile modulus of the hydrogel system was increased from 10 to 32 kPa [101]. However, the incorporation of such nanoparticles disrupts the polymeric network and introduces weak points within that, resulting in a lower tensile strain. In one interesting example, Annabi et al. fabricated human-based tropoelastin hydrogels reinforced by the incorporation of GOs and rGOs (Fig. 20.4) [55]. As expected, the tensile modulus of the hydrogel was increased from 12.6 to 19.3 kPa by the addition of 1% (w/v) GO [55]. They also demonstrated that the fabricated nanocomposite hydrogels showed excessive resilience against torsional forces and cyclic tensile loading. A surprising observation was the enhancement in the tensile and rotational strain of the nanocomposite hydrogels. They demonstrated that nanocomposite hydrogels could tolerate 27 rounds of torsional strain in comparison to pristine tropoelastin that could only tolerate 18 rounds of strain. The observed trend was related to the structure of GO particles and their tendency in forming hydrogen bonds with various points of the tropoelastin networks increasing the elasticity of the structure. Soft nanoparticles can also change the mechanical properties of nanocomposite hydrogel systems. In one example, Wang et al. developed hydrogels with high

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Figure 20.4 Formation and mechanical characterization of MeTro/GO hydrogel. (A) MeTro undergoing uniaxial tension. (B) The expandable MeTro molecule that can attach to GO particles, the mechanism behind the hydrogel’s elasticity. (C) SEM image of porous composite hydrogel. (D) Images showing torsion test of MeTro/GO and MeTro hydrogels showing recovery after 10 rounds noting significant deformation in MeTro hydrogel in comparison to the hybrid. (E) Schematics showing the structure of hybrid hydrogel before and after torsion demonstrating the role of GO nanoparticles and MeTro fibers to respectively act as physical bonding connectors and elastic contributors to the composite. (F) Stress
strain curves from composite hydrogel and MeTro gel. (G) Mechanical tests showing that GO particles can increase the rupture strain and ultimate strength. GO, Graphene oxide. Source: Credit: Adapted from Annabi N, et al. Highly elastic and conductive human-based protein hybrid hydrogels. Adv Mater 2016;28(1):40
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toughness, outstanding stretchability, and good self-healing ability by incorporating acylhydrazone bonds and Pluronic F127 (PF127) micelle cross-linking, as two types of dynamic cross-links, in the uniform system [102]. The improvement of mechanical properties of the hydrogel allows it to stretch more than 117 times of its initial length. Furthermore, imine bonds can be formed from aldehyde end functional groups of PF127 micelle with amine groups in proteins and enhance their biological properties such as cell attachment and expansion in the hydrogel [102]. In another work, Wang et al. reported a novel hydrogel combination of poly(amidoamine) dendrimers and poly(lactic acid)-b-PEG-b-poly(lactic acid) copolymers that showed an extremely interconnected porous network with enhanced mechanical stiffness and reduced swelling ratio [103]. These improvements in this hydrogel’s physical properties resulted in a better attachment of mesenchymal stem cells (MSCs), proliferation, and differentiation more similar to natural 3D ECM structures [103].

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The incorporation of nanofeatures affects the global and local mechanical properties of the materials at the same time. In case the changes within the local properties of the hydrogel system are not amenable for cellular functions and biological response, then nanofeatures should be utilized to reinforce the global properties of the scaffold without affecting the localized properties cells are experiencing [104]. In one example, Akbari et al. fabricated thread reinforced composite hydrogel fibers in which threads formed from nanofilaments were coated by a layer of cell-laden hydrogel system [105]. Nanofibrous substrates generated using electrospinning were also used to mechanically reinforced hydrogel systems in a way that the scaffolds could mimic the mechanical properties of tendons while the soft hydrogel layer could offer properties suitable for cellular proliferation [106]. The rheological properties of hydrogel systems are also an important factor in their use. For example, injecting hydrogels has been considered as a robust method for the delivery of scaffolds for treatment of tissue injuries and diseases [107]. In this approach, shear-thinning materials are typically used which can be further cross-linked postinjection. One of the methods for the fabrication of shear-thinning materials has been the incorporation of nanoparticles. Nanoclays are a class of nanomaterials that has been frequently used to create shear-thinning hydrogel systems [108]. Nanoclays are typically positively charged and once incorporated within a negatively charged hydrogel, the electrostatic interactions and the formed weak bonds result in the formation of shear-thinning materials. Such hydrogels have also been used for arterial embolization [109] as well as creating shearthinning bioinks for 3D bioprinting [110].

20.4.2 Enhanced electrical conductivity Electrical conductivity of scaffolds has shown to have an effect in cellular response and the maturation of cultured cells especially for engineering electrophysiologically active tissues such as neural, cardiac, and skeletal muscle tissues [111,112]. Proper signal propagation and the synchrony in responding or transferring physiological signals are key in the successful function and integration of any newly developed tissues. Although the electrical signals in the human body are more based on ion transfer, it is believed that the use of electrically conductive materials can facilitate the transfer of such signals [113]. Regular hydrogels conventionally used in tissue engineering and regenerative medicine are usually electrically nonconductive and thus are not suitable environment for transferring electrical signals. Once cells are incorporated, electrical conductivity can affect the tissue maturation and function in two ways: (1) at the tissue level, having scaffolds with high electrical conductivity can facilitate the signal propagation within the scaffold and connecting the tissue to the host nervous system and the surrounding tissues [55]; (2) at cellular level, incorporation of small electrically conductive passages between cells might facilitate cell
cell interaction and improve cellular organization and tissue maturation [114]. As a result of these, two different needs researchers have looked into various methods to create hydrogels with only localized (pore-to-pore) or global conductivity (throughout the scaffold).

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Among different approaches for rendering hydrogels locally or globally conductive, the incorporation of conductive micro and nanoparticles has attracted much attention. The addition of metallic or conductive carbon-based particles within polymeric network can facilitate the conduction of electrons through the network. However, if the particles do not have proper contact, the generated resistance can be high and reduce the electron conductance rate in comparison to conductive or semiconductive materials [114]. The problem is more severe for spherical particles where the conduction length of electrons is similar to their diameter and may not be sufficient to connect two adjacent pores. Adds value by providing an example of elongated particles and their ability to negate the lacking conductivity of spherical nanoparticles. In addition, if the concentration of particles is high enough, there is high probability of particles touching each other and increasing the travel distance of electrons. In one example, GNRs were incorporated within alginate hydrogel cultures with cardiomyocytes [114]. It was shown that the incorporation of GNRs did not change the bulk electrical conductivity, but the AFM measurements indicated that localized electrical conductivity was changed at the sites with nanorods. The biological data confirmed a positive role of nanorods on the attachment and maturation of the seeded cells [115]. In another example, GNRs were incorporated into hydrogel to create electrically semiconducting bioink. The bioinks were then used for 3D printing of hydrogel scaffolds, which were successfully used in cardiac tissue engineering (Fig. 20.5) [115]. CNT-based particles have also been incorporated within GelMA-based hydrogels without changing the global conductivity of the scaffolds [101]. The cultured cardiomyocytes showed a lower excitation threshold for synchronized beating [101]. In an interesting example, rGO and GO particles were incorporated with MeTro hydrogels, and it was shown that rGOs can change both local and global electrical conductivity. The presence of the rGO and GO significantly improved the maturation of cardiac tissues [55]. Another interesting experiment showed that the electrically conductive materials can facilitate the signal propagation throughout two pieces of freshly harvested abdominal muscles and their synchronized contraction by applying electrical signals [55]. Shin et al. developed electrically conductive CNT-based inks that could be printed in 2D and 3D configurations on different substrates such as hydrogels and paper [116]. The conductive hydrogels could also be embedded within hydrogel structures to create 3D electrical circuitry. The mechanical, electrical, and biological evaluation of different constructs generated from these inks showed suitable biocompatibility for use in biological application and are able to improve the maturation of electrophysiological-conductive tissues similar to cardiac patches (Fig. 20.6) [116].

20.4.3 Enhanced availability of biological factors and drugs Well-defined microstructure, connected porous networks, and appropriate mechanical properties of hydrogels are important factors for generating a construct that supports cellular growth. However, biological processes such as cell recruitment, proliferation, and differentiation in the body are typically regulated by biological

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Figure 20.5 Preparation of GNRs in GelMA printed hydrogel and the composite’s material and biological characterization as cardiac tissue constructs. (A) Schematic of coating GNRs with GelMA. (B) UV
vis spectra for uncoated (C-)GNRs and GelMA coated (G-)GNRs showing that the coating technique did not change their aspect ratio. (C) Zeta potentials of CGNRs and G-GNRs compared with GelMA showing that the particles were successfully coated. (D) (i) Schematic of incorporation of G-GNRs in GelMA hydrogel, (ii) TEM image from composite that shows homogeneous GNR distribution in the gel, (iii) different concentrations of GNRs in gel. (E) Impedance over a frequency spectrum showing a positive relationship between concentration of GNRs and impendence, showing increased (Continued)

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factors present in the environment [117]. Different scaffolding materials such as hydrogel systems are not usually capable of sustaining the level of biological factors [118]. Numerous research efforts have been devoted to finding methods for enabling hydrogel systems to sustain the level of biological factors during the culture time. Biological factors can easily be mixed with the precursor prior to crosslinking of the hydrogel system [117]. However, often a significant portion of the encapsulated factor can escape the pores of hydrogel system within hours to a few days [119]. This timeline might not be sufficient for inducing the planned biological process. Biological factors and drugs can also be attached to the polymeric backbone of hydrogel systems to modulate their release profile and their bioavailability [120]. This process has been done for a limited number of factors and cannot easily be applied for other therapeutics. In addition, while the loading efficiency of drugs using this approach is limited, the release rate of the therapeutics attached to the polymeric backbone cannot be tailored. One common strategy for overcoming these challenges includes the use of drug carriers capable of controlled release of biological factors and drugs [121]. Nanoparticles offer higher surface-area-to-volume ratios in comparison to their micron-sized counterparts and thus have been the popular selection [122]. Nanogels, polymeric nanoparticles, MSNPs, liposomes, and carbonbased systems are among the most common drug carriers used for functionalization of hydrogel systems. In one study, Posadowska et al. increased drug (sodium alendronate) releasing time by encapsulating them in poly(lactide-co-glycolide) nanoparticles and suspending them in a gellan gum matrix [123]. Their experiments showed the injectability of the hydrogel and its stability after extrusion. They also demonstrated the sustainable drug delivery inhibiting osteoclast differentiation as a potential treatment for bone disorders [123]. However, polymeric systems in general suffer from a quick initial burst release followed by a gradual release. This trend is not ideal for tissue engineering applications, where a sustainable level of biological factors is desired throughout the culture time. In a recent study, VEGF was loaded into nanoclays that were then mixed with collagen hydrogels [124]. The hydrogel systems were then implanted in a murine critical-sized femoral defect for 28 days and showed enhanced angiogenesis. In another example, stem cell conditioned media containing their secretomes were mixed with nanoclays to absorb the compounds [125]. The particles were then conductivity with increasing GNR concentrations. (F) Effect of GNR concentration on Young’s modulus resulting in an increasing stiffness relationship with concentration of GNRs until 0.25 mg/mL. (G) The dry weight of cell encapsulated and printed gel constructs during culture. (H) The topography of the printed constructs. (I) Immunostaining of Cx43 (red), sarcomeric α-actinin (green), and DAPI (blue) showing G-GNR effect on the enhancement of cardiac cells maturation. (J) Percentage expression of Cx43 in fluorescent area. (K) Spontaneous beating rates of the printed cardiac tissue constructs. GNR, Gold nanorod; GelMA, gelatin methacrylate. Source: Credit: Adapted from Zhu K, et al. Gold nanocomposite bioink for printing 3D cardiac constructs. Adv Funct Mater 2017;27(12).

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Figure 20.6 3D printed electrically conductive ink constructs and their cellular and electrical characterization in and out of GelMA hydrogel. (A) Schematic illustration of GelMA/DNAcoated MWCNTs and bonding and interactions between them. (B) Actin/DAPI staining of cardiac fibroblasts cultured on ink printed patterns on PET film coated with PEG. (C) Current
potential curve of fiber network inside GelMA hydrogels. (D) Cell viability assessment after 200 charge and discharge cycles by live/dead assay. (E) Cardiac tissue formation showed by immunostaining of Cx43 (red), sarcomeric α-actinin (green), and cell nuclei (blue). (F) Incorporation of printed fiber network inside GelMA hydrogels and its top view. (G) Immunostaining of Cx-43, sarcomeric α-actinin, and cell nuclei of encapsulated cardiomyocytes in GelMA hydrogel with electrical conductive fibers inside the gel. GelMA, gelatin methacrylate; MWCNTs, multiwalled carbon nanotubes; PEG, poly(ethylene glycol). Source: Credit: Adapted from Shin SR, et al. A bioactive carbon nanotube-based ink for printing 2D and 3D flexible electronics. Adv Mater 2016;28(17):3280
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mixed with GelMA, and the positive role of the encapsulated stem cell secretomes on inducing angiogenesis and protecting cardiomyocytes against reactive oxygen species was demonstrated [125]. Nanoclays have shown to be able to achieve a linear release profile for over 1 month. Chen et al. developed a pH/thermosensitive hydrogel from polyacrylic acid and poly N-isopropylacrylamide that contained functionalized MSNPs in the gel. These silica nanoparticles were loaded with doxorubicin as an anticancer drug and bovine serum albumin as a model protein drug. The in vitro results of this study showed the capability of the hydrogel to sustain

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the release of doxorubicin resulting in the death of the cancer cells. Also, they maintained a time releasing of the bovine serum albumin for helping the surrounding tissues to regenerate, which can be a promising strategy for tumor therapy [126]. Carbon-based particles such as GO and CNTs have also used for controlled release of drugs and biological factors. Derkus et al. investigated the efficacy of protein (myelin basic protein) immobilization in different hydrogels such as gelatin, alginate, chitosan, and their composites both with and without CNTs [127]. Their results showed that the highest protein immobilization was achieved in hydrogels containing CNTs. It was suggested that this observation was due to the electrostatic interaction between the CNTs and their high surface area. The highest protein immobilization was an improvement of about 3.6 times and was achieved in gelatin
chitosan
CNT composite due to the existence of active groups in both hydrogel backbone and CNTs [127]. Liposomal systems have also been frequently used as drug carriers within hydrogel systems. Liposomal systems are capable of encapsulating both hydrophilic and hydrophobic drugs. In one example, Mahajan et al. developed a hydrogel system for localized delivery of paclitaxel for tumor and cancer therapy [128]. They loaded nanoliposomes with paclitaxel and incorporated them into thermosensitive chitosan
dibasic sodium phosphate hydrogels. The in vitro drug release data showed 72 hours extended release time. In vivo studies demonstrated 89.1% 6 3.5% tumor volume reduction that displays the efficacy of this carrier-based hydrogel model for sustained and localized drug delivery [128]. A key difference between drug carriers from hard materials and those from soft materials is the simultaneous changes in the physical properties. For example, the addition of nanoclay particles has shown to affect the rheological properties of the gels, or the incorporation of CNTs can affect the electrical conductivity and mechanical properties of the hydrogels. Thus if the changes in physical properties are not desirable, electrically insulated soft nanoparticles are a better option in comparison to their hard nanoparticle counterparts.

20.4.4 Cellular reprogramming As discussed in the previous section, biological factors and drug delivery can affect cell proliferation, differentiation, and maturation as well as the interaction of host body with the neotissue. However, there are several challenges for local growth factor delivery such as safety of manufacturing and formulation of the factors, their short half-lives, preserving their full bioactivity after their incorporation within delivery systems, and precisely controlling their release profiles in the specific body site. One potential solution would be to program cells to produce the targeted factors locally or to directly deliver plasmids that can direct cellular differentiation without the need for the presence of various biological factors. Plasmid’s DNA structure is stable and can be used in different drug delivery systems; also, they usually do not induce systemic toxicity or localized side effects [129].

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Soft particles such as liposome and other fatty-based nanoparticles have shown to feasibly interact with cell membranes and become internalized [130] (cf. Chapter 13: Diabetic wound healing with engineered biomaterials). Thus such particles are attractive for engineering nanocomposite systems capable of gene delivery. Wu et al. engineered an advanced alginate hydrogel by combining muco-inert PEGylated lipoplexes containing siLamin A/C to assist siRNA delivery into the vaginal epithelium [131]. Their approach exhibited a sixfold increase in the PEGylated lipoplexes uptakes by the vaginal epithelium compared to common lipoplexes intravaginal administration. Results showed significant knockdown of lamin A/C in the vaginal epithelium in mice after lipoplexes delivery which presented a novel strategy with high efficacy for nucleic acid delivery for cervical cancer treatment [131]. In another study, Walsh et al. improved gene delivery by using polyamidoamine dendrimer as GFP gene carrier in 2D and 3D collagen-based scaffolds. They observed an enhancement in transfection of MSCs [132]. Hard nanomaterials have also been used for the intracellular delivery of nucleic acids. In one study, Wang et al. developed a collagen hydrogel-based scaffold containing DNA-polyethylenimine (PEI)-silica nanoparticles [133]. Fibroblasts were then encapsulated within the scaffolds. Genes and plasmids are typically negatively charged and can attach to positively charged PEI and silica nanoparticles. Their results showed that the cells were transfected successfully in the hydrogel and also able to sustain the generation of biomolecules after 1 week resulting in an effective gene delivery system to assist wound healing [133]. Paul et al. developed an injectable GelMA hydrogel by incorporating functionalized GO by PEI which bound to a proangiogenic gene for producing VEGF after transfection to myocardial cells [134]. They injected the composite hydrogel into the myocardial infarcted region of a rat model. Their results showed higher capillary density and a decreased scar area in infarcted myocardium in comparison to other groups that were GelMA and GelMA-DNA. Furthermore, heart functionality was evaluated after 21 days postinjection by echocardiography that was significantly higher in GO-DNA GelMA composite than other groups. This work suggests a promising composite injectable hydrogel for gene delivery for ischemic heart disease [134]. Engineering nanocomposites capable of transfecting cellular cultures with the right plasmids and nucleic acids is an emerging concept, and it is expected to make significant impact in tissue engineering and regenerative medicine.

20.5

Conclusion and future directions

Tissue engineering has emerged as a promising approach for engineering biologically relevant constructs capturing the physiological functions of native tissues. This is expected to expand the use of tissue-engineered constructs beyond a therapeutic tool to a diagnostic tool and even personalized disease models. Fabricating scaffolds that facilitate tissue formation and proper interaction of the new tissue and the host body is a central need for the success of tissue engineering approaches.

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Hydrogels have been a popular material choice due to many favorable properties such as high porosity, high water content, hydrophilicity, and in some cases the ability to recapitulate part of the native ECM composition and biological ligands. However, at the same time, pristine hydrogels suffer from lack of proper mechanical properties, insufficient level of biological factors, the inability to direct cellular fate, and the inability to conduct electrons and ions. Thus a special attention has been paid to engineering hydrogel systems that offer all or some of these characteristics and several different strategies and numerous new hydrogel systems have been developed and tested for tissue engineering applications. The use of nanomaterials has been a popular approach, and various nanomaterials have been incorporated into hydrogels to form nanocomposite systems with superior properties. Despite many promising observations, these strategies have remained limited in terms of fabricating fully functional tissues and translation into clinical practice. This limitation is caused by the lack of proper understanding of nanomaterials with the targeted cells and tissues as well as the immune system. Cytotoxicity and the use of slowly degradable or nondegradable nanomaterials have been another major challenge that has remained a road block against their approval by FDA. We believe substantial research should be dedicated into the understanding of the interaction of particles with different composition, geometry, electrostatic charge, and degradation profile with cells and the immune system at both acute and chronic scales. In addition, improving the specificity and predictability of the changes in the nanocomposite properties and their interaction with the host body are also of great importance. Another area that requires improvement is the use of materials that are biologically degradable. Recently, there have been several studies looking into the use of DNA-based nanoparticles for drug and plasmid delivery. These particles can be degraded through enzymatic processes and are claimed to not leave any trace in the host body. Another challenge in tissue engineering is the inability to control and direct spatial patterning of cellular growth and differentiation in 3D. The realization of hydrogel systems that can limit the ingrowth of unwanted cells or unwanted differentiation would potentially solve this challenge. When comparing various properties of nanocomposites with pristine hydrogels, it has been observed that multiple characteristics change after the incorporation of the nanomaterials. This multifactorial change has affected the predictability of the biological responses to the nanocomposite systems. Developing tools for nanoengineering of composite systems in a way that properties can be independently changed is of great importance.

Acknowledgments Authors declare no conflict of interests in this work. The financial support from the National Institutes of Health (GM126831, AR073822), the University of Nebraska, and Nebraska Tobacco Settlement Biomedical Research Enhancement Funds are gratefully acknowledged.

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Yasamin A. Jodat and Su Ryon Shin Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States

21.1

Introduction

Rapid advances in engineering techniques, biomaterials, and biocompatible electronics have led tissue engineering through massive progress toward organ regeneration and biomimetic hybrids and prostheses. From bioactuators and biosensors to three-dimensional (3D) bioprinted organs and 3D biomimetic tissues, diverse engineered tools are now available to provide a more precise in vitro model of native tissues and cell
cell interactions compared with the limited traditional two-dimensional (2D) cultures. Modeling tissues and organs in vitro is facing an enduring challenge in precise replication of the native extracellular matrix (ECM) environment to match the needs for organ-specific tissue regeneration with original functionality. In other words, each tissue in the human body comprises of specific cell types, each necessitating special biochemical cues, mechanical and biological properties to guide cell growth, signaling and function direction. To recapitulate such intricate nature of tissue ECM, using a single type of material does not often suffice, and therefore, it is necessary to engineer a combination of nano- and biomaterials, hydrogels, and functionalized scaffolds with a wide range of electrochemical, mechanical, and biological properties (cf. Chapter 25: In vitro disease and organ model). While natural and synthetic biopolymers are chosen as the core materials for artificial tissues, providing a wide range of mechanical and sufficient biological properties, nanomaterials such as carbon-, metal-, or polymer-based ones are excellent auxiliary components that can be incorporated into the main matrix to enhance the ECM of the target tissue by providing tunable biophysical, chemical, and biological properties. Excellent electrical and thermal conductivity, high mechanical strength, ease of chemical modification, and large surface area are some of the extraordinary features of carbon-based materials. Comprised of large sheets of carbon honeycomb networks, carbon-based nanomaterials can be developed into a variety of 1D, 2D, and 3D structures each endowed with unique

Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00021-0 © 2020 Elsevier Ltd. All rights reserved.

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physical, chemical, and biological properties. Here, we focus on two main categories of carbon-based materials, carbon nanotubes (CNTs), and graphene, which have been successfully used in fabricating various types of artificial tissues such as cardiac, bone, skeletal muscle, and neural tissues. Replication of the multicellular structure of different injured tissue types necessitates a specific assortment of cell types and given that many of the fully matured cells in the injury site (e.g., neurons) have limited or no regeneration capacity, external cell-based therapies are required to supply specialized cell sources. To this purpose, stem cells have proven to be exceptionally fit due to their high availability, self-renewability, and capability to differentiate into a wide range of subtypes. Through guided and controlled differentiation of stem cells, recent cell therapies are able to activate specialized cells from endoderm, mesoderm, and ectoderm. Various types of stem cells with different developmental capacities have been increasingly used in regenerative tissue engineering to provide cell sources for injured areas. In addition to multipotent stem cells, induced pluripotent stem cells (iPSCs) with distinct infinite regeneration ability can be harvested and cultured using minimally invasive methods and through reprogramming patient’s own somatic cells (cf. Chapter 7: Stem cells: sources, properties, and cell types). As such, iPSCs are highly potent in giving rise to individualized tissue repair and personalized regenerative medicine [1].

21.2

Characteristics of carbon-based materials used for tissue engineering

21.2.1 Graphene Graphene is a derivative of carbon with a single layer of 2D honeycomb network of carbon atoms with hybridized sp2 orbitals. This regular and densely packed carbon lattice structure (Fig. 21.1A) has been found to bring exceptional electrical and thermal conductivity to the nano- and macrostructure, making it suitable for conductive cell culture substrates, electrophysiology, and cell potential measurements [2]. These bonds connecting carbon atoms in the planar lattice give rise to interesting mechanical properties such as high elasticity, strong hardness (greater than diamond) (1 TPa), and fracture strength (130 GPa) [3]. This planar structure can be further modified with chemical or physical methods to create interesting properties such as those in multilayered or oxidized (graphene oxide, GO) or reduced oxide structures (reduced GO, rGO). Each derivative of graphene can be tuned for a wide range of biomedical applications. Specifically, the planar structure includes a large surface area that can be easily modified by various biomolecules, including deoxyribonucleic acid (DNA), proteins, aptamers, and enzymes, to create a controlled release mechanism of drugs, reagents, and active biomolecules. Examples include controlled release of growth factors such as basic fibroblast growth factor in wound-healing applications [4] or

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Figure 21.1 Schematic of the procedure of incorporating carbon-based nanomaterials into biomaterials to enhance tissue regeneration and scaffolding effect. (A) Structure of graphene monolayer. (B) 3D structure of SWCNTs and MWCNTs. (C) Stem cell
laden carbon-based biomaterials prepared, cultured, and differentiated into functional target cells. The cell-laden biomaterial is then implanted into the injured tissue site to provide regenerative therapy and healing. Moreover, the excellent electrical conductivity of carbon-based biomaterial allows for in situ electrophysiology of the implanted site. 3D, Three-dimensional; MWCNTs, multiple wall carbon nanotubes; SWCNTs, single wall carbon nanotubes.

sustained delivery of bone morphogenetic protein (BMP)-2 for osteogenic differentiation and bone regeneration [5]. In addition, this feature can be exploited by drug or gene-delivery vectors to increase the effectiveness of transfection procedures [6,7]. Each graphene compound exhibits different biophysical and chemical properties based on the method of synthesis, modification, and number of layers and thus creates unique interactions with the target cells and tissues [8]. Such interactions include both instructive (functionality and application) or destructive (toxicity and biocompatibility) modes of action and should be carefully considered in the biomaterial optimization stage of any tissue-engineering platform. GO is derived from chemically modified graphene through high oxidation. The structure of GO is composed of a single graphene layer with inplanar functional groups such as hydroxyl, epoxide, and carboxylic acid. These groups provide colloidal stability, pH-responsive negative surface charge, and π
π interactions with bioactive molecules [8]. The amphiphilic structure of GO makes it an excellent surfactant candidate for surface functionalization with hydrophobic biomolecules and modeling organic bilipid layers. On the other hand, GO exhibits lower

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mechanical properties and electrical conductivity due to the presence of the functional groups. To enhance the electrical conductivity, the GO can be thermally, chemically, or UV-optically processed into rGO. Endowed with exceptional electron mobility of 400 S/cm, rGO has been selected for effective implementation of tissue
electrode interaction interfaces useful in electrophysiology of neuronal and cardiac tissues [9,10].

21.2.2 Carbon nanotubes Rolling up graphene sheets into cylindrical configuration creates CNTs with single or multiple “walls” depending on the number of layers (Fig. 21.1B). Single wall CNTs (SWCNTs) have a diameter ranging from 0.4 to 2.5 nm and length ranging from 20 to 1000 nm, while in multiple wall CNTs (MWCNTs), the diameter and length range higher, 1.4
100 nm and 1
50 μm, respectively [11]. Most of the exceptional properties of graphene can also be found in CNTs, namely, high electrical and thermal conductivity, surface functionalization capability, and nanoscale anisotropy. Studies have shown that CNT incorporation into engineered tissue constructs results in improved mechanical strength, electrical conductivity without compromising the gel porosity required for nutrient delivery. Moreover, high adhesion to tissue, long-term stimulation capability, durability, and corrosion-resistance can be favorably exploited to create CNT-based electronic tissue interfaces [12]. A drawback is the lack of solubility of carbon-based nanomaterials in aqueous conditions due to strong hydrophobicity. This insolubility can induce nanoparticle aggregation in drug-delivery systems and lead to reduced biocompatibility and toxic side effects [13]. Upon aggregation, surface area of the nanomaterials is reduced and unfavorable for loading large amounts of biomolecules and drugs. Moreover, large aggregations can lead to reduced endocytosis, impacting the drug-delivery efficiency. Such aggregations can also affect the physical properties of the hybrid carbon-based scaffolds such as electrical conductivity, Young’s modulus, and elasticity. To address these issues, many studies have explored surface modification with various chemicals or biomaterials such as polyethylene glycol (PEG) [14]. Moreover, the toxicity of CNTs can be further aggravated by the small size of CNT nanoparticles which can escape the phagocytosis defense system of the body, causing potential damage to organs, for example, the respiratory system [15].

21.3

Function mimetic carbon-based engineered tissues

Although successful mimicking of the native tissue microenvironment is a crucial step for organ regeneration, the engineered tissues can be most effective only if the functionality of the native organ is fully retained. A drawback of conventional biomaterials is the lack of their ability to provide electrical conductivity, making it difficult to record electrical readouts and cell signaling and failing to represent a

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physiologically functional electroactive tissue such as cardiac, muscle, and brain tissues. An advantage of using carbon-based biomaterials in tissue engineering is the enhanced and controllable retention of functionality due to their outstanding properties as discussed previously. For instance, previous studies have shown that patterning CNTs and graphene can induce directionality and alignment in cell-laden engineered tissues [12,16]. This interesting feature has opened the way toward selective complex cell patterning with the help of lithography methods and modern techniques such as 3D bioprinting. For instance, we have shown that culturing cardiomyocytes on highly aligned microelectrode-integrated CNT
hydrogel constructs can lead to homogenous cell organization and enhanced cell
cell signaling leading to strong spontaneous muscle beatings [12]. What follows is the tissue-specific application of carbon-based materials in regenerative tissue engineering.

21.3.1 Skeletal muscle regeneration Skeletal muscle has a highly limited potential in self-regeneration and is mostly restricted to the recovery of small-scale wounds. Traumatic events such as excessive sports, combats, and surgeries can lead to large-scale abrupt loss of muscle tissue, often referred to as volumetric muscle loss (VML), eventually resulting in functionality impairment of the injured muscle and substantial scaring and tissue fibrosis. Moreover, muscle regeneration is acutely declined by aging due to the decreased renewal and differentiation of satellite cells as well as the changes occurring in the immune cells and the secretion of regeneration biomarkers [17]. Therefore seeking external regenerative resources such as those offered by tissueengineering scaffolds to restore muscle functionality is considered imperative and a recent research focal point. Muscle structure consists of 3D hierarchical architecture of highly aligned myofiber bundles which can create contractile movement with the help of sarcomere units [18]. To guide myofiber orientation and differentiation in artificial muscle scaffolds, a number of biochemical and biophysical cues such as nanoscale roughness, high organization, and alignment are required. Several engineered scaffolds such as gelatin methacryloyl (GelMA), polycaprolactone (PCL), and collagen have been used to date to provide these features [19
21]. However, pristine scaffolds lack flexibility and high conductivity and thus fail to provide sufficient support for muscle movement and functional development. Using carbon nanomaterials such as graphene and CNTs can provide such nanoarchitecture to facilitate cell binding and guidance for myofiber organization while supporting nutrient and oxygen delivery to the cells at a microscale level [22]. A sample of this multiscale hierarchy was developed by Patel et al. through a composition of interconnected microporous carbon foams and aligned CNT mats which exhibited improved myocyte fusion and maturation into multinucleated myotubes (Fig. 21.2A
D) [22]. In addition to biochemical and structural requirements, mimicking the elastomeric behavior of the native skeletal muscle tissue is also an important factor to consider in designing muscle-mimetic engineered scaffolds. While many of the developed scaffolds to

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Figure 21.2 Skeletal muscle regeneration using carbon-based nanomaterials. (A) Fiber alignment in CNT-based scaffolds enhances myoblast differentiation into multinucleated myotubes over a course of 14 days compared to pristine fiber scaffolds. (B) Immunostaining images of MHC (green) and nuclei (Hoechst, blue) show continuous multinucleated myotubes formation in CNT-based fibers. Scale bar represents 100 and 50 μm for main images and insets, respectively. (C and D) Myotube maturation index (ratio of myotubes containing 5 or more nuclei to total number of myotubes) and myotube fusion index (cells containing 2 or more nuclei/total number of cells per image); statistical significance (P , .05) between CNT-based and pristine [22]. (E) Schematic diagram of fabrication of PCE
PCEG nanocomposites to improve myoblast attachment, proliferation and mediate in vivo muscle injury repair. (F) Images of PCEG scaffolds with and without addition of rGO. The latter efficiently reinforces structural stability of PCEG nanocomposites without disturbing the hydrophilicity. (G) H&E staining images of PCEG nanocomposites 4 weeks postimplant. (H and I) PCEG with rGO (1.0) composites show higher regenerated centronucleated myofibers (H) and capillaries (I) in the injury site ( P , .05,  P , .01) [23]. CNT, Carbon nanotube; MHC, myosin heavy chain; PCE, poly(citric acid-octanediolpolyethylene glycol); PCEG, poly(citric acid-octanediol-polyethylene glycol)
grapheme; rGO, reduced graphene oxide.

date have shown successful myogenic expression and support of cell growth, most of these scaffolds cannot replicate the mechano-elastic behavior of the ECM. Recently, Du et al. developed an rGO
poly(citric acid-octanediol-polyethylene glycol) nanocomposite which showed significant improvement in myogenic differentiation, electrochemical activity, and in vivo biocompatibility (Fig. 21.2E
I) [23]. Overall, incorporation of carbon-based materials into traditional engineered construct can improve myogenesis and skeletal muscle functionality and promote the quality of scaffold-based therapies for VML injuries.

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21.3.2 Cardiac tissue regeneration Approximately one in every four annual deaths in the United States is pertained to cardiovascular diseases, making the heart disease the leading morbidity cause after cancer [6,7]. Cardiovascular diseases cover a wide range of disorders affecting the muscular tissue of the heart and the vascular system. Myocardial infarction is the most common heart disorder and a consequence of obstruction in the arteries, which leads to impaired blood supply to the myocardial muscle, creating large-scale necrosis, apoptosis, and fibrosis in the cardiac muscle tissue and eventually leading to cardiac failure and heart attack. The massive loss of cardiomyocytes often leaves irreparable scars in the cardiac tissue due to minimal self-regenerative capacity of the adult cardiac tissue [24]. Heart transplantation is considered the conventional practice but an insufficient solution, as the process critically relies on the availability of donors, a number much lower than that of patients. Tissue engineering has offered novel solutions to this problem through cardiomyoplasty surgery and graft implantation to replace the injured cardiac tissue and retain the contractile capability of the muscle. In these methods an engineered cardiac scaffold is developed by integrating cardiac primary or stem cells into biocompatible materials such as collagen, PCL, and PEG. Some studies have also shown that incorporation of CNTs into GelMA hydrogels can promote the maturation of cardiac tissues as well as an increase in the intracellular Ca21 transition compared with pristine GelMA hydrogels [25,26]. The efforts in engineering cardiac tissues are driven toward bringing electrical function to the patched cardiac tissue to regain the contractile functionality of the myocardium. It has been confirmed that electrical stimulation can effectively promote the differentiation and functionality of stem cell
derived cardiomyocytes [27,28]. However, since most of the biocompatible polymers are often insulator and have minimal electrical conductive capacity, the implanted engineered tissues fail to regenerate the unique electrophysiological behavior of the native tissue [29,30]. Integration of carbon-based nanomaterials such as CNTs or graphene into the engineered ECM can supply the matrix with local electrical conductivity which can lead to enhanced differentiation and beating function of cardiomyocytes. In a recent study a 3D MWCNT-based polydimethylsiloxane scaffold proved a potential in supporting structural and electrophysiological maturation of neonatal rat ventricular myocytes while hampering the proliferation of cardiac fibroblasts [31]. Smith et al. developed a conductive graphene
PEG hybrid to show that structural and physiological maturation of the cardiac cells can be directly impacted by controlling directional conductivity cues [32]. The micropatterning capability of CNT and graphene delivers a new asset to mimic the highly aligned microstructure of the myocardiac tissue and electrophysiological functionalities such as synchronous contraction [16,31]. For instance, chemical vapor deposition has been used to grow well-aligned CNT microelectrode arrays, which are then incorporated into a GelMA
PEG matrix to serve as a culture platform for cardiomyocytes (Fig. 21.3A
D) [12]. The cell-laden structures exhibited partially uniaxial sarcomeric structures and synchronous beating behavior of

Figure 21.3 Enhanced cardiac tissue regeneration with functional and morphological maturation using carbon-based nanomaterials. (A) Schematic diagram of 3D biohybrid actuators with a cardiac tissue layer on top of a multilayer hydrogel sheet embedded with aligned CNT microelectrodes. (B) Immunostaining images of cardiac biomarker (sarcomeric α-actinin, green and Cx-43, red) and nuclei (blue) showing organization of cardiac tissues on 5 days after culture on top of composite hydrogel layers incorporating CNT forest electrodes. (C) Interconnected sarcomeric structures of the cardiac tissue are well depicted at higher magnification. (D) Electrical stimulation of the tissue exhibited displacement of the CNT forest electrode in the multilayer hydrogel construct over time (square wave form, 1.2 V/cm, frequency: 0.5
3 Hz, 50 ms pulse width) [12]. (E) Electrically driven microengineered bioinspired soft robots fabricated using CNT
GelMA
PEG hydrogels and show contraction behavior of cardiac muscle. (F) Immunostaining images (sarcomeric α-actinin: green, Cx-43: red, nuclei: blue) of the cardiomyocytes cultured on the bioinspired robot on day 5 of culture show cardiac tissue maturation [33]. (G) The structure of myocardium can be mimicked using SA-CNTs. (H) Immunostaining image of aligned and elongated cardiomyocytes on the scaffold (Cx-43, red and nuclei, blue). (I and J) Cardiac pacing and resynchronization therapy of SC-CNTs scaffolds performed on isolated neonatal rat hearts using one-piece pacemaker electrode; scale bar, 0.5 cm. [16]. (K) 3D nanofibrous G 1 PCL scaffolds seeded with mES-CM show high cell adhesion, spontaneous contraction, and high cTnT-eGFP biomarker expression. (F-actin, red and nuclei, blue). 1,3 on PCL and 2,4 on PCL 1 G scaffolds. (l) Average beating frequency of cells on PCL 1 G compared to pristine PCL. The former shows higher frequency values than 2D and PCL scaffolds on day 6. However, average beating frequency was significantly lower on day 14 in PCL 1 G compared to day 6 (P 5 .045) [30]. 3D, Three-dimensional; CNT, carbon nanotube; G 1 PCL, graphene and poly(caprolactone); GelMA, gelatin methacryloyl; mES-CM, embryonic stem cell
derived cardiomyocytes; PCL, poly(caprolactone); PEG, polyethylene glycol; SA-CNTs, superaligned carbon-nanotube sheets.

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cardiomyocytes which could be precisely controlled via adjusting the directional electrical stimulation, perpendicular and parallel to CNT array networks. An improved model of this design was recently employed in a self-actuating bioinspired soft robot [33]. Specifically, cardiac myofibers were cultured on a CNT
GelMA and PEG-patterned substrate with incorporated Au electrodes (Fig. 21.3E and F). Such bioinspired biohybrid actuator can provide a novel platform for microelectrode-embedded scaffolding for wireless cardiac tissue engineering. Moreover, Ren et al. developed superaligned CNT one-piece electrodes with enhanced tissue attachment and effective pacing functionality and synchronous beating (Fig. 21.3G
J) which hold a promise in treating heart infarction and cardiac electrical dysfunctionality [16]. In addition to improving cell
cell behavior and cardiac contraction, embedding graphene or CNT can further provide mechanical support to the infarcted site cardiac tissue [34]. Graphene derivatives have shown to provide morphological cues to direct stem cell fate in vivo. For instance, rGO has proven effective in guiding the differentiation of mesenchymal stem cells (MSCs) into cardiomyocytes and thus improving cardiac reparation [35]. Conclusively, incorporation of carbon-based nanomaterials, namely, CNTs and graphene, into the hydrogel scaffolds has proved promising in designing engineered scaffold with enhanced cardiac tissue repair and electrophysiological characteristics similar to those of native cardiac tissue. Moreover, mechanical strength of carbonbased biomaterials is shown to be higher than most conventional engineered constructs. Nevertheless, the fabrication procedure of CNT or graphene scaffolds demands high levels of optimization and stability to yield suitable hybrid scaffolds. For instance, aggregation of CNTs during synthesis and patterning can lead to compromised electrical conductivity and mechanical stability in the final cell-laden scaffold [36]. A poor cell adhesion can sometimes be an issue when seeding cells on carbon-based hybrid substrates [33]. To address this issue, Ameri et al. recently created a 3D graphene foam scaffold capable of reinforcing cell attachment in addition to supporting cell growth and in situ electrical recording [37]. Another issue involves biocompatibility of the nanomaterials in the final cardiac tissue. Compared with CNT-laden scaffolds, graphene-based biomaterials have shown to provide higher biocompatibility for cardiomyocytes derived from stem cells. Hitscherich et al. recently developed a 3D electroactive graphene
PCL scaffold which exhibited enhanced biocompatibility by demonstrating typical cardiomyocyte phenotype and contraction behavior compared to those in nongraphene
PCL scaffolds or 2D cultures (Fig. 21.3K and L) [30]. The biocompatibility and toxicity of the carbonbased nanomaterials are further discussed in the later sections of this chapter.

21.3.3 Neural tissue regeneration The nervous system consists of intricate neuronal networks that relay electrical signals between different parts of the body and coordinate brain
organ communications. Nervous system injuries leave millions of people each year with death or permanent disability and can be categorized into central nervous system (CNS) or

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peripheral nervous system (PNS) injuries. Some examples of causes of CNS injuries include neurodegenerative diseases (e.g., Alzheimer’s and Huntington’s), spin cord injuries, traumatic brain injury, and epilepsy. Second category involves the damage imposed to the peripheral nerves such as diabetic neuropathy, acute motor axonal neuropathy, or the poststroke nerve disorders. Due to the lack of neuron regeneration capacity, regrowth or repair of the nervous tissues damaged upon disease or injury is majorly hampered. Moreover, CNS injuries face extra complexities compared to the PNS injuries as the neuron apoptosis in CNS-affected site often tends to extend to the surrounding unaffected neural areas creating sizeable cell death and severity [38]. Furthermore, CNS regeneration can be largely obstructed by the formation of postinjury glial scars which physically inhibit the axon regeneration and signal apoptosis through secretion of certain biochemical factors [39]. Improving regeneration capacity of neural tissues has been the target of several studies in the past few decades. Currently, the alternative methods for nervous tissue loss remediation involve the use of nerve autografts, allografts, and xenografts, which are challenged by high inefficiency and practical issues such as morbidities at the donor site, immunogenicity, and lack of proper functionality [40]. Alternatively, recent tissue-engineering protocols have been coupled with stem cell technology to enable a more effective regenerative therapy by providing unlimited resource of neural cells capable of replacing the damaged nervous tissue. For this purpose, different types of stems cells such as human embryonic stem, embryonal carcinoma, iPSCs, neural stem cells (NSCs), and MSCs have been programmed to differentiate into various specialized cells and subtypes in the neural tissue depending on the type of injury [41
43]. A critical factor in neuronal differentiation, growth, maturation, and functionality is the electrical signaling and between the neurons and with the other cells in the body through neural networks. Recent studies have shown that electrical stimulation of the nerve can promote stem cell differentiation and neuron regeneration to repair the injured nerve tissue [44,45]. Moreover, nanoscale topographical properties and biological cues have shown to be highly impactful in enhancing the neuronal differentiation [46,47]. Despite the advantageous applications of conventional conductive polymers for neural regeneration, these materials are often not easily tunable at the macro- and nanoscales. As such, carbon-based materials endowed with nanoscale anisotropic features as well as outstanding electrical conductivity can be apposite candidates for directing neural differentiation and nerve regeneration with higher efficiency and retention of functionality. CNTs and graphene derivatives have been used either independently [48] or were incorporated into various types of biomaterial matrices such as silk fibers [46], gelatin [49], and hyaluronic acid (HA) [50] to induce stem cell differentiation into neurons, increasing axonal outgrowth and enhancing the functionality through supporting the expansion of neural networks. Combining rGO with conductive polymers such as poly(3,4-ethylenedi-oxythiophene) has shown to regulate the differentiation of NSCs by providing suitable surface roughness functionalized with biochemical cues which contribute to the outgrowth of neurites and cell migration [51,52]. Moreover, coating the rGO with neuron-adhesive proteins such

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as poly-L-lysine, poly-D-lysine, N-cadherin, or laminin can create well-interconnected neural networks with high continuity as well as increasing cell survival for long-term cultures such as those required for regenerative organ engineering [49,53]. CNTs have shown favorable characteristics for neural differentiation such as providing a nanostructure similar to the fine processes in the neural cell growth. These nanofeatures can enhance the interactions of substrate with NSCs and guide neurite outgrowths as well as contributing to functional maturation of stem cell
derived neurons [54]. Moreover, Shao et al. have confirmed that the interactions between CNTs and NSCs can activate dynamic molecular mechanisms responsible for differentiation which, in turn, activate functions such as cell survival, synaptic development, and modulation of differentiation. In addition, it has been shown that patterning the growth substrate can direct and promote neurite outgrowth [55]. Several studies have exploited the patternable capacity of CNTs and graphene to create regular patterns using photopatterning [56,57]. 3D printing can also be a valuable asset for creating complex neuronal network-like structures and custom patterns. For instance, Lee et al. created a 3D-printed nanoconductive scaffold using MWCNTs to show early neuronal differentiation [58]. Qian et al. developed a multilayer 3D-printed polydopamine/arginylglycylaspartic acid (RGD)
functionalized graphene scaffold showing effective axonal regrowth and myelination after PNS injury (Fig. 21.4A
E) [59]. Moreover, HA
CNT
polypyrrole hydrogels were shown to remarkably promote and accelerate neuronal differentiation and upregulate the calcium channels in a spatially controlled electroconductive 3D system (Fig. 21.4F
I) [50]. In conclusion, graphene- and CNT-based nanoscaffolds have shown promising potential for nerve regeneration by supporting neuronal differentiation and outgrowth in a diverse range of stem cells. Excellent nanoscale anisotropy, high conductivity, structural robustness, and flexibility of the carbon-based nanomaterials can be practically beneficial for developing flexible neural conduits. One of the challenges in engineering carbon-based neural tissues is the inadequate biocompatibility of the biomaterial substrates as a result of poor interaction with the neural cells. This issue can adversely affect the proliferation and differentiation of the cells resulting in low cell functionality outputs. To address this issue the substrate can be modified to include proper topological features. Lee et al. have demonstrated a method to increase the adhesion and interaction of human MSCs with graphenebased substrates by modulating the defects and domain size of graphene to promote neuronal differentiation [60]. Finally, engineered neural tissues can provide valuable tools to study and repair complex and debilitating diseases such as Alzheimer’s, Parkinson’s, and amyotrophic lateral sclerosis.

21.4

Bone regeneration

As an exceptional self-reparable organ, bone can indistinguishably heal and regenerate itself upon fracture and mechanical stress without fibrosis or scar formation

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Figure 21.4 Neural tissue regeneration using carbon-based nanomaterials. (A) Schematic diagram showing the fabrication process of graphene nerve conduits [inner-most and outermost green layers are made of PDA/arginylglycylaspartic acid (RGD)]. SG or MG and PCL composite layer shown in purple. The blue layer is graphene/PCL composite repeated. The conduit is then implanted in the sciatic nerve of a rat defect model. Immunostaining images of the PDA/RGD
SG/PCL scaffolds using (B) Glial fibrillary acidic protein (GFAP), and (C) Tuj1. (D) H&E staining of Schwann cell
laden PDA/RGD
SG/PCL scaffolds showing nerve regeneration 18 weeks postimplant. (E) Transmission electron microscopy (TEM) of the mentioned scaffolds showing myelinated axons [59]. (F) Differentiation of human induced pluripotent stem cell-derived neural progenitor cells (hiPSC-NPCs) embedded in CNT HA
CA hydrogels on day 7 postencapsulation, with or without PPy. First column: immunostaining with Tuj1 (green), MAP2 (red), scale bar 5 50 μm. Columns 2 and 3: intracellular Ca21 influx in hiPSC-NPCs in bare HA
CA (CNT 0/PPy 0) and CNTembedded HA
CA hydrogels (CNT 1/PPy 30) before and after glutamate stimulation, scale bar 5 20 μm. (G and H) qPCR and gene expression of Tuj1, MAP2, and Cav1.2 in scaffold and control samples. Relative gene expression normalized to CNT 0/PPy 0 (n 5 3
4,  P , .01 vs CNT 0/PPy0, 1 P , .05 vs CNT 0/PPy 30). (I) Images of hydrogels 7 days after swelling [50]. CNT, Carbon nanotube; HA, hyaluronic acid; HA
CA, catechol-functionalized HA; PCL, poly(caprolactone); PDA, polydopamine; PPy, polypyrrole; SG or MG, single- or multilayered grapheme.

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[61]. Nevertheless, massive bone loss instigated by infection, traumatic events, insufficient blood circulation, and aging can majorly compromise this restoration capacity [62]. During the past two decades, novel bioengineering techniques have aimed to promote and expedite bone regeneration by introducing biomimetic cellular matrices vested with necessary ECM components to direct bone regeneration. Among various biomaterials, collagen, HA, and chitosan have shown notable effectiveness in inducing bone regeneration [63
65]. Major requirements for bone scaffolding include osteoconduction, osteoinduction, and osteogenesis [66]. In order to generate a functional bone tissue, the biomimetic construct should be rich in osteogenic cell-binding sites to allow cell elongation and migration. The engineered scaffold should also provide a viable stem cell resource and mimic the macro- and micro-architecture and mechanical properties of the bone tissue. The unique hierarchical structure of bone tissue consists of localized macro- and nanofeatures which prove essential in osteogenesis and functional maturation of the bone tissue. Such architectural complexity has throttled traditional hydrogel- or scaffold-based engineering methods as most of these systems recapitulate bone tissue microenvironment, at best. Moreover, bone morphogenesis is highly dependent on nanoscale biochemical cues, growth factors, and physical factors such as timesensitive biodegradability. However, the nanoscale characteristics of the native bone are often left unaddressed with the lack of controllable nanoscale anisotropy in conventional biomaterials. To tackle this issue, several studies have introduced nanomaterials, namely, carbon-based structures into bioengineered constructs [67
69]. Adding nanoscale features to the bone-mimetic engineered constructs can promote efficiency and precision in stem cell differentiation into specific cell types and increase their long-term regenerative capacity [70]. Unlike common bonerepair polymers such as poly-L-lactic acid and PCL, which require chemical modification to compensate for poor cell adhesion, incorporation of graphene derivatives into the mentioned polymers can provide enhanced cell attachment and growth as well as improved biocompatibility [69,71]. For instance, Wang et al. evaluated the in vitro and in vivo capabilities of 3D-printed PCL scaffolds embedded with graphene sheets in bone regeneration compared to PCL scaffolds alone or with no scaffold treatment (Fig. 21.5). As such, graphene-containing scaffolds showed increased cell proliferation, well-organized collagen deposition, and lower immune response. Graphene-based materials developed for bone regeneration have also shown enhanced osteogenesis capability by providing increased π interactions with local dexamethasone content, thus eliminating the need for bone growth factors such as BMPs [2] (cf. Chapter 14: Bone morphogenetic protein-assisted bone regeneration and applications in biofabrication). Furthermore, in conventional bioactive components, such as bioceramics, poly (glycolic acid), and other commonly used polymers, the inherent brittleness or lack of strength in periodic stress loads leads to their low tolerance to dynamic load bearing and fracture. Scaffolds with high hydrogel or soft polymer content such as chitosan and PEG also lack mechanical strength. Addition of graphene and CNTs to these components has shown to improve the mechanical strength and fracture

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Figure 21.5 Enhancement of bone regeneration using carbon-based nanomaterials. (A) Fluorescence staining [40 ,6-diamidino-2-phenylindole (DAPI, blue), F-actin (red)] of scaffolds on day 14 of culture. (B) Immune response of various scaffolds and controls on day 1 and day 3 after cell seeding. (P , .05) between scaffold and positive control groups. (C) SEM images of the cell seeded PCL/graphene (0.50 wt.%) scaffold. (D) Masson trichrome photomicrography of injury site after 60 or 120 days, area of the bone defect, BE, S, CT, BT, G nanosheets, electrically stimulated scaffolds (ES). Connective tissue (collagen) and primary mineralized bone shown in blue, mineralized tissue after 2 or 4 postoperative months shown in red [71]. BE, Bone edge; BT, bone tissue; CT, connective tissue; ES, embryonic stem; G, grapheme; PCL, polycaprolactone; S, scaffold; SEM, scanning electron microscopy.

toughness of the engineered bone scaffolds [2,68,72]. For instance, Xu et al. showed that adding MWCNTs to chitosan
gelatin
hydroxyapatite composites (CS/Gel/nHAp) significantly enhanced the elastic modulus of the scaffolds 3.4 folds compared to that of CS/Gel/nHAp scaffolds with no MWCNT treatment [73]. High cell adhesion
interconnected microporous network and enhanced mechanical properties of the mentioned scaffolds were shown to promote osteogenic differentiation of osteoblasts and lead to bone regeneration. To conclude, carbon-based biomaterials can be tuned to provide nanoscale surface toughness, high mechanical strength, as well as precisely guided stem cell differentiation through functionalization with various proteins and biomolecules. As such, carbon-based scaffolds with high degrees of toughness and complex nanostructures could be designed to better replicate the hierarchical structure of the bone organ. However, such complex designs necessitate incorporation of important factors such as vasculature into the engineering constructs, as the long-term viability of these structures cannot be ensured merely through perfusion [62]. Thus high porosity and proper vascularization should be studied in combination with osteogenesis studies to confirm adequate nutrient and oxygen supply as well as proper scaffold byproduct and waste removal from the treated area [74].

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Considerations for in vivo tissue regeneration

21.5.1 Toxicity While optimal design and engineering carbon-based scaffolds is the key to ensuring effective tissue regeneration, postimplantation challenges also compose a major step of design considerations and should therefore be thoroughly assessed. Despite having an immense potential to promote tissue regeneration, carbon-based nanomaterials are still questioned for their in vivo toxicity and health concerns. There has been a wide debate on the degree of toxicity of carbon-based nanomaterials and whether they adversely affect the tissue site. While some studies report the use of carbon-based materials as safe, others report acute toxicity to DNA and gene expression, activated reactive oxygen species secretion [75]. Specifically, long-term exposure at high concentrations (50 mg/L) can lead to the mentioned toxicities to in vivo and in vitro models. Toxicity can be pertained to a number of factors such as high aspect ratio, particle size, chemical properties, and functional groups on the surface and aggregation conditions [76]. Due to hydrophobicity, large graphene sheets ( . 100 nm) can aggregate on the cell plasma membrane and hinder ion channel activity and nutrient transport which can eventually interfere with natural intracellular equilibrium [1]. Similarly, upon accumulation, CNTs tend to form fibrous structures, causing toxicities such as granuloma formation [77,78]. Studies suggest that careful functionalization and purification of SWCNTs into individualized and particles smaller than 300 nm can be tolerated by in vivo systems and removed facilely through the renal and biliary systems [79]. Moreover, functionalization of carbon-based nanomaterials with biocompatible chains such as carboxyl groups can help to mitigate toxicity [80
82]. Since tissue regeneration highly relies on the use of stem cells, it is important to consider cytotoxic and genotoxic effects of carbon-based scaffolds on stem cells. While exact mechanism of action and scale of damage of carbon nanomaterials is yet to be clarified, it has been shown that long-term exposure of high concentrations of carbon-based materials can radically affect the stem cells through oxidative stress and damages to cell membrane and mitochondrosome [1,83]. A method to prevent such damages is to functionalize the nanomaterial with particular biocompatible materials. For instance, functionalization of MWCNTs with biomaterials such as PEG and HA nanocomposites has shown favorable biocompatibility and induced relatively lower toxicity in bone MSCs compared to pristine scaffolds or those functionalized by acid oxidation [83]. Finally, to ensure biosafety of designed carbonbased scaffolds in clinical practice, it is imperative to conduct further extensive toxicological studies to assess the toxicity level, mechanism of action, and side effects in both long- and short-term courses before such materials can find their way into practical organ and tissue regeneration.

21.5.2 Biodegradability A critical factor for successful development of carbon-based scaffolds after implantation is long-term biocompatibility. Impermeability of graphene against ions and

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Figure 21.6 Cytotoxicity concerns of carbon-based nanomaterials and efforts to address them. (A) cytocompatibility of rGO in mouse mesencyhmal stem cells at different graphene concentrations in medium. Live (green) and dead (red) cells are shown. (B) H&E staining of subcutaneous tissue. The implanted area in rGO after 3 and 30 days is shown with clear signs of monocyte infiltration [86]. (C) X-ray images of rat tibial bone defects 2 weeks postimplant comparing bone healing in control (no filler), filled with CS/Gn scaffold and 0.25% GO/CS/Gn scaffold models with the latter showing improved wound closure. (D) H&E (top) and Masson’s trichrome (bottom) staining in rat tibial bone sections two weeks postimplant shows increased collagen deposition in the 0.25% GO/CS/Gn scaffolds [87]. rGO, Reduced graphene oxide.

small molecules may induce considerable obstructions in ionic transportations and cell
cell interactions [1,84]. Despite hydrophobicity and poor degradability of pristine graphene, functionalization into derivatives such as rGO can enhance the biocompatibility by being hydrophilic and water soluble; thus degrade via aqueous pathways over the course of few months [85,86]. It has been shown that reducing oxidation degree can further lead to faster immune system uptake and clearance from the treated site (Fig. 21.6A and B) [86]. In addition, carbon-based materials can be incorporated with biodegradable materials such as chitosan to reach improved biodegradability [87,88]. For instance, chitosan/gelatin/GO nanocomposites have exhibited cyto-friendly characteristics, improved wound closure, and bone healing two weeks after treatment of rat tibial bone defect with the designed scaffold (Fig. 21.6C and D) [87].

21.6

Conclusion and future perspectives

Carbon-based nanomaterials, in particular, graphene and CNTs, have been increasingly explored in designing tissue-engineering scaffold owing to their unique

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nanoscale features and biophysical and biochemical properties. Integrating these nanomaterials in biocompatible hydrogels and polymers such as PEG, fibrin, and GelMA, a native ECM-mimetic scaffold with anisotropic nanoscale architecture can be achieved which has been proved highly effective in directing stem cell fate and differentiation into a wide range of tissues such as cardiac, neural, bone, and skeletal muscle. Moreover, as opposed to pristine hydrogels that lack good electrical conductivity, carbon-based materials are endowed with excellent conductivity which has shown to promote stem cell maturation in terms of functionality. This advantage is especially favorable for retaining the natural functionality of electrically excitable tissues such as cardiac and neural tissues which often tend to lose functionality undergoing traumatic events or degenerative damages. As such, several multifunctional hybrid engineered tissues have been recently developed, some of which are endowed with augmented capabilities over the traditional scaffolds, namely, high flexibility, chemically functionalized with growth factors for controlled release, and photo-patternable geometries. Recent applications of carbonbased materials for tissue and organ regeneration were highlighted in this chapter followed by a discussion on in vivo and postimplantation considerations such as toxicity and biodegradability of these nanomaterials. Health and environmental hazards and toxicity of carbon-based materials is still a critical undergoing research and has not been fully overcome with the lack of evidence in long-term toxicity and intertissue effects. However, careful purification and surface functionalization can reduce the toxicity and alleviate the biocompatibility concerns. Finally, while the field of hybrid carbon-based tissues is still in developmental stage and is still majorly limited to laboratory research than clinical practice, we envision that their multifaceted capacity will soon be realized in effective and controlled organ regeneration followed by versatile clinical applications.

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904.

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based hydrogels for tissue engineering

22

N. Vijayakameswara Rao Department of Chemical Engineering, National Taiwan University of Science and Technology, Taipei

22.1

Introduction

Hydrogels are highly biocompatible biomaterials due to their low surface tension, high water content, their hydrodynamic properties that are very similar to those of biological tissues, and soft character that reduces mechanical irritation to surrounding tissues and organs [1
6]. Most of the hydrogels are explored in the several biomedical fields, such as diagnostics, bioreactors, drug delivery matrices, and tissue-engineering scaffolds. Hydrogels have a lot of unique features, such as their resemblance to tissue extracellular matrix (ECM), support for cell growth and migration, controlled release of drugs or growth factors, minimal mechanical irritation to surrounding tissue, and nutrient diffusion, which encourage the viability and proliferation of cells [3,5
7]. Injectable hydrogels are promising tools within the area of tissue engineering (TE), as they can target defects in very deep tissues with minimal invasiveness and better conform to defects. Many natural and synthetic polymers have been analyzed as hydrogel precursors for biomedical applications. One of those natural polymers, hyaluronic acid (HA) is a fantastic candidate for fabrication of biocompatible hydrogels because it is highly biocompatible and biologically active. Many researchers have analyzed the biological action of HA and its derivatives [4,5]. HA and sodium hyaluronate are mostly used to prepare scaffolds for TE because they are highly reproducible and affordable. HA is a naturally occurring glycosaminoglycan (GAG), a polysaccharide of large molecular weight that shows intriguing viscoelastic properties, excellent biocompatibility, and biodegradability [5]. All these properties of HA-derived hydrogels make them ideal biomaterials for TE Fig. 22.1. Injectable hydrogels based on HA are ready using various chemical and physical cross-linking procedures. Several chemical alterations, targeted at improving, modulating, or controlling the therapeutic action of HA, are utilized to develop new products [5]. The unmodified HA solution is rapidly degraded and excreted from the body. But HA in a hydrogel form shows many biological activities, and its elimination from the body is delayed. There are many published studies on the functionalization of HA for fabricating HA-based hydrogels. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00014-3 © 2020 Elsevier Ltd. All rights reserved.

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Selfassembled nanoparticle

Drugconjugate

Nanotubeconjugate

n Hyaluronic acid Hydrogel-based implants

Hydrogel

Hydrogel 3D model

Figure 22.1 Schematic illustration of hyaluronic acid
based materials for nanomedicine and tissue engineering.

This chapter concentrates on cross-linking procedures to prepare HA hydrogels for tissue-engineering applications. First, the chemical structure and modifications of HA are outlined. Second, the main functional groups of HA that are widely used to modify the structure and cross-linking agents are suggested. Based on these factors, we then emphasize further cross-linking techniques for hydrogel preparation, including Schiff-base cross-linking chemistry, thiol-modified cross-linking chemistry, Diels
Alder cross-linking chemistry, and photo-cross-linking chemistry. In the final section, we have discussed the HA-based materials for TE.

22.2

Chemical modifications of hyaluronic acid

22.2.1 Hyaluronic acid In 1934 Karl Meyer and John Palmer described a process for isolating a new GAG in the vitreous humor of bovine eyes. They revealed that this substance comprised an uronic acid and an aminosugar; they suggested the name HA. HA, also called hyaluronan, is a naturally occurring nonsulfate linear polysaccharide. HA consists

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of repeating disaccharide units of D-glucuronic acid and N-acetyl-D-glucosamine linked with β-1-3 and β-1-4 glycosidic bonds [5]. HA is the main GAG component of the ECM of human connective tissues. It is a significant structural component in the skin and has several cell surface receptor interactions; it also has immunosuppressive and antiangiogenic activities. HA is within brain tissue, hyaline cartilage, and synovial joint fluid. Due to the strong hydrophilic character and its high molecular weight in biological cells which could absorb a lot of water up to 1000 times its strong volume, HA displays significant structural and functional roles in the body [5,6]. HA is also used clinically in soft tissue replacement and enhancement, in addition to surgical procedures and diagnostics. However, HA is highly soluble and often displays very poor mechanical properties with rapid degradation behavior in vivo. Thus HA has been chemically and cross-linker-modified to improve its own properties, such as mechanical properties, viscosity, solubility, degradation, and biologic properties. HA derivatives have been used in scaffolds for TE, in soft tissue surgery such as vocal fold augmentation, drug delivery, intracellular delivery of small interfering RNA, wound healing, and as a device in many surgical procedures [5]. The preparation of HA materials has been achieved by using many different chemical modifications to give mechanically robust materials. All these HA derivatives possess physicochemical properties that might be substantially different from the native polymer, but many derivatives have the biocompatibility and biodegradability of indigenous HA. The most common method of modification of HA is cross-linking to form a hydrogel. But, the durability of HA hydrogels is based upon their ability to resist degradation from hyaluronidases and reactive oxygen and nitrogen species, thereby limiting their efficient usage. To overcome the problems of HA degradation, functional groups on HA are manipulated in the preparation of HA-based materials. In the chemical structure of HA, the three most often used sites of covalent modifications are carboxylic group, a hydroxyl group, and
NHCOCH3 group as shown in Fig. 22.2. The chemical modification of HA can be performed on the carboxylic acid group and the hydroxyl group functional sites (Fig. 22.1) [4,5,8]. The amino group can also be recovered by deacetylation of the N-acetyl group. Carboxylic group reactions include amidation and esterification, while the hydroxyl groups give origin to ester and ether linkages. HA can bear reductive amination in the one aldehyde group at the reducing end of the polymer. Through deacetylation of the N-acetyl group, it is possible to recover an amine and by periodate oxidation dialdehyde groups via ring opening of the D-glucoronic acid residue [1]. In the literature, there are many reagents that couple carboxyl and amino groups to form amide or ester bonds. But for preparing cross-linked hydrogels, the most commonly used reagents are carbodiimides, carbonyldiimidazole, and others [8]. The formation of the new bond will proceed in two process steps. In the first step the reagent forms a reactive adduct with the carboxyl group. In the second step, nucleophilic attack at the activated species eliminates the activating moiety, resulting in the formation of a new bond. The amino group or alcohol group is implied as the nucleophile in these reactions [4,5,9,10]. The chemical modification of
OH

Figure 22.2 Chemical modifications of hyaluronic acid.

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Figure 22.3 Schematic illustration of injectable hydrogels with gelatin and hyaluronic acid cross-linked by Schiff’s base formation. (A) Synthesis of gelatin-modified hydrazide. (B) Synthesis of aldehyde-modified HA. Source: Reprinted from Hozumi T, Kageyama T, Ohta S, Fukuda J, Ito T. Injectable hydrogel with slow degradability composed of gelatin and hyaluronic acid crosslinked by Schiff’s base formation. Biomacromolecules 2018;19:288
97 with the permission of ACS Publications.

groups can be prepared by ether formation, ester formation, hemiacetal formation, and oxidation. Reagents used in these reactions are summarized in Fig. 22.3. The modification reactions of the
NHCOCH3 group comprise deacetylation, amidation, hemiacetylation, and hemiacetal formation, among others. Amidation methods are utilized for deacetylation of the N-acetyl groups of HA that may then react with acid and are generally conducted using hydrazine sulfate [4].

22.3

Cross-linking chemistry of hyaluronic acid

The cross-linking of the polymer is a method to increase mechanical stability [11,12]. For clinical uses of HA, there is a requirement to increase the stability of HA solutions after injection. Several research teams have investigated the cross-linking of HA-amine or HA-hydrazide derivatives obtained with homo- or hetero-functional cross-linking agents, including bis(sulfosuccinimidyl) suberate, 3,3-dithiobis(sulfosuccinimidyl) propionate, or 2-methyl suberimidate. They react with HA-hydrazide derivatives at pH values of 5.0, or HA-amine derivatives at pH values above 8.0 because of the greater pKa of the amino groups [11]. The in situ

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generations of possibly toxic degradation hydrazide derivatives may lead to unwanted effects, which will need to be evaluated with in vivo procedures. Many chemical cross-linking methods, including Michael addition, click chemistry, and Schiff’s base formation, were studied [8,11,12]. One of these, Schiff’s base formation was widely researched, due to rapid cross-linking and superb biocompatibility. Despite numerous previous studies, management of the degradation rate has turned into a challenge in Schiff’s base cross-linked hydrogels. Recent studies have proven that the hydrolytic stability of Schiff’s base is enhanced by reducing the electrophilicity of carbonyl carbon [11]. A growth of nucleophilicity of the amine derivatives by electron donation to C 5 N and π
π conjugation contributed to the hydrolytic stability. Hydrazone or oxime cross-linked hydrogels with a benzoic aldehyde as carbonyls, whereas hydroxylamines and hydrazides were included as counterparts and were stable to hydrolysis at physiological pH. To control the degradation rate of injectable hydrogels, the effect of this cross-linking structure demands additional study.

22.3.1 Schiff-base cross-linking hydrogels Schiff-base reactions are among the most widely accepted approaches for the preparation of hydrogels, especially due to their mild reaction conditions. Schiff bases are prepared by facile condensation of an aldehyde or a ketone with primary amines. The formula for Schiff bases is RN 5 CR0 Rv, where R, R0 , and Rv may be alkyl, aryl, heteroaryl, or cycloalkyl. The
C 5 N
imine bond from Schiff bases plays a particular role in conferring broad-spectrum biologic activities to these compounds [13]. Hozumi et al. developed an injectable HA hydrogel with gelatin which consisted of novel injectable hydrogel composed of carbohydrazide-modified gelatin (Gel-CDH) and monoaldehyde-modified HA (HA-mCHO). Upon mixing the two precursor polymers, the hydrogel was fabricated through Schiff’s base formation. The ex vivo rat aortic-ring assay also demonstrated the potential of the GelCDH/HA-mCHO hydrogel for tissue-engineering scaffolds (Fig. 22.3). The in situ forming amine-modified HA (HA-NH2) and CHO-HA hydrogels were synthesized via Schiff-base reaction. HA-NH2 was synthesized by coupling ethylenediamine with 1-ethyl-3-(3-dimethyl aminopropyl) carbodiimide and hydroxybenzotriazole at pH 6.8 for 24 hours. Genipin was used for double cross-linking, resulting in a more compact microstructure, slower mass loss, and higher compressive modulus of hydrogels [13]. A new hydrogel hemostat consists of HA-conjugated with inorganic polyphosphate (PolyP). A hemostatic hydrogel, HAX-PolyP, was formed quickly by mixing aldehyde-modified HA and hydrazide-modified HA-conjugated with PolyP. HAX-PolyP revealed similar biocompatibility using the HA hydrogel without PolyP conjugation in vitro and in vivo. The HA-PolyP accelerated the coagulation rate of human plasma ex vivo, and HAX-PolyP showed as powerful that the hemostatic effect as fibrin adhesive in a mouse liver bleeding version in vivo (Fig. 22.4) [14]. The injectable HA hydrogel was prepared for in vivo bone augmentation by Martı´nez-Sanz et al. They prepared HA modified with 3-amino-1,2propanediol and subsequently reacted with NaIO4 to provide aldehyde functional

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Figure 22.4 A synthetic scheme of the injectable hemostat hydrogel by using hyaluronic acid and polyphosphate. (A) The HAX-PolyP illustration as an injectable hemostat. (B) Synthesis of the (i) HA-CHO, (ii) HA-ADH, and (iii) HA-PolyP polymers. Source: Reprinted from Sakoda M, Kaneko M, Ohta S, Qi P, Ichimura S, Yatomi Y, et al. Injectable hemostat composed of a polyphosphate-conjugated hyaluronan hydrogel. Biomacromolecules 2018;19:3280
90 with the permission of ACS Publications.

groups. Mixing equal volumes of HA-aldehyde derivative and HA-hydrazide derivative formed a hydrazine-cross-linked hydrogel within 30 seconds [15]. Naturally derived chitosan (CS), polysaccharide, was composited with HA to decrease the erosion and degradation behaviors of hydrogels. Insulin has been entrapped inside an N-succinyl-CS and aldehyde HA hydrogel leading to functional adipose tissue. In another interesting study, injectable OHA-gelation-adipic acid dihydrazide (oxiHAG-ADH) in situ forming hydrogel was developed by using high molecular weight HA and gelatin for pulposus regeneration [16]. The results showed that oxiHAG-ADH is biocompatible and has highly viscoelastic properties. The CS
HA hydrogel and investigated its effects on abdominal tissue regeneration [17]. The physical and biological properties of the hydrogel were demonstrated to be suitable for application in abdominal wounds. The newly designed hydrogel

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exhibited a rapid cellular response, sufficient ECM deposition, and marked neovascularization were found after the application of the hydrogel, compared to the control group and the fibrin gel group (Fig. 22.5) [17]. Injectable HA-based hydrogels were widely used for cartilage TE. In the literature arginyl-glycyl-aspartic acid (RGD) peptide-functionalized pectin [18], type I collagen [19], thiolated HA derivative [20], and the glycol CS [21] are suitable compounds to combine with HA for cartilage TE. Glycol CS combined with oxidized HA through Schiff-base formation has been used as injectable hydrogel for cartilage TE. ATDC5 chondrocytes were seeded inside the hydrogels and results indicated good biocompatibility and proliferation. Thus oxidized HA/glycol CS injectable hydrogels can be used as cell-delivery vehicles [21]. The HA-based in situ forming hydrogels have been prepared to promote angiogenesis. The development of a slow degradable HA/gelatin hydrogel is suitable for inducing the formation of new blood vessels. HA and gelatin were functionalized with a monoaldehyde and carbohydrazide, respectively. To fabricate the hydrogels a double-barreled syringe fibrin glue applicator was used to generate a Schiff’s base formation. Hydrogel biocompatibility was demonstrated via in vitro cell viability tests. The ex vivo assay using a dissected aortic rat ring revealed that the synthesized hydrogel is a suitable support for the microvascular extension [22]. In another report a hybrid injectable hydrogel, consisting of deferoxamine-loaded poly(lacticco-glycolic acid) nanoparticles incorporated into an HA/CS hydrogel. The angiogenesis was induced by deferoxamine drug release. The HA/CS hydrogel showed cytocompatibility and was inducive to cell proliferation, and maximal blood vessels formation through in vivo tests. Subcutaneous injection of the hydrogel into mice proved the beneficial effect of deferoxamine for neovascularization when compared to HA/CS hydrogel [23]. The HA/CS in situ forming hydrogel has been used for abdominal tissue regeneration. The physical and biological properties of the hydrogel were suitable for application in abdominal wounds. The in vivo test indicated that the use of HA/CS hydrogel led to faster tissue regeneration increased cellular accumulation and ECM deposition. Besides the increased amount of fibroblasts and endothelial cells, the regenerated tissue in the presence of the composite hydrogel showed a greater thickness and capillarity when compared to the naturally regenerated tissue. These findings concluded that CS/HA hydrogel was promising for abdominal tissue regeneration. The injectable composite hydrogel consisting of HA and methylcellulose was prepared using polyethylene glycol as a cross-linker to study the central nervous system TE. Hydrogel preparation involved the dissolving of the compounds into the artificial cerebrospinal fluid, followed by cross-linking with polyethylene glycol (PEG). The in vitro tests displayed better cytocompatibility for lower concentrations of hydrogels. Cell viability results showed that HAMC was cytocompatible for further applications in vivo and would be a promising choice for neural TE in the future. The HA/RGD-functionalized injectable hydrogel was prepared by using cross-linked HA and pectin dialdehyde with adipic acid dihydrazide and G4RGDS oligopeptide by carbodiimide chemistry, respectively. When chondrocytes were encapsulated into hydrogels, the in vitro results showed

Figure 22.5 (A) Schematic illustration of the preparation of CS/HA hydrogel via Schiff’s base reaction. Chemical structures of chitosan (CS) (a) carboxymethyl chitosan (NOCC) (c), hyaluronic acid (HA) (b), and aldehyde hyaluronic acid (A-HA) (d). (B) Schematic diagram of the animal model and the experimental procedure: (a) the abdominal wall defect model, (b) the Control group, (c) fibrin gel or CS/HA hydrogel employed over the defect, (d) the Fibrin gel or CS/HA hydrogel group. (C) Gross observation: (a,d) the established defect wound, (b) the Control group (black arrow: PP mesh), (c,e) the employed fibrin gel or CS/HA hydrogel (green arrow: fibrin gel or CS/HA hydrogel), (f) the Fibrin gel or CS/HA hydrogel group. Reprinted from Deng Y, Ren J, Chen G, Li G, Wu X, Wang G, et al. Injectable in situ cross-linking chitosan-hyaluronic acid based hydrogels for abdominal tissue regeneration. Sci Rep 2017;7:2699. with the permission of ACS publications.

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improved phenotype maintenance and chondrogenesis with increased RGDfunctionalized pectin, as cell membrane integrins recognize RGD oligopeptide. Also, the biomimetic hydrogel exhibited acceptable tissue compatibility when delivered to a mouse model. Though Schiff-base reaction hydrogels are simple to translate to clinical applications due to their simple methodology, one of the constraints of utilizing such injectable hydrogels is that their pH sensitivity. Schiff-base (imine) linkages are most likely to hydrolyze under acidic conditions [24]; consequently, these hydrogels cannot be utilized for biomedical applications, in which hydrogel stability is crucial in disease conditions, where the pH is generally slightly acidic [25]. Current crosslinking chemistries frequently require a coupling agent, catalyst, or photoinitiator, which may be cytotoxic, or require a multistep synthesis of functionalized-HA, increasing the complexity of the system. With the goal of designing a more straightforward one-step, aqueous-based cross-linking system, we synthesized HA hydrogels through Diels
Alder “click” chemistry [26].

22.3.2 Diels
Alder click cross-linked hydrogel The current HA cross-linking chemistries often require a coupling agent, catalyst, or photoinitiator, which may be cytotoxic, or involve a multistep synthesis of functionalized HA, increasing the complexity of the system. With the goal of designing a simpler one-step reaction, the Diels
Alder (DA) cross-linking of HA is designed. The DA reaction is a [4 1 2] cycloaddition involves the covalent coupling of a conjugated diene with a substituted alkene to form a six-membered ring product in the absence of a catalyst. The electron-withdrawing groups on the alkene and the electron donating groups on the diene are important factors for increasing reaction rate. Furan-modified HA derivatives were synthesized and cross-linked via dimaleimide poly(ethylene glycol) [27]. By controlling the furan-to-maleimide molar ratio, both the degradation and mechanical properties of the consequent DA cross-linked hydrogels can be tuned. Rheological and degradation studies demonstrate that the DA click response is a suitable cross-linking way for HA. All these HA crosslinked hydrogels have been shown to be cytocompatible and might represent a promising material for soft TE [27]. HA/PEG hydrogel was fabricated for the first time by incorporating two cross-linking procedures, such as first enzymatic crosslinking and following DA click chemistry. The enzymatic cross-linking led to speedy gelation of HA/PEG in 5 minutes, causing the creation of an injectable substance. In addition, the DA click response cross-linking created a hydrogel that has outstanding contour memory and antifatigue attributes [28]. New HA-based hydrogels have been obtained by means of a “click chemistry reaction,” between an azide derivative and an alkyne derivative of HA obtained by amidation in an aqueous solution of HA using an amino-azide bifunctional linker and propargylamine, respectively. The HA-based click-gels could be useful materials for controlled drug release of therapeutically relevant biomolecules as well as for cells scaffolding in TE [26].

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22.3.3 Photo-cross-linking A variety of alterations of native hyaluronan have been invented to provide mechanically and chemically robust substances through compound cross-linking. As an example, modification of HA can be achieved by covalent derivatization of the carboxylic acid or hydroxyl functionalities of the polymer. The resulting hyaluronan derivatives possess physicochemical properties that are substantially different from a native polymer. Nevertheless, most derivatives are both biocompatible and biodegradable. Reducing the degradation levels of HA-based substances is necessary for several biomedical applications. Bencherif et al. [29] and Baier Leach et al. [30] have synthesized the HA hydrogels by the photoinduced crosslinking approach. The methacrylated-HA obtained through glycidyl methacrylate or via methacrylate anhydride was cross-linked by free-radical photoinduced polymerization procedure when subjected to under UV-light (365 nm). In this case a photoinitiator, for example, 2-oxo-ketoglutaric acid or 4-(2-hydroxyethoxy) phenyl-(2-hydroxy-2-propyl) ketone is essential to initiate radicals. Fenn and Oldinski have developed visible light cross-linking of HA-methacrylate hydrogels for injectable tissue repair [31].

22.3.4 The hyaluronic acid
disulfide cross-linking hydrogels A disulfide cross-linking chemistry was used to prepare HA hydrogel from thiolmodified HA [12]. The dithiobis(propanoic dihydrazide) and dithiobis(butyric dihydrazide) were synthesized and then conjugated to HA with carbodiimide chemistry. Hydrogels were subsequently formed under moderate conditions by air oxidation of thiols to disulfides. This type of reaction is interesting as it does not rely on synthetic cross-linking agents and will be formed in physiological conditions. Thiolated-HA can be cross-linked via disulfide bond formation upon oxidation. Nonetheless, the synthesis of thiolated-HA is a time-consuming, multistep process that can negatively impact the native structure of HA by decreasing the molecular weight [12]. Moreover, thiol-disulfide is crucial cellular functions such as adhesion and proliferation, and consequently a sterile hydrogel canvas without thiols allows more control over cellular behavior.

22.4

Hyaluronic acid as a biomaterial in tissue engineering

22.4.1 Hyaluronic acid
based scaffolds Hydrogels have drawn fantastic attention on account of their applicability in an assortment of fields, such as TE [3]. HA hydrogels can mimic the individual tissue in terms of water content and also exchange oxygen, nutrients, and metabolic waste. The usage of HA-based hydrogels has been intensively studied for cartilage TE, as it is an integral component distributed ubiquitously through the cartilaginous ECM

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and they are able to keep the morphology of pure chondrocytes. Chondrocytes abundantly secrete type II collagen and GAGs in the ECM of the native cartilage tissue. The demand for the development of the right biomaterial relies upon the limited capacity for repair of the articular cartilage, likely because of its avascularity and composition. Hydrogels acquired from pure HA are utilized for individual umbilical mesenchymal stem cells implantation into the articular site. Studies on rabbit models for cartilage tissue regeneration revealed a better cartilage repair effect for seeded hydrogels. The repaired tissue had similar cellular architecture and type II collagen arrangements to the natural tissue [32]. Using CS, an elastin-like protein and tauroursodeoxycholic acid-poly(lactic-co-glycolic acid) microspheres [33] as substances for HA-based hydrogels synthesis also have been analyzed. Hydrogels prepared from oxidized HA and CS were used as scaffolds for chondrocytes encapsulation. The simulated in vitro microenvironment showed good cell viability and also an appropriate ECM production, thus demonstrating the potential of the system for further in vivo studies. Implantation of HA
CS hydrogels in osteochondral defects in rabbit knee joints demonstrated the possibility of the system for tissue repair. Histology tests affirmed tissue regeneration from the lack of hydrogel traces from the defect site [34]. The combination of HA and also an elastin-like protein revealed positive effects in stimulating the regeneration of cartilage tissue. After the in vitro testing, it had been observed that using the hydrogel helped to preserve cell phenotype while enhancing cell proliferation and ECM generation [35]. Tauroursodeoxycholic acid is employed for its properties to cause neovasculogenesis and the inhibition of adipogenic differentiation of mesenchymal stem cells. Implantation at osteochondral defects in rats revealed the filling of the defect and also the formation of new tissue following 10 weeks. The usage of tauroursodeoxycholic acid enhanced tissue regeneration analyzed by raised proteoglycans from the superficial region and calcified cartilage. Poly(Nε-acryloyl-L-lysine) has been assessed for bone TE, with an HA-based hydrogel. The usage of this copolymer is anticipated to demonstrate cytocompatibility because both its components promote cell attachment and proliferation. The in vitro tests conducted on osteoblast precursor cells from mouse revealed cell viability and enhanced cell proliferation, as anticipated. In vivo biocompatibility was assessed after 7 days utilizing the dorsal skin-fold room technique, and the results revealed that the inclusion of the copolymer into the HA hydrogel enhanced biocompatibility. Furthermore, histological tests demonstrated a much better cell infiltration with the greater crosslinking of this hydrogel, thus demonstrating the connection between the mechanical properties of compact structures and enhanced native mobile penetration [36]. An aqueous solution comprising phenolic-substituted HA, horseradish peroxidase, and catalase was utilized to manufacture hollow hydrogel fibers via microfluidic rotation technique. These hydrogel fibers supply a biomimetic microenvironment, permitting cell encapsulation. Cell viability tests demonstrated their potential for engineering complex artificial tissues with appropriate biocompatibility and biodegradability. HA coupled with methylcellulose demonstrated that it could be applied to 3D bioprinting, as it is a versatile material that could be applied as a bioink. Encapsulated mesenchymal stem cells survived the process of bioprinting of the

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hydrogels, staying viable for one more week. These results reveal that HA
methylcellulose combinations are promising materials for 3D bioprinting of artificial tissues [37].

22.5

Conclusion

HA is one of the most appropriate natural materials in hydrogel scaffolds. HA has a similar water content as human tissue, has good tissue compatibility, and it plays an important role in promoting the proliferation and differentiation of seed cells. Natural HA has a short reservation time in vivo due to the rapid degradation by hyaluronidase. Therefore appropriate cross-linking and modification of the active group is essential. HA after cross-linking and modification has been widely used alone or compounded with other materials in the TE. The current problem is that there are a variety of cross-linking and modification methods of HA. Furthermore, the raw materials from different origins owning divergent molecular weight, thus exhibiting different biological properties, which brings about a lot of issues into the research and applications of HA. HA and HA-based substances are extensively employed for the preparation of scaffolds for TE. These scaffolds offer the benefits of increased biocompatibility, controllable degradation rate by using a cross-linker, and appropriate porosity for cell encapsulation, differentiation, and proliferation. HA-based scaffolds are used for a variety of tissues, where the cartilage tissue being the most intensively studied.

Acknowledgement My sincere thanks to National Taiwan University of Science and Technology (NTUST) for providing initial startup fund.

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11. [13] Tan H, Li H, Rubin JP, Marra KG. Controlled gelation and degradation rates of injectable hyaluronic acid-based hydrogels through a double crosslinking strategy. J Tissue Eng Regener Med 2011;5:790
7. [14] Sakoda M, Kaneko M, Ohta S, Qi P, Ichimura S, Yatomi Y, et al. Injectable hemostat composed of a polyphosphate-conjugated hyaluronan hydrogel. Biomacromolecules 2018;19:3280
90. [15] Martinez-Sanz E, Ossipov DA, Hilborn J, Larsson S, Jonsson KB, Varghese OP. Bone reservoir: injectable hyaluronic acid hydrogel for minimal invasive bone augmentation. J Control Release 2011;152:232
40. [16] Chen YC, Su WY, Yang SH, Gefen A, Lin FH. In situ forming hydrogels composed of oxidized high molecular weight hyaluronic acid and gelatin for nucleus pulposus regeneration. Acta Biomater 2013;9:5181
93. [17] Deng Y, Ren J, Chen G, Li G, Wu X, Wang G, et al. Injectable in situ cross-linking chitosan-hyaluronic acid based hydrogels for abdominal tissue regeneration. Sci Rep 2017;7:2699. [18] Chen F, Ni Y, Liu B, Zhou T, Yu C, Su Y, et al. Self-crosslinking and injectable hyaluronic acid/RGD-functionalized pectin hydrogel for cartilage tissue engineering. Carbohydr Polym 2017;166:31
44. [19] Chen Y, Sui J, Wang Q, Yin Y, Liu J, Wang Q, et al. Injectable self-crosslinking HASH/Col I blend hydrogels for in vitro construction of engineered cartilage. Carbohydr Polym 2018;190:57
66. [20] Shaoquan B, He M, Junhui S, Cai H, Sun Y, Liang J, et al. The self-crosslinking smart hyaluronic acid hydrogels as injectable three-dimensional scaffolds for cells culture. Colloids Surf B Biointerfaces 2016;140:392
402. [21] Kim DY, Park H, Kim SW, Lee JW, Lee KY. Injectable hydrogels prepared from partially oxidized hyaluronate and glycol chitosan for chondrocyte encapsulation. Carbohydr Polym 2017;157:1281
7. [22] Hozumi T, Kageyama T, Ohta S, Fukuda J, Ito T. Injectable hydrogel with slow degradability composed of gelatin and hyaluronic acid cross-linked by Schiff’s base formation. Biomacromolecules 2018;19:288
97. [23] S V, A S, Annapoorna M, R J, Subramania I, Shantikumar VN, et al. Injectable deferoxamine nanoparticles loaded chitosan-hyaluronic acid coacervate hydrogel for therapeutic angiogenesis. Colloids Surf B Biointerfaces 2018;161:129
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[24] Billman JH, Diesing AC. Reduction of Schiff bases with sodium borohydride. J Org Chem 1957;22:1068
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Singhal, Reena. Polym Technol Eng 2016;55:54
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Alder click cross-linked hyaluronic acid hydrogels for tissue engineering. Biomacromolecules 2011;12:824
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Alder “click chemistry”. Polym Chem 2014;5:1082
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49. [30] Baier Leach J, Bivens KA, Patrick Jr. CW, Schmidt CE. Photocrosslinked hyaluronic acid hydrogels: natural, biodegradable tissue engineering scaffolds. Biotechnol Bioeng 2003;82:578
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36. [32] Park YB, Ha CW, Kim JA, Han WJ, Rhim JH, Lee HJ, et al. Single-stage cell-based cartilage repair in a rabbit model: cell tracking and in vivo chondrogenesis of human umbilical cord blood-derived mesenchymal stem cells and hyaluronic acid hydrogel composite. Osteoarthritis Cartilage 2017;25:570
80. [33] Kim BJ, Arai Y, Choi B, Park SH, Ahn JS, Han IB, et al. Restoration of articular osteochondral defects in rat by a bi-layered hyaluronic acid hydrogel plug with TUDCAPLGA microsphere. J Ind Eng Chem 2018;61:295
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Microfluidics in tissue engineering

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Sudip Kumar Sinha1 and Arindam Bit 2 1 Department of Metallurgical Engineering, National Institute of Technology, Raipur, India, 2 Department of Biomedical Engineering, National Institute of Technology, Raipur, India

23.1

Introduction

Tissue engineering is a field of engineering that uses integrated form of different types of cells with an objective to build a functional biological constructs for preclinical and clinical applications. Most of the tissue engineering constructs are being developed at in vitro culture conditions with limited access to micronreconstruction of native biological pattern. Recent studies have shown that perfusion is a vital parameter for effective cellular interaction with the surface and sublayers of scaffold [1]. However, perfused scaffolds face challenges, including continuous flow of media, large quantity of media, and nonuniform distribution of stress for the adhered cells over surfaces of scaffold [2]. Therefore micro-transport phenomena at the site of scaffold
cell interface are an important physical event which needs to have comprehensive understanding. Boundary layer phenomena of fluid transported through microchannels have shown potential impact on cellular morphology during its interaction with substrates and scaffolds [3]. Fluid transported through microchannels occurs at very low Reynolds number (typically less than 0.1 Re). Transported fluid contains colloids (e.g., red blood cells/platelets, and cell medium) with mean diameter of these constituents comparable to the diameter of the channel. It results in the formation of greater drag forces at the periphery of these constituents. A relative upstream of fluidic layers near the boundaries thus enhances the thickness of boundary layers near the wall of the channel. It increases the relative residence time of the fluid at wall. This stationary layer of fluid is also known as lubrication layer [4]. This is the typical regime of microfluidic channel that facilitates the mimicking of capillaries within the context of tissue engineering. Simultaneously, the lubrication layer in microchannels also provides regulated transition of the medium over the embedded scaffold, resulting in optimum interaction of cell-culture media for developing tissue-engineered constructs.

23.2

Design considerations of microfluidics chips

Connection of microchannels on a substrate is a cumbersome task and needs to be appreciated at each stage. Manual connections of channel heads are not effectively Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00010-6 © 2020 Elsevier Ltd. All rights reserved.

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successful. Multiport connectors can be employed to enhance effectiveness in design. Dimension of chip and position of ports are primary markings for introducing multiport connectors. Since industrial application of microfluidics is at nascent stage, universal dimensions are challenges for implementation by developers. Therefore for activating interoperability of any microfluidic design, dimension of chip and connectors need to be standardized. In such case, chip is always considered to be a flat device that can be made of different materials and will host fluidic inlets and outlets at different locations. The connectors possess two sides: one connected to inlet or outlet port, and other either with a delivery tube, or another fluid circuit board, or a sensor. These connectors should also possess flexibility to use other types of components such as adhesives, O-rings with a clamping system, or mini-Luer. Thus it enables interoperability between different types of connectors and chips. A proper nomenclature of chip-geometry enables us to visualize the pathways between chip to fluidic circuit board, the connectors tubing circuits, as well as the development of a transparent concept of the device between the client and manufacturer. Fig. 23.1 represents a schematic diagram with all the components of the microfluidic chip. O-ring facilitates leakage proof to the microfluidic circuit when tubings are connected to the inlet and outlet nozzles of the chip. Mini-Luer is an alternative to O-ring. It is used in specific cases when a mismatch is found between channel height and thickness of the O-ring. Adhesive used at the inlet and outlet ports provide secondary confirmation of the completeness of the fluidic circuit. It also helps the ports to affirm themselves at specific regions on the channel.

23.2.1 Photolithography The process of “Lithography” in general involves transferring a replica of a base or master pattern onto a solid substrate which is typically an Si-wafer.

Figure 23.1 A schematic representation of different components of a microfluidic platform.

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The lexicon meaning of the word “lithography” arises from Greek word stone [lithos] and to write [gra´phein]. The process was first demonstrated by Aloys Senefelder in 1796, and in a crude method, he inked Bavarian limestone and relocated an imprinted image from stone onto paper. Photolithography is the most extensively used lithographic technique so far and has been used as a major method for the fabrication of microfluidic devices. Apart from micropatterning of microfluidic-based device applications, this highly developed technology has been successfully used in the integrated circuits (ICs) industry for microfabrication [5,6]. In the latter case the pattern transfer from masks onto thin films is exclusively done by photolithography. The usual steps involve in manufacturing of ICs and in miniaturization science are photomasking and subsequent chemical processing. Photolithography has been used as a major method for the fabrication of microfluidic devices. It involves exposing a substrate coated with photoresist to light such that the selectively developed regions transmit incident light, thereby etching the surface of the substrate. The steps involve in photolithography process commence with the construction of a mask [normally a “chromium (Cr)” pattern layer on a glass shield], and casting the silicon, glass, or gallium arsenide (GaAs)
based substrate with an active photoresist polymer substance. Next, a high-power, collimated ultraviolet (UV) light is focused via the photomask onto the photoresist layer and it is subsequently cured, transporting the mask pattern to the photoresist (SU-8 or PMMA) on top of the substrate [7,8]. Recently Ma et al. [8] have used sunlight as the UV source to fabricate a simple, low cost, easy to operate, and environmental friendly photolithographic device for microfluidic applications for electrophoretic separation, multiple gradient generator, and pneumatic valve
based cell culture. The photoresist layer is classified in two major categories: positive and negative. In a positive resist the UVray exposed zones are eliminated in the successive growth stage, whereas in a negative type, the active areas remain unaltered after completion of the entire process [5]. Photolithography has proven to be a highly successful technique with numerous advantages. Nevertheless, this practice has some drawbacks once it is applied in microfluidic devices in biomedical engineering. First of all, the process is extremely expensive since the instruments required for this method is associated with the fabrication of state-of-the-art microelectronic devices. Thus clean rooms are an integral part of any photolithography technique, and the huge amount of capital expenditure involve in its construction makes this process practically unavailable in various fields of chemistry, electronics, biochemists, and biomedical engineering. As an alternative, soft lithography technique has become extremely popular in the recent past since its production operation is feasible outside a clean room with the help of a printing or molding process. However, the master pattern has to be made by the use of conventional photolithography technique. Once the master is available, most of the fabrication tasks can be continued outside a clean room with the use of only a printing or molding procedure. Therefore this method is mostly applicable for a limited number of materials.

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Second, the method is mostly performed by the projection of a pattern on a photomask on top of a photoresist film. The photomasks are expensive, and the time required for designing and fabricating them with user-defined architecture causes a significant obstacle for their use in fast, efficient, and commercially viable patterns and devices. Finally, this technique has a very little (or none in some case) ability to control the chemistry of the top layer to be masked and, also, curved or nonplanar substrates cannot be used to pattern through this method of fabrication. Optical or photolithography technique traditionally involves several steps in a very controlled manner. The technique is flexible enough since the desired cells can be grown on top of the electrodes, printed circuit board, and some other micro/nanoelectronic devices. The reason behind this is the typically used substrate materials of micro/nanoelectronic and micro
electro-mechanical system (MEMS) components are based on SiO2, Cu, Pt, Au, glass, quartz, polyimide, etc., and these electronic devices are not only biocompatible but also conducive toward specific applications [9]. Jain et al. have produced miniature sized structures on Si-wafer substrates as a prototype to understand the behavior of cardiac failure of various natures [10]. Other groups have also worked on cardiovascular devices fabricated by photolithography technique. Yamaguchi et al. [11] have selective transformation of patterned APTES
COOH surfaces by means of vacuum-aided photolithography technique. The surface in a specific area/region is chemically modified for patterned culture of rat cardiac cells. The various drawbacks found in traditional photolithography technique for biological and biomedical applications have been successfully overcome by a simple and inexpensive process, often known as “soft lithography technology.” This technique permits the control of the surface molecular structure, micropatterning of multifaceted molecules, and the production of microfluidics channel structures. Fig. 23.2 shows the schematic representation of photolithography process.

23.2.2 Microcontact printing In comparison to photolithography, soft lithography techniques offer a fundamentally different route to fabricate lab-on-a-chip devices on plane or curved flexible and soft substrates in a cost-effective manner. Among the various types of soft lithography techniques in practice [e.g., replica molding (REM), nanoimprint lithography, microtransfer molding, nanotransfer printing, capillary force lithography, solvent-assisted micromolding, decal transfer lithography, nanoskiving, phaseshifting edge lithography, and nanotransfer printing] which are principally based on the operation of printing, molding, and embossing by the help of an elastomeric impression or stamps, micropattern printing is traditionally the foremost runner. All these techniques offer an accurate control of cell behaviors and, therefore, have been extremely effective in developing tools for cell biology and regenerative medicine. Microcontact printing (μCP) is a well-established and one of the major soft lithography-based micropatterning techniques, which has the ability to create

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Figure 23.2 Schematic demonstration of the photolithography technique.

cellular patterns on both plane and curved surfaces with a chemical moiety by either influencing biological causes or by means of topographical cues. This technique was first introduced by the Whiteside group [12]. The technique is based on using a stamp, typically made up of poly(dimethylsiloxane) (PDMS), with patterned relief constructions on its surface. The size of the features embedded in the patterns that is to be fabricated could range down to 40 nm line grating [13] and even 2 nm using nanotubes to mold the stamp [14]. PDMS is a soft transparent material with superior oxygen permeability and excellent biocompatibility. In microfludic chips, the material has the ability to create appropriate microenvironment for cell population, growth, and timely inspection due to its transparency. Although this is a versatile technique for patterning large coverage areas with nanoscale-like features, there are certain drawbacks that limit its application in some areas. A proper alignment and overlaying of several patterning steps in a programmed and precision manner on a solitary substrate is a challenging task while maintaining flexibility in substrate. Another common problem arises from the difficulty to control the printing process, owing to the high compressibility of highly sensitive PDMS. The super sensitive nature of PDMS layer leads to recurrent unmanageable disintegration of the stamps for higher aspect ratios, thereby deteriorating the printing precision and reproducibility. As already mentioned, PDMS-based elastomeric impression or stamps are used for required chemical modification of a plane. The stamp absorbs self-assembled

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monolayer or molecular ink to patterned gold surfaces forming covalent linkage with a protein molecule that steers a cell-pattern formation. In a different study, Choi et al. have shown bacterial array has been grown on target cell adhesion, by virtue of self-assembling of polyelectrolyte multilayers and micromolded poly(ethylene glycol)
poly(lactide)
based block copolymers to support target cell adhesion [15]. Scaffolds and gels are increasingly being popular for tissue engineering applications and to study cells and tissue mechanobiology. Vedula et al. [16] have performed some innovative studies in this perspective and their analysis shows that cells separated by tracks can form epithelial bridges that can sustain tissue integrity throughout cell migration. This scheme has been used to develop fibronectin layer on top of a nonadherent polymer track by μCP. In the present era, automation [17] and robotics [18,19] have revolutionized the integrity and stability of this process. These latest technology-based systems can support prints with sub-10 μm accuracy repetitively while printing diverse molecules as required. The overall benefit arises due to the fabrication of 3D complex architecture of biomechanical devices with different specific cells and subsequently understanding cell signaling, proliferation or cell migration. The efficiency of the printing process is totally controlled by the bridging among the ink and the surface.

23.2.3 Micropatterning of cells on microchannels Micropatterning of mammalian cells are important in tissue engineering studies in order to perform biomimetic in vitro assays and thus evaluating biochemical processes by incorporating bioMEMS sensors. These sensors will have impacts on neurological, oncological, or tissue engineering issues in the case of personalized medicine. Cell patterning is a crucial step in tissue engineering studies in vitro. It brings a larger myriad of materials and techniques interacting, which allow one to immobilize and manipulate cells, imitating the 3D in vivo milieu. Micropatterning deals with enhanced control, a deeper understanding, and implementation of these techniques in bioMEMS systems. On this note, nanotechnologybased novel methodologies have been adopted to fix cells on substrates in highly precise and controlled manner. Cellular micropatterning following manipulation is considered as a primary step to perform drug-testing experiments [20], for understanding biochemical processes [21], to develop microfluidic devices for medical applications [22]. Membrane functionalities can be understood by modeling single-cell manipulator. In a similar way, cellular interaction with microenvironmental particles in extracellular space as well as its response to drugs and chemotactins can also be evaluated using cellular manipulators [23]. Biomimetic in vivo conditions can also be achieved by parallel-cell manipulators. It indicates the huge potentiality of micromanipulators for 3D tissue engineering [24]. Substrate features such as hydrophobicity, hydrophilicity, conductivity, thermal, and environmental factors also drive cellular integration with the substrates [25]. Functionalization of polymeric substrates such as PDMS, poly (methyl methacrylate)

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(PMMA), and peri-implantitis can also be made as functional biomaterials in form of alginate, chitosan, or functionalized surfaces for enhancing cellular micropatterning [26]. Physical cell patterning includes inkjet printing, optical tweezers, laser-based cell patterning, acoustic force patterning, and electrokinetic force
based cell patterning. Cellular ink is used for bioprinting in inkjet printing technology. Printing can be done in three technologies: (1) continuous inkjet printing, (2) drop-on-demand (DOD) printing, and (3) electrohydrodynamic jet printing. A schematic representation made in Fig. 23.3 illustrates the above techniques. DOD printing technology is used widely for 3D bioprinting due to its ability of high precision and lowest chance of contamination. Yusof et al. [27] had reported a noncontact approach to print single cells by using inkjet printing technique which comprises dispenser chip to deposit droplets, a sensor to detect the cells, and an automation tool to print on specific substrates. Optical and optoelectronics tweezer technology guide cellular movement by optical forces. Optical tweezers consist of laser that emits laser beam, creating radiation pressure for cellular migration. Dielectric properties of cells and intrinsic charge across cellular membrane make it feasible to manipulate cell printing remotely [28]. This technology also provides high precision of positioning small arrays of cellular bodies in define lengths. However, due to its limited manipulator zone, this technology may be limited in precision while considering printing for homogeneous geometrical patterns [29]. Radiation forces can be reduced by using optoelectronics technology. It reduces the radiation energy by 100,000 times. It uses high-resolution virtual electrodes for manipulating single-cell structures by intervening their dielectric properties. Laser energy is also used to propel cells from one source film to an acceptor substrate in laser-based cell patterning technology. It comprises the following subsets: (1) laser-induced forward transfer (LIFT), (2) absorbing film-assisted LIFT, (3) biological laser processing, (4) matrix-assisted pulsed laser evaporation direct

Figure 23.3 Schematic representation of (A) continuous inkjet printing: printing is made through an ink-collector located above the substrate, and the flow of ink-drop is regulated by charge-electrode at the tip of the syringe; (B) drop-on-demand printing: flow of ink is being regulated by the resistive or piezoelectric element present at the tip of the nozzle; and (C) electrohydrodynamic jet printing: speed and thickness of printing are regulated by the applied potential field between the tip of the nozzle and the substrate bed.

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writing, and (5) laser-guided direct writing [30]. Local patterning of breast cancer and cell mobilization during the stage of angiogenesis at the region of tumor can also be modeled using this printing technology [31]. Surface acoustic waves (SAWs) are being used in acoustic force patterning. SAWs are being generated by electrodes, which are excited at different frequencies (10
1000 MHz) and deposited on piezoelectric substrates [32]. In such kind of manipulation system, displacement resolution of cells depends on number of nodes and frequencies formed [33]. Cell functionalities and viabilities as a function of time and energy are required for cell patterning had been studied by Ding et al. [33]. Electrokinetic forces, also known as dielectrophoresis (DEP) is considered as an active method of cell manipulation. A combination of electrokinetic forces and hydrodynamic effects has been utilized to achieve cell trapping to specific areas without damaging them. Polarization and dipole moment features of cell membranes have been complemented in this DEP for patterning cells to guided channels [34].

23.2.4 Cryopreservation techniques of cells for tissue engineering Cryopreservation of cells is a key part in routine cell culturing due to the need to store particular cell types for long time periods. Cryopreservation protocols involve medium exchanges, centrifugation, addition of proteolytic enzymes to detach adherent cells from their substrates, addition of cryoprotective chemicals such as dimethyl sulfoxide, and slow freezing (21 C/min). After thawing, preparing cell cultures for experiments can take days to weeks [35], depending on the cell proliferation rate and other biological processes. These laborious procedures are necessary to reduce lethal damage to the cells caused by freezing. Two major obstacles in cryopreservation are the growth of ice crystals during thawing, which cause physical damage to cellular membranes and internal organelles, and the toxic effects of cryoprotectants and local high solutes concentrations during the procedures [36
38]. Success of the cryopreservation can only be estimated after thawing, by assessing the viability and functionality of the thawed cells. Cryopreservation of cell cultures is therefore expensive and time consuming. At present, many cell types (e.g., neurogenic stem cells) are considered as unsuitable for cell culture use due to the inability of cryopreserving them [39]. Delivery and storage of cells that are difficult to cryopreserve are a bottle neck in cell-based therapies and diagnostics. Cryopreservation is the process of preserving the live cells at very low temperatures by using certain preservative agents that enter the cells, prevent dehydration, and form intracellular ice crystals. There are many methods of cryopreservation such as ultrarapid freezing and thawing, controlled-rate freezes, freezing with nonpenetrating polymers, vitrification, and equilibrium freezing. The main disadvantage of cryopreservation is that it may lead to cell death and destruction of cell organelles during the freezing process due to the formation of ice crystals. To overcome the damage occurring during cryopreservation, certain agents such as ice-binding proteins and antifreeze proteins (AFPs) are used during the process of cryopreservation that will reduce the damage of the cells

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by limiting the effect of ice crystal at cellular microenvironment and keep the cells alive for a longer duration. A good example of development of a microfluidic device platform for tissue engineering applications is micro-fingered cryo chamber for stem-cell preservation. This device is used for the cryopreservation of intact cells, specifically stem cells, attached to a substrate that is ready to use upon thawing using inductive nanowarming via nanoparticles. This system involves a precisely controlled freezing method using a directional freezing technique in combination with microfluidic cold-finger device (for live imaging) that imposes minimal chemical and physical damages to the cells culture, thereby facilitating high percentage of viability upon nano-conductive thawing. The gradual freezing of the sample enables control over ice crystals shape, size, and position. Incorporation of various types of AFPs in the freezing process may further improve control over the ice-recrystallization parameters. However, due to their contradictory effects to certain cell types (such as immobilization, interference in intercellular signal transduction, increasing cytotoxicity) and high cost for minimal quantity, their usage is limited. Microfluidics (MCFs) were fabricated using in-house 3D printer and REM with PDMS. The microfluidic channel network contains one to three inlets and one outlet. A copper wire serving as the cold finger will be embedded in the center of an elliptical pool. The volume of the device is of 0.2 mL and the height of the device will be 10
20 mm, and it will be verified by a surface profiler. The low height of the device will be deliberately kept in order to ensure fluorescence imaging with low background signal. A schematic of this platform is given in Fig. 23.4. The vitrified samples are heated in 1-kW hotshot inductive heating systems with 2.75-turn, water-cooled copper coil or a 15-kW custom-built system depending on the volume of the sample (1
18 mL). The 1-kW system is operated at a magnetic field of 25.13 mT (peak, volume-averaged field strength) and 360 kHz. Higher fields up to 60 kA/m at 175 kHz will be achieved on the 15-kW system. Nano-

Figure 23.4 (A) Design of the microfluidic chip with three inlets and one outlet. The copper tip acts as a cold finger. (B) Experimental setup. The MCF device is placed on a copper plate, and two thermoelectric coolers determine its temperature (T1). The cold finger is cooled by a third thermoelectric cooler shown in purple (T2) that is externally controlled independently of T1.

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warming rates from samples of interest are subtracted from controls (no msIONPs) to estimate the heating rate and specific absorption rate by the temperature rise method as previously reported [40].

23.3

Biomaterials at microscale

Microfluidic platforms can be systematically used to synthesize micro to nanometerscale functional biomaterials aiming at their variable configurations, multiplex structural configurations, and indigenous properties in a versatile and user-friendly route. Microfluidic systems offer numerous advantages when compared with conventional techniques. The biomaterial size (dimensionality), shape, morphology, composition, and crystallinity can be finely tuned, based on requirements. Once the particles convert into the micro-and/or nanoscale regime in a specific biomedical device, the activity could be enhanced because of its low supply of reagents, fast heat, mass transfer, etc.

23.3.1 Composite microparticles Over the years, significant research has been done for organic and inorganic microparticles as microfluidics for interactive drug delivery of proteins, peptides, and drugs [41]. Microparticle-based assemblies are extremely effective in in situ drug delivering for specific requirements to circumvent biological barriers. The overall benefit is to lead the system into highest level of productivity of drug delivery, reduction of doses, and minimization of side effects [42,43]. Li et al. [44] have shown a novel composite core
shell microparticle assembly constructed of gelatin methacrylate (GelMa) as the core in combination with poly (lactic-co-glycolic acid) (PLGA) shells for sustained drug delivery. A colloidal solution of GelMa and PLGA oil has been used together as the raw materials for the synthesis of these microparticle assemblies. In addition, the microfluidic dual emulsion templates, comprising both hydrophilic and hydrophobic moieties, for example, hydrophilic doxorubicin hydrochloride and hydrophobic camptothecin, might be supplied, respectively. In tissue engineering and other areas of biomedical applications, scaffold-based substances with induced porosity have found lot of attention in recent times. The scaffolds based on micro- or nanofibers have the unique advantages over others for having the ability to control the exact pore size and distribution, the orientation of the pore network, and hence can encourage cell proliferation and growth with tuned microenvironments that simulate the physiological conditions. Photo-polymerization is one of the popular cross-linking techniques that are potentially used for microfluidic spinning. Cross-linking via chemical and ionic polymers is popularly used to prepare a variety of microfibers which are inherently biodegradable and biocompatible. The formation mechanism is based on the creation

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of covalent or ionic/secondary bonds within the polymer chains and cross-linkers. Typically, the polymers are alginate, PLGA, chitosan, gelatin
hydroxyphenylpropionic acid, or a combination of multiple polymers of this family of polymers. Zhao et al. [45] have successfully synthesized photo-cross-linked scaffolds
based porous chitosan
gelatin agglomerate for the application in chondrocyte culture in vitro. This unique combination of biopolymeric materials fabricated using chitosan
gelatin hydrogel system offers excellent biocompatibility and can foster the growth of chondrocytes useful for cartilage tissue engineering.

23.3.2 Particulate biomaterials at the nanoscale Various polymer molecules with low molecular weights or increased monodispersity of inorganic nanostructured species have shown that this technology can be extremely advantageous and convenient in various applications related to biomedical devices. It is worth to mention that the accurate control over particle sizes ranging from microto nanoscale, composition, and surface properties, their morphologies, uniformity, and favorable structures is essential in development of microfluidic systems. A diverse range of materials has been synthesized based on particles of inorganic origin for microfluidic systems. These include metallic and metal oxide nanoparticles (NPs), organic
inorganic hybrid nanostructures, functionalized NPs, quantum dots, NPs with core
shell structure, and conventional porous structured materials. The crystallinity, dimensionality, and morphology of these particulate materials in nanoscale make the difference. Once a particle transforms in the nanometric range, it demonstrates radically different properties compared to bulk materials of same composition. Metallic NPs (e.g., Ag, Fe, and Au) have the propensity of continuous ion release, and these factors is occasionally utilized in various range of products [46], including antimicrobials, where these particles act as a storehouse that liberates cytotoxic ions by reacting in an aerobic atmosphere. Various biomedical devices, prosthesis, and transplantable parts have been coated with silver NPs to achieve bacterial, yeast, or fungi population growth control. Microfluidic synthesis in biomedical applications is aimed at coating the as-synthesized NPs or stabilizing them in order to reduce their cytotoxic effect for human cells at the same time to be antimicrobial against pathogenic prokaryotic cells. Chen et al. [47] have successfully synthesized microfibers carrying silver NPs and graphene oxide inserted in bacterial cellulose operated in a microfluidic platform. Their results demonstrate excellent antimicrobial action for Gram 1 and Gram 2 bacteria without any cytotoxic effect on human liver cells because of the controlled kinetics of Ag-ion release. Yang et al. [48] have shown that chitosan microparticles attached with silver NPs have been synthesized by droplet microfluidics with enhanced antimicrobial reaction against Escherichia coli and less cytotoxic effect against fibroblast cells. Several groups have designed integrated as-synthesized “Au” NPs for developing microdevices. These NPs have the ability to respond via confined surface plasmon

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resonance for biosensing of protein and polypeptide molecules [49]. The novel metal NPs offer superior sensitivity and very low detection limits for the target species. On the other hand, Knauer et al. have synthesized NPs based on noble metals (Pt and Ag) with core
shell and multishell type structural configurations [50] via a segmented flow microfluidic approach. These outstanding NPs display excellent optical properties and have the potential in plasmonic applications. Biomaterials functionalized with various species to inherit desired physicochemical properties is essential to pass natural biological obstacles (such as gastrointestinal barrier, stratum corneum located on the upper skin layer) which are providing a safeguard against incursion of foreign effects. In order to supply drugs to the affected organs, circumvention of the barriers is a prerequisite by constructing drug transporters with special characteristics. Hence, appropriate modification in the physical and chemical nature of drug species is essential. In the initial stages the particle size of the drugs was reduced from micron to nanometer size ranges. It has been found that small-scale particles are effective in overcoming the barriers; smaller NPs with diameter ,6 nm can be discharged by kidneys [51]. On the contrary, particles possessing sizes larger than 200 nm are prone to settle in organs such as spleen and liver. However, particle size does not always possess a threat in developing advanced drug delivery system for biomedical applications. Albanese et al. [52] have shown the interdependence between basic NPs characteristics such as its morphology, size, chemical activity, compositions, and free surface charge and their biomolecular response, kinetics, dispersion, and toxicity.

23.3.3 Fibrous biomaterials at micro- and nanoscale The intricate biomaterial structures are made possible by engineering the biological materials from a zero-dimensional to 3D structures with the help of microfluidic technology. This technology enables to produce unique nanostructures in the microfluidic-based system. Thus construction of functional hybrid microfibers is possible with the opportunity of the microfluidic synthesis method [53]. The current challenge is to culture cells in most of the polymeric based continuous microfluidic substrates. However, poor cell interaction with the materials is a major problem. Shi et al. used a method to fabricate a hydrogel photo-crosslinkable cell-responsive methacrylamide-modified gelatin (GelMa) fiber with the help of the microfluidic method to promote cell adhesion. The manufactured hydrogel has the quality to promote higher cell viability by micro-grooved channels and has a potential application in tissue engineering [54]. Zhang et al. developed nanofibers with variable gradients by modifying its surface which has a great impact on tissue engineering. Electrospinning method has been used to produce these nanofibers with the assistance of a microfluidic approach [55]. Kang et al. developed hydrogel microfibers with features that can be tuned digitally with respect to chemical, structural, and morphological properties with the aid of a microfluidic system along with digitally controlled program flow control. The fibers have been made by using a spinning method that mimics the spinning mechanism of spiders. The process relies on the

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microfluidic chip combined with a digital fluid controller where fibers have biomimetic properties [56]. Ahn et al. used alginate to develop one-dimensional microfibers and designed in a flow focusing way as drug carriers. In addition, ampicillin was introduced into fiber to improve the antibiotic property and tested in vivo in rats for the wound healing purpose [57] (Table 23.1).

23.3.4 Sheet biomaterials Two-dimensional (2D) sheet materials are characterized by larger dimension in two primary directions (x
y) which is several orders magnitude smaller than those in other direction (z), representing a flat or “sheet”-like surface topography. Elsayed et al. created monodispersed microbubbles with the aid of polymeric alginate shell (as shown in Fig. 23.5). The method used to create microbubbles was T-junction micro-bubbling method. In order to make porous polymeric films combined with coarse capillaries, these microbubbles have been used. Polymeric films were composed of alginate and phospholipids along with a surfactant. Moreover, developed porous films can be used as coatings over implants as well as in drug-delivery applications [67]. Microfluidic channel plays a major role for the encapsulation of cells within it, especially in the tissue engineering field. These microfluidic channels act as a reservoir that helps in controlling cell activity. In order to make cost-effective microfluidic devices via cell-laden agarose hydrogel, the opportunity of the soft lithography method was used by the researchers. In addition, variable dimensions of microchannels were microfabricated for the cell encapsulation in the developed molds for cell culture, and it had shown higher cell viability near the microchannels after 3 days [68]. Researchers fabricated SiO2 nanotubes with micropatterning. Silica-coated hydrogels (collagen fibrils) were put in microgrooves in a Teflon-based microfluidic chip and have shown better osteoblast differentiation when cell cultured with MC3T3 preosteoblasts [69] (Table 23.2).

23.4

Methods for cell patterning and cultivation

Spatial control of cell position is a crucial task for multiple aspects of tissue engineering, including generation of radial pattern of liver lobules, linear pattern of smooth muscle, and radial form of anal sphincter. This section represents novel methods capable of patterning cells in a certain shape on a substrate. One possible way of patterning cells over substrate is along the nodal positions of resonance vibration over cellular substrate during its cultivation process. A metallic cellular cultivation device can be used in such cases for generating resonance frequency using a piezoelectric material fixed on various surfaces of the substrate. These resonance vibrations can also be used for cellular differentiation, confirmation and maturity of specific types of cells from their mesenchymal stem cells (MSCs) type by mesenchymal lineage process. In this process of cell patterning, viable cells are

Table 23.1 Tabular representation of composition, morphology, and application area of fibrous biomaterials at micro- and nanoscales. S.I.

Composition

Morphology

Application area

References

1.

Hydrogel Calcium alginate

3.

PLGA

Fibrous

4.

Methyacrylamide-modified gelatin (GelMA) BMP-2 incorporated PLLA Laminin Chitosan, PCL, PEO, PU, and PVP PGA polymer, PEUU, and PGA/PLLA composites PLGA hydrogel

Microgroove surface fiber Membrane Fibrous Nanofibrous mats Fibrous

Core
shell fibrous scaffold for codelivery of Co and BMP-2 Designing scaffold to regulate the arrangement of cells for tissue engineering Design and construction for regenerative medicine and tissue engineering Templates for the production of fiber-shaped tissues Osteoblast differentiation of MSC Nerve regeneration Wound dressings and skin tissue Tissue-engineered heart valves

[58]

2.

Cross-linking polymeric chain Flat microgroove pattern

Thin sheets

3D polymer scaffolds

[66]

5. 6. 7. 8. 9.

[59] [60] [54] [61] [62] [63,64] [65]

BMP, Bone morphogenic protein;3D, Three-dimensional; MSC, mesenchymal stem cell; PCL, polycaprolactone; PEO, poly (ethylene oxide);PEUU, poly(ester urethane) urea; PGA, polyglycolic acid; PLGA, poly(lactic-co-glycolic acid); PLLA, poly-L-lactide; PU, poly urethane; PVP, poly(vinyl pyrrolidone).

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Figure 23.5 Schematic diagram of microfluidic T-junction. G, Gas; L, liquid; LG, liquid/gas.

patterned along the nodes of vibration. Pressure-head amplitude optimization is also a critical criterion for cellular patterning.

23.4.1 Cell-patterning techniques A soft lithography
based technique is used to prepare layers with cell responsive and cell-repellent components. The pattern is having grooves and ridges to provide cell guidance [70
72]. These grooves and ridges also help the alignment of cells and even their orientation. In this objective the main aim is to see the effect of grooves on the shape, size, orientation, and growth of cells on the surface of the substrate [72]. For example, this is relevant in the context of vascular tissue engineering. There are different types of blood vessel that may vary in size but have common general features. Each vessel type has three tissue layers: tunica intima (the innermost layer), tunica media (middle layer), and tunica adventitia (the outermost layer). The tunica intima consists of endothelial cells (ECs) (0.2
0.5 μm thick, 10
15 μm wide, and 20
50 μm long) and connective tissues. The basal lamina (80 nm) is in between ECs and the underlying connective tissue. Next is the tunica media layer which is composed of layers of smooth muscle cells (SMCs) and is 200 μm
1 mm thick. The outermost layer, tunica adventitia, consists of collagenous fibers, and its thickness is 10% of the total thickness of the blood vessel in the case of elastic arteries and 50% in the case of muscular artery. The most popular material for the fabrication of microfluidic system is PDMS, but its hydrophobic nature and permeability to water vapor, resulting in evaporation inside the device, are incompatible with long-term cell viability. Hence, more biofriendly materials such as PMMA or gelatin are preferred in current studies for developing microfluidic chips. A schematic diagram is shown in Fig. 23.6.

23.4.2 Bioreactors Microfluidic-based bioreactors are most convenient bioreactors which can be used for developing various cellular constructs. These tissue-engineered constructs can

Table 23.2 Tabular representation of composition, morphology, and application area sheet
based biomaterials. Sl. no.

Morphology

Composition details

Application areas

References

1. 2.

Sheet Sheet

Cardiovascular tissues Wound dressing

[6] Azuma et al.

3. 4.

Sheet Sheet

5.

Sheets

Poly(ester urethane) urea Alginate, chitin/chitosan, and fucoidan powder blend on hydrogel sheet Poly(-caprolactone) layered by Schwann cells Gelatin methacrylate (GelMA) incorporated with carbon nanotubes AlginateCalcium chloride

6.

Sheet Scaffold

7.

9.

Tri-layered fibrous sheet scaffold Stripe-patterned heterogeneous hydrogel sheets Specific patterns

Silica

10.

Electrospun sheets

Chitosan hydrogel

8.

Poly(lactic-co-glycolic) acid sheets coated with fibronectin Microfabricated poly(glycerol sebacate) and fibrous poly(caprolactone) Poly(dimethylsiloxane)

Nerve regeneration Myocardial tissue engineering Scaffolds Tissue Engineering Drug delivery Adhesion of cells Mosaic hydrogel sheets Barcoding Cell patterning Myocardial tissue engineering Tissue-engineered heart valves

[8] Junka et al. [9] Shin et al. [53], Jun et al. [10] Leng et al.

[5] Kitsara et al. [11] Capulli et al. [65]

Coculture of multiple cell types Tissue engineering

[7] Kobayashi et al.

Cell guidance Drug delivery Cartilage tissue engineering

[53] Limongi et al.

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Figure 23.6 Cell-patterning technique for development of blood vessel of different layers integrating over a single microfluidic chip.

Figure 23.7 Microfluidic chip containing tissue construct and connected to external fluidic circuitry to construct a microfluidic bioreactor.

utilize low-set-up environment for successful in vitro characterization. These bioreactors also house different layers and segments for performing distinct activities of cellular construct. Fig. 23.7 shows a schematic diagram of a bioreactor with microfluidic housed tissue-engineered constructs.

23.4.3 Microfluidic devices for cell manipulation Microfludic device fabrication for cell manipulation includes microchannels with wells of graded depth. The incorporation of features afterward within microfluidic chamber in connection with microchannels is a challenging task. It requires technology such as pseudo-grayscale with backside diffused light lithography

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(pGS BDLL). pGS BDLL produces multilevel microfluidic devices [1]. Isotropic pattern of binary opaque and transparent rectangular patterns were created at lowresolution photomask on transparent paper to produce long microchannels with SU-8 negative photoresist. Numerical computer aided design (CAD) modeling was used for creating pGS masks, which was further printed by CAD-Art services (Bandon, OR) at high density dot-per-inch. Distinct pseudo-gray areas were fabricated using the rectangular mask mentioned previously. In this way, 80%
100% areas of transparent zones were created on photomask for pGS BDLL. Differential graded transparent pGS masks were used for BDLL. Negative photoresist, SU-8, was painted on cover-slides at a thickness of 100 μm with an electrospinning machine. These differentially graded transparent masks undergo UV treatment. The resultant PDMS channel heights were examined under fluorescence microscopy. The method of fabrication is given in Fig. 23.8.

23.4.4 Microenvironment on cell integrity Microenvironment plays a crucial role for cellular integrity. An enhancement in cell viability and metabolic activity of MSCs within microchannels under induced shear stress might be a resultant of enhanced cell
extracellular matrix (ECM) integration and increase in shear-stress induced focal adhesion points. The organization and uniform distribution of focal points is the main component in cell
ECM interaction, and thus it plays an important role in temperature-induced stress of adherent cells to maintain cell attachment and sustainability. An accumulation of depolarized focal point complexes indicates the role of focal point of adhesion (FPA) complex in outcomes of thermal stress. The MSCs were stained with F-actin and vinculin to identify F-actin and FPAs complex organization. When stress applied to the

Figure 23.8 Schematic of BDLL: (I) represents the UV transilluminator from which UV light passes through (II) glass substrate and (III) pGS photomask (a mask on which the pattern of the microchannel is drawn) to expose (IV) SU-8 substrate to produce microchannel of exactly same pattern by regulating the amount of energy allowed to pass through the spatial region of pGS photomask. (V) Quartz glass is used to cover SU-8
coated glass, whereas (VI) a tinted film is used at the top of the set-up for preventing it from exposed on light. BDLL, backside diffused light lithography; pGS, pseudo-grayscale; UV, ultraviolet.

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cytoskeleton of cells during thermal stress by virtue of intracellular ice formation beyond the threshold, focal points can withstand and cells eventually die. In addition, adhered MSCs, in presence of shear stress induced FPA, display low density of FPAs before and after thermal stress.

23.5

Microfluidic cell culture models for tissue engineering

23.5.1 Basal lamina Talking about biomaterials, 3D hydrogel cell cultures using microchip technology has become an active field [73]. Klotz et al. had also investigated the influence of hydrogel (GelMA) on 3D cell culture techniques, and its biofunctionality and mechanical tenability was explored [74]. In another study, Sart et al. [75] observed that these bioreactors provide enhanced control of hMSC microenvironment. In another study, feasibility of self-organization of perfusable capillary networks was explored on a microfluidic platform [76]. This self-organizing capillary bed was used to connect separate culture wells that can contain various tissue explants [77]. Yi et al. had developed complex in vivo environment of central nervous system using 2D and3D neuronal cell culture. Effective role of exosome delivering neurogenic microRNA enables induction of efficient differentiation process during neurogenesis. In a recent study, Kilic et al. [78] had used a new multilayer silicone elastomer device, and differentiate pluripotent human (NTERA2) cells into neuronal clusters interconnected with thick axonal bundles and interspersed with astrocytes, resembling the brain parenchyma. In another study, Karimi et al. [79] had discussed about the role of nanoparticles in microdevices for studying toxicity on neural cells. A microfluidic-based basal lamina mold fabrication technique is shown in Fig. 23.9. A 3D skin model for checking the skin regeneration is established in COMSOL (as shown in Fig. 23.10). To know the physiology of the silk fibroin
based hydrogel for the treatment of skin wound, the model of skin injury will be developed for testing the 3D skin substitute developed. For this, finite element method is proposed for numerically stimulating the healing process. In this part, focus will be on developing skin as a model organ to illustrate the mechanism of wound healing using silk fibroin - chitosan-based hydrogel.

23.5.2 Vascular tissue Vasculogenesis is an essential activity for the successful implementation of cellseeded scaffold. It provides continuous supply of nutrients and other essential factors for growth and proliferation of seeded cells. Growth factors are often being used to trigger the process of vasculogenesis in 3D scaffolds, but they produce uncontrolled and scattered generation of vascular structures within the scaffold.

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Figure 23.9 Schematic representation of the processing step involved in micromoldingbased pattering technique.

Figure 23.10 Multilayered microfluidic chip.

Microfluidic technology has recently extended to the generation of patterned vasculature. But these structures are limited to 2D forms. A microfluidic chip housing 3D-patterned blood vessel synthesis along with its quality monitoring system is a specialized synthetic vascular unit. Microchannels

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will be fabricated on a biocompatible silicone polymer with multilayered perforation to enrich the channels with nutrients and continuous oxygen supply. 3D porous scaffolds will be drawn over the patterned microchannels at different depths (with existing fractals-like interconnections between each layer). Then MSCs are seeded on each of these layers of 3D scaffold with sufficient perfusion. Defined layers of microchannels are coated with tropoelastin to ensure construction of capillary patterned aggregation of cell types. Elasticity and conjugation of the blood vessel materials with flowing blood is evaluated in order to test the feasibility of implanting the vessel construct. Continuous monitoring is done using oxidase biosensors in order to evaluate the concentration of nitrous oxide in blood over a period of time. Shear stress developed near the inner wall of the fabricated vessel in presence of flowing medium is monitored by laser Doppler velocimetry. A microchannel circuit
based chip is designed using COMSOL Multiphysics software and transport equations for capillary based flow and diffusion-based flow is solved using Semi-Implicit Method for Pressure Linked Equation method. The solution of the model is represented in the form of evaluation of pressure distribution, normal and shear-stress distribution at various sections of all the wall of microchannel. Three inlets are considered at three different levels, whereas all the outlets follow a unique height (level) after leaving the section of scaffold platform. Evaluation of abovementioned parameters helps us to optimize the fabrication of the chip that will offer less cell death and delamintion during the flow of media through the channels during dynamic incubation. Fig. 23.11 shows a schematic of the geometry with incorporation of the process of differentiation of MSCs into different form of cells. In blood vessels, along with the ECM proteins, the inner intimal layer consists of a layer of EC, the middle layer tunica media is composed of SMCs, and the outer adventitial layer mainly composed of fibroblast cells. In this study to replicate these

Figure 23.11 A schematic representation of multichannel microfluidic chamber housing vascular graft and sensory devices (biosensors) for measuring flow parameters for real time evaluation of metabolic state of vascular cellular matter.

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multilayer microchannels of varying stiffness and cells, we used soft lithography techniques. Three layers of PDMS (Sylgard 184 Silicone Elastomer Kit) with varying stiffness and geometrical features are fabricated based on the computer design proposed in the previous section. To fabricate each layer, silicon wafers with the required pattern was obtained by spin-coating negative photoresist, followed by soft bake, UV exposure, post-baking, development, and hard baking (or the PDMS channels can be fabricated by following technique—to fabricate the channel polyvinyl alcohol layer was 3D printed and PDMS is casted on the Poly vinyl alcohol (PVA) sacrificial mold). Long and narrow PDMS microchannels of diameter 200 μm and aspect ratio of 0.4 were cast using the patterned silicon wafer (PVA mold). Different ratios of PDMS cross-linker have been used to fabricate three individual layers that represent mechanical properties of each layer of the blood vessel. Further, the PDMS microchannel layers were assembled and sealed using oxygen plasma treatment. To enhance the attachment of the cells to the microchannels, a polycaprolactone (PCL) scaffold were 3D printed on the PDMS microchannels. Then prepolymer of tropoelastin containing photo-initiators was introduced into the scaffold and crosslinked under UV light. The porous PCL scaffold ensures the confinement of the cells and absorption of the nutrients. Moreover, to accurately mimic the blood vessel, the microchannel layers were cocultured with ECs and MSCs. The MSCs will be obtained from different sources, including umbilical cord, bone marrow, and adipose tissue and maintained in Dulbecco’s modified Eagle’s medium supplemented with fetal bovine serum. The microchannels were first seeded with ECs and cultured until the cells become confluent and line the inner surface of the channel. Then MSCs cell suspension was injected into the EC lined channels and incubated for 6 hours. The MSCs that are not attached to the inner surface are removed. The fabrication of tropoelastincoated 3D scaffold inside the microchannel improves the cells adhesion, and the interconnecting channels provide the distribution and flow of nutrients. MSCs cocultured with ECs in the presence of fibrogenic growth factor and vascular endothelial growth factors supplement the media favors the vasculogenesis.

23.5.3 Liver Liver is an organ present only in vertebrates that has roles in detoxification, synthesis of proteins, and production of biochemicals useful for digestion. All these functions are carried out by interactions between blood and hepatocytes. An adult human being will have liver weighing approximately 1.75 kg of which 1.2 kg is represented by hepatocytes. Hepatocytes are arranged in single-cell thickness and have blood supply from portal vein and hepatic artery. These two blood vessels create a space called liver sinusoid lined by ECs and phagocytic Kupffer cells. These ECs were interrupted by pores allowing the passage of solutes into a precapillary space. This precapillary space is rich in ECM molecules, collagen type IV, fibronectin, and heparin sulfate. Major causes of liver failure are alcohol consumption and hepatitis virus. Among hepatitis viruses, hepatitis B and C virus have capability to develop chronic hepatitis.

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Figure 23.12 Sectional view of microfluidic chip showing liver cells regeneration chamber.

Although various methods are available to treat liver failure, whole organ transplantation is the ultimate choice. This technique has many limitations such as less availability of donors, risks involved while transplanting, and cost effectiveness. These limitations lead to the use of tissue-engineered liver devices that will serve as an alternative to treat reversible liver failure. Major components of these liver assist devices will be hepatocytes. Artificial liver should provide 10% of the liver mass. Artificial liver device designed for normal 70 kg patient requires approximately 120 g of liver tissue which is not feasible in every case. Due to this limitation, alternatives are explored such as transformed hepatocytes, freshly isolated, cultured, and partially transformed hepatocytes. In the case of transformed hepatocytes, neoplastic hepatocytes under culture conditions are assessed for purity, function, and pathogenicity. Major concerns among these cells are the potential of transformed hepatocytes to show the changes in gene expression under culture conditions and the risk of tumorigenic cell infusion into the system of the patient. Freshly isolated hepatocytes are another source used where primary hepatocytes are isolated which can be further cryopreserved. Porcine hepatocytes are widely used as they are readily available, reach maturity within 6 months, have similarity with human organs regarding size, and are relatively disease resistant. When using cultured hepatocytes, cells are cultured by replacing tissue culture media with plasma or blood of the patients. In partially transformed hepatocytes, viral oncogenes are introduced, which generate cell lines that have the property of immortalization where they divide expressing differentiation. The devices designed for liver assist majorly concentrates on providing environment to attain high cell viability and their function. The major requirement of device design is the transport of nutrient, toxins, supply of oxygen to the patient. Currently several devices for liver assist systems were authorized by Food and Drug Administration to ensure and evaluate the safety of human clinical trials. A sectional view of microfluidic chip for liver segment is shown in Fig. 23.12.

23.5.4 Bone Bone is a complex tissue for supporting internal vertebrates. The two major cells responsible for bone formation are osteoblast and osteoclast. The bone cell, osteoblast,

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Figure 23.13 Schematic representation of different layers of microchannels for housing different concentric layers of bone segment.

secrets bone substance whereas osteoclast absorb bone tissue during growth. The bone disorder can be caused due to trauma, disease, or injury. The bone graft is the most effective therapy for bone replacement; it can be autogenous bone graft and allogenic bone graft. In the case of autologous bone graft, it is difficult to obtain donor bone. For the successful regeneration of bone using tissue engineering, the basic requirement is the synchronization between the bone cells, biocompatible polymer matrix, and bioactive signaling molecules. For bone regeneration, porosity must be greater than 90% and pore size must be in between 100 and 350 μm [80]. With recent advancement in biomaterial fabrication, scaffolds with drug-delivery capacity, growth factors, and antibiotics are fabricated to enhance bone ingrowth [81]. The addition of growth factor such as fibroblast growth factor, bone morphogenic protein, plateletderived growth factor, and transforming growth factor-beta into scaffold promotes angiogenesis and osteogenesis. And the autologous stem cell sources used for tissue engineering are umbilical cord blood, adipose tissue, and bone marrow. Microfludicdriven bone tissue engineering construct is schematized in Fig. 23.13.

23.6

Conclusion

Tissue engineering is a useful tool in the field of regenerative medicine. Fabrication of tissue engineering constructs is supported by two fundamental pillars: miotic microenvironment provider and biomaterial. Cost as well as the acuteness of the regenerative tool is highly governed by the abovementioned pillars. In this chapter, special emphasis was made on developing tissue-engineered constructs at a

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micro-confined zone. These zones were defined as microfluidic channels which facilitates regulated flow of culture media at multivectorial form. Microfluidic principles in correlation to tissue-engineering bioreactors were well described in the first section of this chapter. It was followed by a comprehensive illustration of various biomaterials that can either be used within micro-confined environment or can be directly printed within defined microchannels. Lastly, a section was dedicated for the development of microfluidic-based bioreactors for different regenerative segments of tissue constructs. In this chapter a section was also dedicated describing the procedure of preservation of tissue constructs using microfluidic devices such as cryo device. At the end, this chapter reveals a new dimension and scope for further development and understanding of scientific insights of behavior of biomaterials and cellular constructs within synthetic microfluidic channels. Also, an insight can be develop to understand the role/influence of material and topographical features of microfluidic channels for integrating biomaterials with seeded cells.

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2. Credi C, De Marco C, Molena E, Pla Roca M, Samitier Martı´ J, Marques J, et al. Heparin micropatterning onto fouling-release perfluoropolyether-based polymers via photobiotin activation. Colloids Surf, B: Biointerfaces 2016;146:250
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Michael Gasik1,2 1 School of Chemical Engineering, Aalto University Foundation, Espoo, Finland, 2School of Health Sciences, University of Eastern Piedmont, Novara, Italy

24.1

Introduction

The need to develop tissue substitutes and regeneration platforms is one of the most demanding and challenging applications in modern tissue engineering [1,2]. Many types of biomaterials are presently available for use in different implants [3,4]: metallic alloys, ceramics, composites, polymers are all used, whether with or without living cells, medical substances, or some other additions such as antifouling or antibacterial factors. There is also a growing trend of use of different scaffolds in tissue engineering applications to support and promote correct tissue formation [5]. Three-dimensional biomaterial structures (scaffolds) are highly desirable matching the biomechanical properties of the tissue [6] and closely mimicking in vivo behavior (facilitating cell adhesion, growth, and tissue formation [7]). Such biomaterials assist the body to rebuild the damaged tissue, and eventually they minimize associated pain and healing time [8,9], and their mechanical properties are becoming one of the most critical issues to address [5]. Mechanical and biomechanical stimuli play a very important role in an organism development, homeodynamics, and homeostasis, starting from an embryo [10]. On the level of a living cell, mechanical pathways are one of the three fundamental pathways [electrical, (bio)chemical, and mechanical] to communicate with the cell and with its environment. On the organ level, we may note a difference between biomechanics itself (also as a scientific discipline), mechanobiology (as a set of phenomena acting on the cell level) and practical engineering biomechanical cues for characterization of biomaterials, medical devices and implants, as well as for analysis of their interaction with tissues and the host systems. Traditional biomechanics focuses on how tissues and biomaterials would actually perform the structural, functional, and locomotory actions, and it involves a great deal of body and organs physics, mathematics, and computational analysis [10,11]. Biomechanics aligns with somewhat simplified models of both materials and tissues, complying with available the computational power and simulation time, as a complete picture on subcell, cell, tissue, organ, and body levels would be too Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00024-6 © 2020 Elsevier Ltd. All rights reserved.

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difficult to calculate in necessary details. For example, a soft tissue can be approximated with many various mechanistic approaches (hyperelastic or viscoelastic, with tensegrity model, and a number of empirical equations) yet mainly serving the purpose of description of experimental data used to derive these models [11], because addition of multiscale features would require too many assumptions or too many resources to be of a practical value. One interesting comparison could be made here with the mechanics of solids—eventually, the strength of any material specimen is thermodynamically determined with a great precision by mutual positions of constituting atoms, molecules, grains, and boundaries, but such computation would require enormous efforts with not much practical outcome as it is impossible to make another specimen with exactly the same structure and composition. Mechanobiology can be described as complex, combined, and synergetic effects of acting mechanical forces that modulate morphological and structural fitness and features of the tissues—especially bone, cartilage, ligament, and tendon [10]. The role of mechanobiology phenomena on tissue formation, development, maturation, and regeneration is huge, and not yet well understood [10
13]. For skeletal tissues, mechanobiology addresses the question how these same load-bearing tissues are produced, maintained, and adapted by cells as an active response to biophysical stimuli in their environment [10,11,14
16]. Biomechanics and mechanobiology at large are combining in musculoskeletal tissue engineering, in treatment and prevention of conditions such as deformities, osteoporosis, osteoarthritis, and bone fractures, also involving interaction with biomaterials as implants [10,14,17
19]. These topics are already covered by a vast number of literature, specialized journals, and conferences as well as handbooks and textbooks, to which reader is recommended to refer for more details, for example, Refs. [19
27]. In this chapter the main focus is on the third area, providing a review of some practical applications of biomechanical methods for the purpose of design, characterization, analysis, quantification, and regulatory approval of biomaterials. Due to the vast number of biomaterials and their forms, one usually need to make a choice of the biomechanical testing methods, and this choice might not be so evident as it may look. There are certain constraints in existing standards (easier to make a test but more difficult to understand the meaning of the data for clinical applications) and protocols (ad hoc test devices and methods that are not easy to replicate in another laboratory). One important gap is still a lack of proper understanding of biomechanical properties and their meaning in medical care, perhaps with the exception of some cases [5]. An example from the author’s experience was a question at orthopedic conference, asking how many surgeons think that the elastic modulus of a biomaterial is very important for the implant. This answer got a full consensus, which was expected. However, the second question asking how many doctors present did actually use this knowledge of elastic modulus in selection of the implant for a patient, resulted in zero positive answers. Thus one may ask—are we measuring correct properties of a biomaterial, and if “yes,” why and how these data have to be used in a clinical practice?

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24.2

601

Biomechanics and mechanobiology

From a physiological point of view, any solid component of the organism (from bones to cells, and even viruses) may be considered as a living tissue [11]. One can roughly distinguish “soft” tissues (muscles, connective tissues, gut, brain, and other organs) and “hard” tissues (bones) by their mechanical flexibility and other “softer” mechanical properties. Naturally, the quantitative metrics of what is soft and what is hard are not very clear as such classification requires comparison with some references. On the cell level, mechanical forces cause changes in size, shape, number (induction of cell death or proliferation), position, and gene expression of cells [28]. Transduction of mechanical forces between neighboring cells or across different tissues has been shown to be critical for the coordination of cell contractility and dynamics, and it is known that coordination and integration of forces by longrange force transmission and mechanosensing of cells within tissues produce largescale tissue shape changes [28,29]. Extrinsic mechanical forces also control tissue patterning by modulating cell fate specification and differentiation, and this can be used in directing, for example, stem cells fate by mechanical cues [30,31]. The spectrum of these mechanical and micromechanical cues and their impact on the cell behavior are often referred to as mechanotranduction [31
33]. All cells and organisms are mechanosensitive [31,32], relying on conversion of mechanical stimuli from their physical surroundings or from within the organism to electrochemical or biochemical signals, which then regulate a wide repertoire of physiological responses, Fig. 24.1. Mechanotransduction also has a fundamental role in regulating physiological phenomena in other specialized tissues that are not directly involved in sensory functions [33]. Mechanotransduction objective is to describe the cellular processes that translate mechanical stimuli into bio(electro) chemical signals, enabling cells to adapt to their environment [28]. The most straightforward action of cellular mechanosensing is force-induced conformational changes in mechanosensitive proteins, resulting in opening of

Figure 24.1 Simplified schematic of biochemical and mechanical signaling transduction between the cells and substrates they are attached to. When a cell expresses both at the same time, the neighboring cell may receive them at different time points (first and fast mechanical, and then more slow biochemical) leading to multiplexing response. Biochemical transport might be facilitated by increasing fluid flow, also imposing additional shear on the cells.

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membrane channels or altered affinities to binding partners, thereby activating next signaling pathways [33]. External mechanical strain transduced by specific molecules expressed on cell surfaces (such as integrins) induces the instantaneous disturbance and the reorganization of actin cytoskeleton, which influences cell shape, stimulates signal transduction pathways, and induces transcription of genes resulting in a switch between cell growth and behavior [29,34]. Any changes in normal intracellular force transmission through changes in cellular (or extracellular) structure and organization can lead to altered molecular forces acting on these proteins, resulting in attenuated or increased mechanosensitive signals [31,33]. Despite extensive research over the past decades, many components of cellular mechanotransduction remain incompletely defined or understood [28,33,35]. Specific challenges still remain in the repair or regeneration of tissues that predominantly serve a load-bearing biomechanical function [36]—they are not limited to only bone, cartilage, tendons, or muscular tissues but are very critical in cardiac, cochlear, dental, pulmonologic, gastrointestinal, or gynecologic applications. This biomechanical function might be also combined with electric signal generation or transduction (bone, heart) or enhancement of fluid flow (cartilage, hemodynamics). Mechanobiological interactions between cells and scaffolds can have a critical influence on cell behavior, even in tissues and organs that do not serve an apparent biomechanical role in the body [36,37]. Many researches in cellular mechanotransduction focus on extracellular physical forces conversion into intracellular chemical signals, but mechanical forces that are exerted on surface are known to be also channeled along cytoskeletal filaments and concentrated at distant sites in the cytoplasm and nucleus [35]. Surface membrane receptors (integrins, cadherins) were shown to have a central role in mechanotransduction [32]. The ability of cells to perceive extrinsic mechanical forces influences tissue size and architecture not only by changing their adhesive and cytoskeletal organization on short timescales but also by influencing their fate specification and differentiation on longer timescales [28,34,38]. Why mechanotransduction presents an interesting and somewhat unique way of driving the development of cells and tissues? A simplified comparison case [35] shows that a small ion such as Ca21 with an apparent diffusion coefficient of B100 μm2/s needs about 25 seconds to pass across 50 μm distance of a cytoplasm, and the triggered cellular response molecule needs another 50 seconds to travel back (at average velocity of B1 μm/s). On the contrary to this, mechanical stress waves propagate along tensed cytoskeletal filaments on the same distance within B2 μs (at a velocity of B30 m/s), Fig. 24.1. Therefore cells can receive many more same or other biomechanical signals during the time when one chemical signal is still traveling, indication about a million times faster response [35]. Thus physiological local stresses of B10
100 Pa can lead to a local range of deformations up to 10
100 μm covering one or more adjacent cells, which might be seen as “mechano-paracrine” signaling. Inside the cell, nuclear stiffness is clearly higher than of a cytoplasm [35,39,40], and it might also facilitate longdistance force propagation in the nucleus (one can imagine a stiffer ball inside the viscous fluid, which is subjected to dynamic excitations—stress waves and inertia

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effects will help to these forces transfer when the conditions are favorable). This was reported to facilitate mechanotransduction compared with a nonprestressed homogeneous structure with the same total input mechanical energy [35]. Besides the “good” cells constituting common tissues, biomechanical factors are also important for bacteria and bacterial attachment to surfaces, leading to biofilms [41]. It is also logical to assume they act on virions and viruses inside the cells, but that scale is of nanometer level and may require some different biophysics to be analyzed. Mechanical environment affects dispersal of biofilm organisms due to a body fluids flow and associated motion and micromotion of the surfaces in question [41], which has also implications on treatment of biofilm. These mechanoenvironmental factors determine the degree of deformation of a biofilm due to an applied force (shear and more complex deformations [42]), and challenge biofilms during growth [41]. Biofilms thus might be considered as some kinds of “composite hydrogels” [43]. Increased fluid-related or mechanical shear resistance [44], presence of locking cations [45] (e.g., Ca21, exopolymeric matrix production, and composition [46,47]), and changes in quorum sensing [48] are all known to affect the strength and stiffness of bacterial biofilms [41]. From mechanistic point of view, most practically important medical and environmental biofilms do behave as viscoelastic materials [44,49], even though values of the commonly measurable viscoelastic properties vary by many orders of magnitude [50]. However, there are no yet extensive consistent studies of viscoelasticity of biofilms, which might be very different for biofilms inside porous or on a rough surface implants, where traditional rheology test methods are not suitable [41,51]. Even less is still known about biomechanical properties of complex structures when biofilms and tissue components are present together, and how these properties affect development and reactivity of such co-cultures.

24.3

Mechanobiology and biomaterials functionality

Given the difficulties in matching all of the complex behaviors and architecture of native tissues, a key issue in the development of engineered repairs and implants is the prioritization of various biomechanical properties as design parameters, as it will be difficult (if not impossible) to completely match all of the material properties of native tissues [36]. On the other hand, there is also an understanding that may not be necessary to match all of the material properties of native tissue a priori in view of the remodeling potential of the implanted tissue in vivo—indeed, some degree of “mismatch” is required to ensure thermodynamic-like reasons to have necessary driving forces for tissue regeneration aiming on the restoring functionality [36,52,53]. Although mechanical strain usually exhibits complex patterns in vivo, its impacts and related mechanisms can be investigated with simplified in vitro models [54]. Depending on system geometry, a combination of translational, rotational and/or multiaxial strain can be applied to cells to study the interaction between strain and cellular responses [29]. Priority has often been placed on a

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single parameter such as the compressive or tensile “proper” modulus (often confused or refereed as “stiffness”) with of a tissue [55
57], with yet unknown relative importance of all of the different properties [29]. Many characteristics of cell (morphology, function, fate) can depend strongly on substrate “stiffness,” that is, (in the most common sense) the ability of the adherent material to undergo deformation in response to applied force [29]. Here, it is important to define what is meant by the term “stiffness.” There is a confusion in literature where stiffness is expressed in Pa (unit of pressure) or in N/m (unit of surfacespecific energy), being also different for static and dynamic loading. When testing non-uniform specimens (such as explants or ex vivo samples), data are often provided in units of force (N), as it is not evident how this force values could be normalized or converted, respectively, into stress and material (and not that particular specimen) properties. The authors prefer to avoid “stiffness values” unless they are more rigorously obtained and compared with proper control or baseline. Instead, as shown next, invariant or idempotent properties (those representing the true material functions) might be much more beneficial in practical use. As an example, one may consider the use of elastic modulus, which definition originally fits only linear elastic materials and only for very small deformations, as that was prescribed by the theory of elasticity for centuries. Elasticity theory does not tell anything about the time-dependency of this property, neither about the rate (speed) of loading or unloading nor its hysteresis due to cyclic load. As almost all materials and tissues one deals with are clearly not elastic, it is a great (and clearly not correct) simplification to artificially reduce experimental data to some fixed numbers. When such numbers are published and cited, their connections to conditions and assumptions, where they have been deduced, is usually lost and thus practical value minimized. When loading
unloading cycles are applied on the tissue successively up to the same stress level, the stress
strain curve is gradually shifted until the mechanical response of the tissue enters a stationary phase and the results become reproducible from one cycle to the next [11]. This phenomenon is due to the changes occurring in the internal structure of the tissue, until a steady state of cycling is reached. This initial phase of behavior common to all living tissues is usually used as preconditioning of the tissues prior to experimentation [58]. The purpose of testing is to obtain simple, general laws describing the macroscopic behavior of the materials, in order to determine their mechanical properties, and predict their response under defined conditions [11]. If these preconditioning steps are not properly documented or followed, the stiffness analysis can produce rather different results. It is however not easy to define terminologically what is “good” and what is “proper,” even for implants of the same kind, such as orthopedic or dental cases [5]. What can be more precisely defined is the type and grade of biomaterial used in that implant, biomaterial composition, surface state, chemistry, etc., as these parameters are now being measured with a great precision and can be also documented to allow follow-up. As a classical example, one may consider tissue regeneration during fracture healing or an orthopedic operation. It was assumed that tissue formation guidance

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in a local biomechanical environment is guided by two main factors: local mechanical strain and local fluid rate [17,18]. The combination of these leads to mechano-regulative index (MRIX), the value of which comprises the sum of octahedral resolved shear strain γ in the solid phase (Eq. 24.1) of the tissue and the relative local velocity magnitude |U| between the fluid and solid components of the tissue (Eq. 24.2): γ5

2 3

ffi qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
2
2 εx 2εy 1 εz 2εy 1 ðεx 2εz Þ2

MRIX 5

80 γ 1 jUj 3

(24.1)

(24.2)

where resolved strain γ is calculated from differences of the principal (x,y,z) components of the strain tensor [11,24]. Based on analyses and validation outcomes [17,18,25
27], lower tissue strains and lower fluid velocities favor bone formation (MRIX , 1) and higher values favor fibrous tissue (MRIX . 3), whereas intermediate values are favoring tissue regeneration toward cartilage. MRIX values have been successfully used in modeling and simulation of the tissue formation [19]. In real life the velocities of the fluid and mechanical deformations are not independent variables, as one might anticipate from Eq. (24.2). When biomaterial or tissue is dynamically stimulated, structure
fluid interaction generates its own fluid flow, shear waves, or (at high rates) even shock waves, which cannot be easily included in the previous case [5,51]. It is indeed would be highly oversimplified to assume that MRIX value alone is sufficient to define the relationship between the mechanical parameters and tissue formation. Many tissue models are limited by number of assumptions made in those models (tissue properties and their dynamics). They do not much take into account, for example, implant surface geometry or tissue
implant interface development [20,21]. Gross characterizations of threshold mechanical stimuli for periimplant bone formation (such as micromovement) are implant-specific, surface geometry specific, and site specific. Implant surface geometry was shown [20] to influence early tissue formation and consequently the early (primary) mechanical stability of implants and differences in the pattern of bone formation with different implant surface geometries—identifying the relationship between tissue strain and tissue formation has important implications to the design of bone fracture repair devices and engineered skeletal tissues [20]. The clinical implications of the surface geometry dependence are in preference of one implant or fixation tool versus another in terms of their osseointegration ability and rate, as well as in tolerances of micromovement or loading [20]. Although the mechanoregulatory model’s predictions [17,18,24] are consistent with experimental observations, it is important to note that models account only for a part of the regulation of tissue formation by mechanical factors, even they may be more important than mechanical stimuli in initial bone formation [5,23]. Considering the numerous approaches which have been followed toward soft tissue modeling, as well as the specific nature and conditions of the experiments, it is probably not

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possible to present one model as more reliable for any case than the others. It is also not possible to collect and review all the existing biomechanical models without losing the understanding of the natural mechanical behavior common to all of them [11]. Another example of practical significance is in effect of local (micro, nano) topology of the surface of any material to reactivity of cells and tissues [59
62]. Nontransformed adherent tissue cells generally must attach to a solid substrate for their survival in contrast to suspended cells [29]. It is commonly referred that as a reaction, the substrate “resists” this deformation proportionally to its stiffness and this repulsive force influences cellular behavior of more compliant and more stiff substrates [63
66]. Experimental data show that motile cells cultured on elastic substrates trend to align their shape along the direction of highest stiffness and move toward stiffer regions [67
69]. Normal fibroblasts on flexible substrate were shown to exhibit a decrease in the DNA synthesis rate as well as an increase in the apoptosis rate [70]. A simplified concept was that since each tissue in the body has a different “stiffness” ranging from brain tissue (B1
5 kPa) to cartilage ( . 40 kPa) and to calcified bone ( . 2000 kPa) [30,63,64], cells located on different tissue may show particular properties and reactions. For example, “appropriate” substrate elasticity can be used to guide cell proliferation, and expansion/control of differentiation of stem cells, and proliferation and apoptosis of various cell types [70
73]. There is also a direct effect of mechanical factors on immunological response of cells. Recently, the coined term mechanoimmunology [74] comprises study of mechanical forces acting at the level of ligand
receptor interaction leading to immune cells activation. In the most relevant to biomaterial interactions the rigidity of the substrate was shown to influence phagocytic ability, cell morphology, and elasticity [75,76]. The interaction via mechanical forces is not likely to be straightforward: there are data indicating that stiffer substrates (modulus of B240 kPa) lead to more proinflammatory mediators than very soft (modulus B0.3 kPa) substrates [77]. However, when the substrate is made to be highly anisotropic (e.g., [37], differences in tangential elastic modulus 400 GPa/5 kPa 5 80  106 times), it leads to significant downregulation of proinflammatory TNFα, IL1ß, IL12ß, IL6, CCL2, and COX2 cytokines for peripheral mononuclear blood cells [37]. We may conclude that it is reasonable to consider that biomaterial “stiffness” (meant here to be shown as elastic modulus) alone is not sufficient to characterize or forecast how immune cells would potentially react on it. One should take into account that stiffness in dynamical conditions might be rather different than on one in static conditions [5,11,51]. Recalling that major range of frequencies in the body is low (B1 Hz for blood flow pulsations, 0.5
10 Hz for main brain waves and up to B50 Hz in cilia beating frequencies), biomaterials stiffness need to be measured in these ranges for higher clinical relevancy [43,54]. When one adds to this mentioned speed of the mechanotransduction (B2
5 μs, so formally of B200 kHz rate), superposition of these harmonic signals might, respectively, amplify or cancel certain waveforms, leading to different outcomes.

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Another often referred property in mechanotransduction, mechanobiology, and the mechanical biomaterial
tissue interaction is “fluid stress” or “fluid shear stress” [11,13,21,23,24,78]. There is a common acceptance that acting fluid flows have a substantial impact on cell deformation and hence mechanotransduction, even the deformations caused by fluid
structural interaction are not equivalent to the same level of deformations caused by direct strain stimulus [78,79]. In biomechanics it is known that pure shear stress does not cause volumetric deformation of any material, and thus its effect is not equivalent (idempotent) to other modes of stresses acting on a cell, tissue, or biomaterial [11]. The main challenge in shear stress estimation is that the shear stress cannot be actually measured and its calculation usually involves too many assumptions that cannot be reasonably justified (expect maybe for a very simple models). Many studies rely on assumption of the tissue or a scaffold as a porous body with “equivalent channel diameter,” ensuring “laminar flow” with constant (and often “adopted”) properties of fluids, materials, and tissues. For mechanical side, there is also a variety of means how stresses and strains are expressed and calculated—the more deformation tissue faces, the more deviations come between different strain definitions [5,11,80]. This finally leads to very different values of the stress and shear strains, and it is no surprising that a wide range of the values is found in publications. In computer simulations, physical coupling is better as there is a possibility to link different phenomena and count for their cross-interactions [21,81], but more complex models require more parameters and this increases uncertainty in the outcomes. Since such models are commonly based on differential equations systems [16], solutions are also sensitive to selected initial and boundary conditions, sometimes lacking proper justification. As mentioned previously, in dynamic tests there is also a controversy in expression of strain rate (or shear strain rate), often used in rheological analysis [11,82]. In study [83], this was expressed in terms of acceleration faced by the cells. However, as mathematically acceleration is a second derivative of the displacement in time, the same acceleration is achievable by changing frequency, amplitude, or both, and if this is not specified, the uniqueness of its value is lost (Eq. 24.3): Strain : γðτÞ 5 γ 0 sinðωτÞ @γ 5 A sinðωτ 1 φÞ; @τ   qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi γ ω A 5 ðγ 0 ωÞ2 1 ðγ_ 0 Þ2 ; φ 5 tan21 0 γ_ 0 _ 5 Strain rate : γðτÞ

€ 5 Acceleration :γðτÞ

@γ_ 5 f ðτ; ω; γ_ 0 Þ @τ

(24.3)

Strain rate and acceleration cannot be identified whether due to varied frequency or varied strain amplitude, or both. The effect of frequency and deformation on the

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cells (and biological materials) are not equivalent—one may recall, for example, their different linking with the induced fluid flow and fluid stress. It is thus highly advisable to perform experiments and model in a consistent way and properly document major variables, as the reverse decomposition is often impossible [5]. The main questions on biomechanics of biomaterials and their reactions with cells and tissues [36] are thus both quantitative and qualitative. From the data point of view, one need to know the ranges of stress, strains, associated fluid flows, etc. (both intrinsic and as external stimuli) in tissues, biomaterials, and biomaterial
tissue interfaces under physiologic or pathologic conditions [36]. We may not fully understand how exactly mechanical factor and mechanotransduction works in vivo but we can select and modify some properties as design parameters in the development of engineered tissues [5,36]. There are several challenges in that way, and one of them is associated with the correct data availability. The present level of evaluation of mechanical function in biomaterials and tissue engineering studies can be considered as highly insufficient [84]. For example, of 205 analyzed articles on cartilage tissue engineering, mentioning a kind of some applied mechanical stimulation, only 29% has some quantified material properties [85]. But even then, the recommended and optimized values often have their scatter comparable to the average value [86
88]. Correct and detailed biomaterial testing is rather time-consuming and requires expertise to properly quantify nonelastic material behavior of tissue, which is also scarce in many dedicated biology labs [89,90]. This also has impact on clinical applications [91], as viscoelastic response of tissues (besides, e.g., bone, cartilage, or heart) are also very important for diagnostics and functionality of lungs [92], liver [93,94], brain [95], breast [96], skin [97], prostate [98], vocal folds [99], etc. Indeed, there is huge potential to use viscoelastic measurements in the assessment of the quality of tissues, in terms of both positive and negative physiological or pathological states [91]. However, even simple viscoelastic analysis of modern biomaterials and biological tissues requires the knowledge of entire history of deformation in order to take into account also its rate-dependency [11,91]. This consequently imposes the use of numerical or phenomenological models which might fit experimental data very well but nevertheless fail to predict tissue behavior outside of this range [43,53,89].

24.4

Methods and challenges

Conventional mechanical testing or characterization of the material itself, usually involves the determination of strength, hardness, fatigue, coatings adhesion strength, etc. as well as so-called materials properties such as elastic (Young) modulus, shear modulus, viscosity, and loss tangent (for dynamic loading). Many biomaterials, including those for implants, are being nowadays tested under different mechanical loading schemes, specified by various standards [89,100] such as ASTM F384, F1160, F2977, F2900, and F2150. Despite the standards and combinations of the various mechanical testing methods available, provided quantitative

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assessment of native and engineered tissues are still hardly compatible. Highly complex and unique structure function relationships of tissues present the challenges inherent to translate these characteristics in engineered tissues [36,89]. Besides conventional (tensile, compressive, bending, etc.) tests for materials themselves, there are also dedicated tests for implant materials such as fatigue tests (e.g., ISO 14801 for dental implants). These tests are targeting on the determination of a few parameters only, such as tensile strength, high-cycle fatigue limit, and they are mostly destructive [5,89]. Their main purpose is to determine the practical limits of materials in service conditions from mechanical point of view only. Standard mechanical tests do not usually involve any kind of biological factors and are of a very limited use for mechanobiology. Biocompatibility and other biological type in vitro tests evaluate biomaterials’ ability to work in vitro, such as ISO 10993. Tests are being carried out in respective culture wells or similar devices with only goal to access the effect of materials [in direct contact or via an extract (indirect testing)] on living cells in static conditions (the standard ISO 10993 does not allow any mechanical stimulation, and its use was consequently criticized in literature [101]). There are known many attempts of simultaneous application of biological objects (cell cultures) and movable materials specimens [102]. In bioreactors for tissue engineering applications, mechanical forces are usually poorly controlled, and, as a result, realistic stresses and strain acting on the material are not possible to evaluate properly [84,103]. Such an approach is limited by the laws of physics as these stresses and strains cannot be measured in principle, but only calculated (only real forces and displacements can be measured directly). Too high uncontrolled local stress or deformation is known to cause cellular toxicity, involving multiple yet unclear mechanisms [104]. The combined static and dynamic biomechanical properties of biomaterials are crucial for the final success of the treatment, and there, a high correlation between in vitro conditions and expected in vivo tissue regeneration must be sought [15,105,106]. As pointed out [107], the synergetic effect of correct mechanical stimulation is greatly dependent on the scaffolding material, its environment, and the cell presence. This shows the need for consistent simultaneous analysis to compare different biomaterials and to get conclusions about these features. In general, the testing and evaluation of biomaterials and tissues always have a challenge as most of them have nonlinear mechanical behavior [3,5,11,12,88,108,109]. For example, hydrated collagenous tissues possess anisotropic and nonlinear mechanical and transport properties that vary significantly with site [110
112]. Not only these properties are critical to the overall function of the tissue, but the interactions among the different phases can result in sophisticated stress
strain and pressure fields that can only be predicted using nonlinear inhomogeneous or fiber-reinforced models [11,36]. Hence, one always has to consider two sides of biomechanical characterization: a proper experimental method and a consistent relevant material model (or its numerical approximation) to interpret the results of testing. The first challenge is with the selection of experimental methods of loading or deformation of a biomaterial or tissue specimen. Apart from nonlinearity and need of

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preconditioning (which duration and parameters are not known a priori [58]), there are also side effects that might be significant. For example, test grips pressure on soft tissues might affect the apparent modulus of the specimen to a great extent [80]. In view of the lateral specimen size and the difficulty to have effective clamps on soft tissues, tractions would induce important artifacts rendering the process imprecise [113]. Uncertainty in models approximating the specimen behavior, properties variations, and their interpretation often lead to undesired results: it is very difficult to design of biomaterials or a scaffold that has to correspond to “optimized” permittivity of annulus fibrosus being 180 6 127 nm2/(Pa s) [86] or meniscus modulus 80 6 110 kPa and permittivity of 70 6 80 nm2/(Pa s) [88]. For bone permeability values between 1027 and 107 nm2 (so even 14 orders of magnitude!) were reported depending on the specimen, model used, and method of measurement [87]. The second challenge exists in understanding of experimental results and extraction of the relevant information from that data. As mentioned previously, due to nonlinearity and preconditioning there are usually simple or straightforward models or approximations one can take to calculate required parameters [5,11]. For example, for bone stress analysis, different poroelastic and viscoelastic models were applied [11,13,100,114
116]. In terms of general behavior of viscoelastic materials, they to some extent might be approximated with a modified biphasic theory used for biomechanical description of highly hydrated tissues such as cartilage [51,53,89]. There solid and fluid constituents might be presented as nearly idealized solid and fluid, linked with share of stresses between the phases, whilst keeping the compliance in deformation and shear rate. Soft tissues in unconfined conditions are highly compressible, and since fluid is almost incompressible at these physiological conditions, it is squeezed out of the channels. This also constitutes the reason as to why tissue elasticity and viscosity vary when derived in different modes of stress application. In many experimental systems, stress is applied locally over a relatively small surface area, allowing water to flow away from the stressed region to its surrounding as also possible in natural conditions [41]. Here, one would observe the size effect caused by diameter of the indenter or a compression plate, as boundary conditions for such test would be less predictable; that is why the use of indentation method (as, e.g., used in analysis of cartilage biomechanics) might not be recommended. The outcomes of the models usually are results of compromise between computational costs, ability of validation, degree of simplification, and need of many fitting or phenomenological parameters [11,114], for which physical sense might be lost after becoming just a set of mathematical coefficients. The implementation of simplified constitutive law to model soft tissues should be performed with caution in numerical simulations [113]. Misleading numerical results could be obtained in situations where multiaxial stresses are present, but even more difficult is the application of various numerical approximations (such as Prony series, chained Maxwell
Kelvin models or just a phenomenological combinations for viscoelasticity) which require quite a number of fitting parameters, by which physical sense is lost and for which applicability for forecasting the behavior is very weak [113,117,118].

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In experimental validation of models, in vivo tests are commonly considered as a “gold standard” to determine how an implant or biomaterials would act in such conditions. However, a great burden of the variables and uneven conditions makes most of the animal tests essentially useless, raising the costs and slowing the introduction of new solutions. Most of the “well-went” in vivo tests fail in human clinical trials (success rates of 2%
15% only were reported [119]). The decisions from in vivo tests are often made on weakly controlled limited population of subjects with a high scatter. In Europe the Directive 2010/63/EC requires medical device manufacturers to move toward “3Rs” (reduction, refinement, and replacement) approach for human science. This drives manufacturers, developers, and clinicians to strengthen the application of alternative in vitro and in silico methods to obtain the maximal information from the intended product before undergoing clinical trials [120,121]. Most of the existing in vitro methods still rely on very limited conditions, are fragmented, and thus not sufficiently translatable to patient circumstances (recalling, e.g., the abovementioned biocompatibility evaluation using ISO 10993). It is notable that tissue engineering and biomaterials journals often warn that these biocompatibility evaluations have no scientific reasoning and done purely for regulatory purposes [101]. From practical point of view [3,5,80], implantation site is hardly reasonably squeezable down to microliters volume, where studies proceed without important natural-mimicking stimuli. It was anticipated that another approach should be introduced as “high-output screening” for medical devices [5]. On the contrary to known high-throughput screening (5massive number of specimens but with a limited number of readouts), it should aim at maximizing the amount of consistent data and information from the minimal number of tests or specimens. This could provide a great assistance for medical device producers, researchers, and medical doctors as it will minimize clinical tests, shorten time to market and improve lives, without compromising patient safety [5,53].

24.5

Biomaterials evaluation: a practical example

It is of a common understanding that biomechanical properties (and mechanical cues in general) have very important implications for cells, tissues, and organs, but the questions remain what exactly these properties are; how do we measure and quantify them; and how this knowledge would be translated into biomaterials design, development, and translation to clinical practice. From the previously discussed overview, it can be seen that too simplified approaches lead either to limited conclusions (more stiff substrates lead to osteogenic tissue differentiation) or require too complex models (with dozens of fitting or assumed parameters), generating too many drawbacks and uncertainties [5,30,37,61,114]. As of one practical case, one may consider articular cartilage (AC) repair, which is still one of the most challenging applications of biomaterial scaffolds. The damage and degradation of AC are not only progressing with age, obesity, or systemic

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diseases but also in the young and active population due to physical causes such as injury [52]. If untreated, these defects may progress toward osteoarthritis, affecting over 150 million people worldwide, mainly by the degeneration of hyaline cartilage in synovial joint lacking the ability of self-regeneration [122]. Natural wound healing, in full-thickness defects of cartilage, often leads to the formation of fibrocartilage [105
107,122], which is functionally and biomechanically inferior to the original hyaline cartilage making the tissue more prone to further deterioration and osteoarthritic changes of the joint [52]. Chondro-conductive and -inductive biomaterials are highly desirable to treat cartilage lesions at early stages before manifestation of the disease. The structure, functions, and biomechanical behavior of AC are very complex, highly anisotropic, and time- and loading history-dependent [123]. The AC consists of a relatively small number of chondrocytes surrounded by a multicomponent matrix, which can be imaged as a composite with 70%
85% water and remaining proteoglycans (proteins with glycosaminoglycans attached as a bottlebrush-like structure) and collagen [80,124]. Proteoglycans and water concentration vary through the depth of the cartilage tissue. Proteoglycans can bind or aggregate to a backbone of sodium hyaluronate (NaHA) of molecular weight of 2
4 MDa to form a macromolecule weighting up to 200 MDa [125]. The biomechanics of AC and synovial fluid is also complex and essentially nonlinear [8,126,127]. Not so many studies have coherently and systematically analyzed AC properties [57,128,129] due to variability of the samples, local inhomogeneity, and differences in applied biomechanical methods. Complex loading schemes are associated with significant variations of interstitial fluid pressure and fluid flow, making the interpretation of results more difficult [129,130]. The collagen-rich matrix behavior is highly nonlinear and requires rather sophisticated models to be described as a composite material [108]. Synovial fluid is well known to have non-Newtonian viscosity versus its composition, shear rate, mode of loading and the presence of other factors [52,131], and its proper characterization presents a challenge of its own. Most of the biomechanical properties of AC tissue reported experimentally are obtained with either confined compression [132] or indentation [133]. These measurements data are approximated with biphasic [61,134] or triphasic [114,135] theories, or even more simplified viscoelastic models. However, due to peculiarities of the AC tissue properties [136], it is difficult to compare results published with different studies, using various specimen types, methods, and testing devices. It was also reported [58] that fluid flow and flow-dependent phenomena may dominate the AC behavior in different testing regimes, and thus it is impossible to determine in general required recovery time. Aggregate modulus in a range of 50
120 kPa was reported for human, bovine and canine tissues by different sources [105,122,124,126,132,133,136], but often full test data were not available to make correct comparison (indentation usually produces much larger values [108,128]). Formal models of AC are missing essential features that limit their practical application only to specimens analyzed in those studies. Therefore it is a great oversimplification to characterize AC or scaffolds for AC repair by a set of one or two

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numbers without exact data on the test method and data analysis [52,114]. How one could solve these challenges? As there will be obviously many specimens and their versions to be tested, one would be interested in maximizing the number of outcomes and readouts to get first screening (lead development). Whether with or without added cells, the specimens need to be loaded into a sample holder and dynamically compressed at prescribed pattern, aiming at simultaneous measurement of bioactivity of the cells and viscoelastic properties of the specimens [89]. Here one may be tempted trying to load many specimens at the same time as being done in bioreactors. However, when many specimens are being stimulated simultaneously, there is less limited control of individual strain and stress states. Calculation of the properties of these materials will be only possible by fitting deformation curves to (preselected) theory [11] (such as biphasic [61,134] or triphasic [80,114,135] models). Application of small elastic strain model [11] to materials which clearly do not follow this rule (as was mentioned earlier) has led to oversimplifications decreasing the quality and questioning the relevance of the data. As reported by the authors [137], these theories indeed have failed to describe the behavior of biomaterials in such specific conditions. Similar methods [85,137
139] were found requiring too many assumptions, more new theories or additional independent experiments to extract true material data. Such approach has intrinsic flaws in measurement precision as none of them properly subtracts the support stiffness contribution, sample holder correction or effect of the intermediate layers introduced (resistance sensors, adhesives, magnets, etc.). Increasing number of specimens thus is paid off with less read-outs—for example, in method described in Ref. [139], only approximate elastic modulus was possible to measure, and the ability to use these data in clinical translation was very limiting. In order to get a comprehensive picture of a biomaterial behavior, one need to (1) define the application conditions and what exactly is required from that biomaterial, implant, or medical device in general; (2) select the proper testing protocol which mimics the most of those critical parameters and variables, relevant to clinical targets (structure, functionality, operability, safety, etc.); and (3) develop an approach how required variables would be obtained or extracted from the tests readouts and how these data could be used to make a reasonable prediction. The list of the features of the biomaterials testing, which would be important to know about, is shown in Table 24.1. Here one can highlight how important is the assumption of linearity, applicability of a material model (simple in some cases, but not so straightforward in others), or coherency. The latter term means that for many material properties that are linked together, they must not be measured in separate methods, as the outcomes might not be compatible. For example, for a scaffold of AC repair fluid transport is essential since the cartilage is avascular and transport of nutrients and metabolic products relies much on diffusion [80,127
129]. The relevancy of static tests for clinical conditions is that scaffolds need to better supportive for weight-bearing and undergo smaller deformation when surrounding tissue [52]. At dynamic (walking, running) loads, the situation reverses: under dynamic loading, one has to aim at more active fluid

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Table 24.1 Important features of a system for testing of biomaterials on their biomechanical properties. Conditions of the tests

What it means

Biomaterial linearity Geometric linearity Intermode linearity Specimen homogeneity Model-free Coherency in testing

A linear response of a material to an applied stimulus Small deformations acting on the sample Only one mode of loading is acting at one time Specimen structure/composition is homogeneous

History dependence High-output testing Nondestructive Invariant analysis Causality Clinical (physiological) relevancy

Do not need assumption of a material mechanistic model Obtaining of several (usually interdependent) properties in one experiment simultaneously to ensure they are linked Effect of the previous stimulation history of the biomaterial specimen Maximum data from the single test/specimen Specimen is not destroyed and could be retested if needed Extraction of material specimen parameters which are generalized for all relevant conditions Specimen response cannot come before stimulus Test conditions are close to clinical targets rather than prescribed by some standards

exchange to provide biomechanical stimulus to chondrocytes, to ensure fluid and nutrients supply and removal of metabolic products, promoting tissue regeneration [62]. Thus lower dynamic modulus and better fluid diffusivities are desired. These features have to be, however, compatible to the above requirements for static conditions as cartilage must work well in both these extremes. Porosity depends on the degree of cartilage compression which affects permeability, permittivity, and finally the local availability of the fluid source [122,124,126]. If the porosity or permeability of a scaffold is measured in one device and method, but viscoelastic properties with another, there is a risk that these data will refer to rather different conditions that will be difficult to match (as they could relate to different pressure gradients, for instance). The coherency approach does not require commonly used materials models (such as elastic, viscoelastic, or hyperelastic ones), as there is no objective to reveal exact mechanisms. For biological systems one often cannot set up experiments to measure all of the state variables [89]. Another important feature of measured properties is their invariancy—often referred to as “true” properties of a material. For example, elastic modulus can be strain and strain-rate (frequency) dependent, and it could also vary with testing time [5,11,140]. These values cannot be used for prediction before a proper correlation with time and other parameters is made, and even when made, it does not guarantee that this correlation will be suitable for prediction. Properly controlled prescribed mechanical loading of a biomaterial specimen with measurement of resulting strain

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(or vice versa) via same single probe-sensor element can be used to evaluate true time-invariant biomaterial functions [89,140,141]. Mathematically, time convolution operation (including loading history dependence) with idempotent analysis (which also includes nonlocality and obey causality, Table 24.1) allows bypassing the need of complex Fourier transform (usually employed in viscoelasticity), as it is known that a complex transformation is only feasible when the “stiffness” does not depend on frequency [142]—seldom the case in biomaterials and tissues. The test results should be capable to answer whether a biomaterial is good for its intended application as medical device or other product, how different biomaterials relate to each other in expected clinical performance, and how close the biomaterial specimen is to the properties of control specimen or natural tissue it aims to correct or replace. The example of the list of the properties obtained as readouts or outcomes (midpoints or endpoints) and their meaning are shown in Table 24.2. For the above example of scaffold for AC such tests were carried out in two different simulated synovial fluids with 1 and 4 mg/mL of sodium hyaluronate (NaHA) which mimics osteoarthritic and normal joint conditions, respectively [52]. The viscosities of these fluids are substantially higher than water: about 20 and 100 times for 1 and 4 mg/mL, respectively. This means that even small changes in permeability (a feature of the material structure) will affect changes in permittivity (a feature of a specific fluid flow thought the material structure [52]). Fig. 24.2 shows that under dynamic loading, there is indeed a nonlinear dependence of permeability of rhCol-polylactic acid (PLA) materials versus memory value [52,89] which is not the case for PLA without recombinant human collagen additions. Here memory value represents a “fading memory” showing how much specimen “remembers” its history from the previous cycle. It is interesting to note that effective permeability for rhCol-PLA increases with fluid viscosity on the contrary to hydrophobic PLA. Comparison of invariant dynamic modulus (at dynamic cycle load at 1 Hz) versus aggregate static modulus (at constant applied stress) also shows that PLA material static behavior is practically independent on NaHA concentration, whereas for rhCol-PLA it decreases (Fig. 24.3) [52]. Remarkably, dynamic parameter is highly NaHA-dependent for both cases (unlike shown above for permeability, measured coherently at the same conditions). This suggests that synovial fluid will stay likely more in rhCol-PLA scaffold (5 higher “stiffness”) than in conventional PLA, whereas the latter will lose more fluid at the same loading with the same time span. In other words, rhCol-PLA scaffold will better bear static loading (standing). In dynamics, rhCol-PLA has lower “stiffness” which in combination with higher permeability allows rhCol-PLA to provide better fluid exchange to supply nutrients to chondrocytes and to ensure hyaline cartilage regeneration [52]. With such a simple example, one may faster analyze and compare different materials combinations and optimize scaffolds, for example, osteoarthritic or normal cartilage correction as they clearly require different approaches. For high-output screening purposes many other parameters can be also coherently measured (directly or as a postprocessing of the data) just in one experiment: dynamic (shear modulus, bending modulus, fluid diffusivity, cyclic decay, etc.) and pseudo-static (displacement change, creeping compliance, history-dependent viscoelastic properties, equivalent fluid channel size, fluid diffusivity, etc.).

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Table 24.2 Some readouts and outcomes likely available from the enhanced simulation testing. Variables

What it means

Aggregate modulus Characteristic times Channel size

Apparent elastic modulus (Pa) of a biomaterial filled with a fluid when the fluid does not move Times describing a duration of a transient process like diffusion (s)

Compliance Dynamic modulus Fluid diffusivity Fluid source Injection kinetics Material permeability Material permittivity Material memory

Shear modulus Slope modulus Stress Strain Swelling pressure System viscosity

Equivalent diameter of the channel (μm) for fluid transfer in a porous or permeable specimen under specific test conditions Function of time as the ratio of strain to constant applied stress (1/Pa) Ratio of the amplitudes of stress to strain (Pa) Fluid diffusion coefficient (mm2/s) Rate of the fluid supply (mL/cm3 s) under dynamics For an injectable material from a syringe Characteristic of a porous material, showing its ability to conduct fluids (mm2 or nm2) Characteristic of a permeable material showing its ability to conduct specific fluid (m4/s N or nm2/Pa s) Function showing how much is a share of a material history “stored” during its testing cycle (from zero—no memory—to unity, full memory). Values over unity mean that recent material state overrides all previous states, although this seldom found in practice Property of the material interface under true shearing conditions (Pa) Ratio of stress to strain over the linear segment of the stress
strain curve (Pa) Ratio of acting force to area it uses (Pa) True logarithmic strain (Lagrange strain) Pressure (Pa) required to balance the fluid inflow causing material to swell when freely immersed in such fluid Real (invariant) viscosity of a biomaterial
fluid system (Pa s), that is, not the viscosity of fluid alone

When bacteria are added, or pH, temperature, and media composition are changed, new data sets can be compared and the decision made for the selection of the best biomaterial solution [5,54,89]. This can drastically reduce the time and efforts required for in vivo or clinical tests, as the solutions not showing a “good quality” would unlikely succeed in any extended deeper tests. Therefore some main features of the biomaterials enhanced simulation testing approach might be summarized as follows: G

Dynamic conditions are not equivalent to static tests, so are materials properties. Both tests need to be run coherently to get the complete picture. To be remembered: all materials properties are not single values but functions of the state variables and (very often) also of time.

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180

Dynamic permeability (nm2)

160 4 140 120 100

1

PLA

80

rhCol-PLA

60 40

0

20

1

0 0.00

4 0.10

0.05

0.15

0.20

0.25

Dynamic memory value at 1 Hz

Figure 24.2 Dynamic permeability of the two different scaffold materials versus dynamic memory value [89] and NaHA concentration (numbers near points, mg/mL). Note PLA scaffold permeability is a material structural feature (constant) whereas collagen-loaded PLA effective permeability increases with NaHA concentration [52]. 140

Invariant 1 Hz modulus (kPa)

120

PLA

4 4

rhCol-PLA

100 80 60 1

40

1 0

20

0

0 0

10

20

30

40

50

60

70

80

Static aggregate modulus (kPa)

Figure 24.3 Comparison of dynamic invariant modulus (1 Hz) and aggregate static modulus for PLA and rhCOl-PLA materials at different NaHA concentrations (numbers in boxes, mg/ mL) as reported in Ref. [52]. Lower dynamic modulus and higher static one lead to optimal combination of scaffold for cartilage repair. Note these values are time-invariant (independent on testing time) until material degradation will occur. G

The effect of sample preconditioning (especially for tissues) is rather strong and thus preconditioning protocol should be always justified and reported. It is also desirable that effect of preconditioning is quantified and its optimal parameters are described.

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Table 24.3 Reasons to consider correct enhanced simulation testing of biomaterials. G

G

G

G

G

G

G

G

G

G

G

G

G

G

G

G

G

Discover hidden materials features such as variation of fluid flow versus deformation Get many properties in one coherent test (not measured separately)—high-output screening with less costs and time Quantify, for example, materials permeability and swelling without external fluid pressure gradient Mimic, for example, surgery and postsurgery protocol or any relevant tissue conditions Eliminate unfeasible solutions at earliest stage Enable a long-term prediction of material behavior Improve risks reduction and quality assessment Increase ethical value (not tested on animals) Combine with other techniques in lab—enabling “companion testing” Beat competitors with more solid scientific/technical evidence Convince final users that new biomaterials are indeed better than old ones Present legally valid evidence in a case of litigation

All these should be done at proper conditions (the “PHRASE” concept: PHysiologically Relevant And Sufficient Environment) to assist translation of the test results to practice— the closer, the better. It is not enough to state “stiffness” or similar parameter in the data report, because also the conditions and details about how this parameter was measured are required. Similarly, such properties of the target tissues need to be known, but there should be some differences between tissue properties and biomaterial product to secure that tissue regeneration driving forces will act. The quantification of the specimen properties should be preferably done with as least assumptions as possible and likely model-free to ensure greater stability and enable predictive features of the data analysis. Testing ideally would be done on organ level if the main objective is to restore its functionality, and not just replicate the structure—in clinical reality, prompt, irreversible, and cost-effective recovery of patient without adverse effects and risks might be seen as an ultimate goal. Testing should be naturally compliant to regulatory guidelines but it does not need to adhere to standards—actually, it should go beyond the standards and being economically affordable (such as in the high-output screening).

To conclude, it might be useful to list also the reasons why properly design biomechanical tests need to be deployed in the laboratory or a production facility (Table 24.3). It is seen that besides scientific and technical values a proper test can also bring substantial savings in costs and time, as well as decrease (sometimes even eliminate) use of animal tests in biomaterials screening and discovery.

24.6

Conclusion and outlook

Biomaterials and tissue biomechanics have traditionally been too separate research areas with relatively little overlap in terms of methodological approach or research

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questions [143]. The development of new biomaterials and tissue engineered products has primarily been a research area for applied biochemistry, materials science, and biology. Biomechanics on the contrary relies much on the applied physics and mechanics, with very little incorporation of mechanobiology. It is now anticipated that the boundaries between biomaterials development, biomaterials, and tissue biomechanics are getting lower, and there is much more understanding in role of biomechanical cues in biomaterials applications [143]. For the purpose of the better integration between these disciplines, we suggest a definition of biomechanology, which might be postulated as a new engineering/ scientific discipline combining practically feasible, controllable, and measurable biomechanical stimuli and parameters on the level of tissues and organs, which has closer clinical and physiological relevancy. It would for example be focused on how mechanical stimulation would modulate tissue regeneration and biomaterial
tissue interaction, comprising essential parts of practical biomechanics, mechanobiology, mechanoimmunology and perhaps mechano-bacteriology (to add effects of pathogens and formation of biofilms) complemented by proper and correct biomechanical testing. The methods of biomechanical characterization of biomaterials are needed to be moved toward more coherent evaluation, which is aimed to go much beyond existing standards and with more focus on the purpose (application conditions). The same alloy or composite, previously known or certified with some tests, might not be anymore sufficient to be translated into clinical practice. One example could be in 3D printed and conventional titanium alloys, which were found to exhibit significant differences in mechanical properties [144] not originally expected. Such differences have to be quantified and taken into account already at the design stage as the new implant might generate different biomechanical environments at the same anatomical location, potentially leading to undesired outcomes. The “take home” message might be expressed with a motto do what’s right, and not what’s easy. Assessment of biomaterials, implants, medical devices, and tissue engineering products must be sufficiently simple, robust, with a high physiological relevance, and produce stable, invariant and idempotent outcomes (ideally suitable for model-free prediction) needed in development and clinical practices [5,53,54,84,89]. Such analysis should follow more practical, clinically relevant guidelines rather than sticking to (often outdated) standards. This may lead the way to discovery of new biomaterials and to significant improvement of benefits/risk ratios and safety of the patients.

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In vitro disease and organ model

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Emal Lesha1, Sheyda Darouie2, Amir Seyfoori3, Alireza DolatshahiPirouz 2,4 and Mohsen Akbari3,5 1 Tufts University School of Medicine, Boston, MA, United States, 2Radboud University Medical Center, Radboud Institute for Molecular Life Sciences, Department of Dentistry-Regenerative Biomaterials, Philips van Leydenlaan 25, Nijmegen, The Netherlands, 3Laboratory for Innovations in Microengineering (LiME), Department of Mechanical Engineering, University of Victoria, Victoria, BC, Canada, 4Department of Health Technology, Center for Intestinal Absorption and Transport of Biopharmaceuticals, Technical University of Denmark, Lyngby, Kgs, Denmark, 5Center for Advanced Materials and Related Technologies, University of Victoria, Victoria, BC, Canada

25.1

Model development

Modeling the complex microenvironment of human tissue and disease physiology is a difficult task that requires precise knowledge and understanding of the dimensionality, physiology, and cellularity of the specific tissue or organ of interest. Biomaterial and bioengineering research during the past few decades has attempted to tackle the understanding of such intricate systems by developing different bioengineering technologies and strategies. In this section, we will describe how recent advances in microgel and microfluidic technologies, stem cell and organoid research, and 3D bioprinting have improved the understanding of tissue architecture and vascular anatomy, which is leading the path to studying drug release and testing drug effects in vitro. Such advancements have potentiated the study of various conditions by improving disease modeling, such as cardiac valve disease, hepatotoxicity, or cancer. Finally, by increasing the understanding of basic physiology, we will discuss how recent bioengineering advances can led the path to improve and more diverse treatment modalities (Fig. 25.1).

25.1.1 Microengineered tissues 25.1.1.1 Photo-patterned microgels Modeling human organs is of great importance for the millions of patients who are affected by disease each year. Using in vitro developed microtissues to be assembled in a 3D microarchitecture, with the same cell
microenvironmental interactions as occurs in vivo, has shown great promise for the development of organ modeling research [1,2]. Herein, encapsulating cells within hydrogels is important for Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00020-9 © 2020 Elsevier Ltd. All rights reserved.

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Figure 25.1 Representation of different methods of model development, including photopatterned microgels [173], emulsion-based microgels [174], and bioactive fibers [35]. (1) Schematics of the fabrication process of glioma spheroids on agarose microwells. (2) (i) Fabrication of pH sensing alginate microfibers with pH responsing beads, (ii) microscope view of the generated microfiber, (iii) color variegation of the fibers based on pH of solution. (3) Schematics of the fabrication process of vortex-induced emulsion encapsulation method using a cross-linking light source. Source: (1) Reproduced with permission from Mirab F, Kang YJ, Majd S. Preparation and characterization of size-controlled glioma spheroids using agarose hydrogel microwells. PLoS One 2019;14(1):e0211078; (2) Reproduced with permission from Tamayol A, Akbari M, Zilberman Y, et al. Flexible pH-sensing hydrogel fibers for epidermal applications. Adv Healthc Mater 2016;5(6):711
9. (3) Reproduced with permission from Franco C, Price J, West J. Development and optimization of a dual-photoinitiator, emulsion-based technique for rapid generation of cell-laden hydrogel microspheres. Acta Biomater 2011;7(9):3267
76.

regenerating 3D tissues [3]. Microengineered hydrogels can potentially be applied to recreate the complexities of in vivo tissue constructs by assembling the building blocks in the shape of microgel tissue units to generate larger structures [4,5]. There are several methods for cell-laden microgel fabrication, such as micromolding and photolithography, by which patterned microgels with diverse shapes and sizes are developed [6
8]. However, due to fabrication restrictions and difficulties working with monomer solutions due to low viscosity and low process yield, other lithography systems such as continuous flow lithography and stop flow lithography (SFL) were developed [9
13]. These methods are highly compatible with different photo-crosslinkable polymers that can be used as cell matrix carriers. Poly(ethylene glycol) diacrylate (PEGDA) is one of the photo-polymers used by different groups

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to fabricate SFL-based cell-laden microgels [13,14]. Different organs have been regenerated using cell-laden microgels. As an example, in vitro bone regeneration was conducted from bone marrow-derived mesenchymal stem cells (MSCs)
laden gelatin methacrylate (GelMA) microgels (160 μm) [15]. They have been also used for cardiac disease applications using cardiac side population coated on the surface of the GelMA microgels [16]. 3D microgel patterns can be formed by a recently developed method called photo mold patterning, in which photopolymer hydrogel patterns are created using polydimethylsiloxane (PDMS) micromolds [17,18]. This technique is suitable for cytocompatible and geometrically controlled environment generation, and favorable for recapitulation of cell behavior in a specific organ model [18]. Assembling of photo-crosslinked microgels is of great importance for developing functional microtissue architecture that can resemble native organ behavior. To this regard, different assembly methods have addressed the importance of hydrophobic effects [19], hydrophilic templates [20], and physical templates [7,21]. In a study conducted by Liu et al., photo-patterned PEGDA cell-laden microgels physically assembled in the presence of a cross-linker carrying multiple thiol groups, which serves to stabilize the final microtissue structure [22].

25.1.1.2 Emulsion-based microgels Emulsion-derived microgels can serve as powerful building blocks for organ model generation. This technique is promising for generation of cell-encapsulating materials, relying on hydrodynamics to encapsulate single or multiple cells inside polymeric droplets [23
26]. Several polymers such as PEGDA [27,28], GelMA (gelatin methacrylate) [29,30], alginate [31,32], and PNIPAM [Poly(N-isopropyl acrylamide)] [33,34] can be used in this platform to create cell encapsulated micrometersized emulsion droplets that have the ability of controlling the cellular microenvironment, an important first step of the in vitro micro-organ model. Different studies have shown the viability and functionality of the encapsulated cells in these microgels. Franco et al. showed the capability of the PEGDA microgel as a matrix for various functional cell encapsulation. NIH3T3 fibroblasts, MHP36 neural stem cells, and bEnd.3 brain endothelial cells were successfully encapsulated in a PEGDA microgel modified with the fibronectin-derived Arg
Gly
Asp (RGD) adhesive sequences, as well as a matrix metalloproteinase (MMP)-sensitive sequence; all cell lines preserved their viability and function in vivo [35]. A liver tissue model in a threedimensional (3D) core
shell cell-laden microgel droplet was developed through a droplet microfluidic method, which included a hepatocyte-filled aqueous core and a fibroblast-laden alginate shell [36]. They used hepatocytes and fibroblasts cocultured with water
water
oil double emulsion droplets to represent an artificial 3D liver organ in a droplet. Liver specific related assays such as the rate of albumin secretion and urea metabolism all showed advantages of this method as compared to two-dimensional (2D) monoculture of hepatocytes. Among the different disease models, cancers models have been studied the most using the droplet microfluidic method [37,38]. Droplet

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microfluidic devices are considered an exceptional method for providing tumor cell spheroids encapsulated in an extracellular matrix. Through this method, Wang et al developed a capillary microfluidic device for fabricating multicore double emulsions with external
internal connected porosities [39]. These microcarriers included a suspension of a hepatocellular cell line, HepG2, in an extracellular matrix (ECM)
like biopolymer containing Matrigel and alginate, which formed liver tumor spheroids in the pores with the capability of exchange of nutrients and lack of applied shear forces to cultured cells. In other experiments, with the aid of a droplet microfluidic chip and alginate microcarriers, a hepatocellular carcinoma spheroid was developed that could be controlled for the size and shape [40]. They demonstrated that the shape of the aggregates varies from spheroid to spindle and branch by increasing the alginate concentration from 2% to 5%. Cytoskeletal analysis and scanning electron microscopy observation of cell
cell contact showed that cells were densely packed and interconnected with viability of about 96%, making this a reliable model of liver tumor tissue.

25.1.1.3 Bioactive microfibers When considering bottom-up assembly strategies for generating spatio-organized comprehensive 3D tissue models, cell-laden microfibers (CLM) are considered suitable platforms for imitating organs in vivo [41,42]. CLM provide advantages for constructing critical building blocks of tissue constructs, such as blood and lymph vessels, nerve networks, and muscle fibers [43,44]. Moreover, assembly of fibers in a 3D structure is much easier than some of the techniques mentioned earlier [41,45]. Various polymers have been considered for creating CLM that were ultimately used for the development of a specific tissue model, such as alginate [46
48], poly(lactic-co-glycolic) acid (PLGA) [49,50], chitosan [51], PNIPAM [52], and even composite materials [53,54]. All of these hydrogels must accommodate cell attachment, either by their inherent structural chemical moieties or additive surface treatments. Moreover, each microfiber material must be compatible with specific processing protocols, which are generally categorized as extrusion, microfluidic based, interfacial polyelectrolyte (polyion) complexation (MIPC), and electrospinning [55
59]. Although all these methods use pregel polymer as a precursor of the microfiber, each one faces specific challenges. To date, microfluidicbased methods have been recognized as the most feasible route to form microfibers controlled by size and shape [60,61]. A major limitation for the MIPC method is the limited availability of anionic and cationic biomaterials that can be used to create bioactive microfibers [62]. In addition, the scalability of the approach is limited by the need for manual pulling for the fibers from the interface between the anionic and cationic droplets. Electrospinning is another method for producing cell-laden fibers by which different cells, including immortalized human embryonic kidney cells or embryonic stem cells (ESCs), have been encapsulated in polyvinyl alcohol
or polyethylene glycol (PEG)
based electrospun fibers [63].

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In regard to the generation of a primary organ model using microfibers, there are different methods for incorporating living cells. These comprise surface-based and encapsulating-based methods, the latter being further subcategorized to tubular, core
shell, and compartmentalized groups [41,51,64
68]. To date, extensive efforts have been made to model normal or diseased organs using assembled microfibers. In a liver model developed by Yamada et al. [65], alginate microfibers were cultured with hepatocytes and fibroblasts; functionality of the model was confirmed based on albumin and urea secretion assays, as well as relative ribonucleic acid expression levels [65]. Another study was done by Onoe et al. to develop novel meter long core
shell microfibers composed of ECM polymers and different cell lines, with the goal of recapitulating blood vessels and neural tissues [69]. Results from the study showed that double-coaxial microfluidic-derived fibers provide a microenvironment that properly regulates cell
cell interactions. Progress has also been made in recreating muscle tissue, as flat fibers with grooved patterns were developed that resembled the patterning of a myoblast cell [70]. In another study focused on type I diabetes treatment a pancreatic islets model was developed in which islet cells isolated from sprague dawley rats were encapsulated in collagenalginate composite fibers with a diameter of less than 250 μm. It was confirmed that encapsulated islets showed high cell viability and normal insulin secretion [71].

25.1.2 Microfluidic tissue models and microphysiological systems Microfluidic-based research platforms have innovated the way we think about tissue engineering and design. Emerging in the late 1990s with the advent of PDMS, microfluidics quickly became a critical component of bioengineering research, as 2D cell culture techniques had often not been able to create an accurate representation of 3D tissue properties. Microfluidic systems provide many advantages to the study of in vitro cell-to-cell interactions, including but not limited to controlled delivery of nutrients to the cells, oxygen control, and endothelial shear stress [72]. The ability to control these parameters is critical and allows the tissue engineer to create an environment that mimics the unique physiology and nutrient requirement for a specific tissue or organ system. Microfluidics also allows for manipulation of fluids in microscale channels, with specific characteristics that include laminar liquid flow and domination of viscous over inertial forces [73]. Diffusion is the dominant mechanism of mass transport for molecules within microchannels, which mimics the delivery mechanism of human tissues in vivo, including the tumor microenvironment [74]. Microfluidics has the capacity to combine multiple cell types in a biomimetic ECM with mass transport system and under appropriate mechanical and biochemical stimuli [75
77]. Microfluidic systems have successfully reflected the in vivo complexity of human tissues with a high precision and created attractive alternatives to the traditional in vitro tumor models. With regards to screening new therapeutic agents and their side effects on various organ models, microfluidic platforms introduce a new concept of microphysiological systems

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recapitulation, by which real in vivo cell
cell and cell
matrix interactions are mimicked in vitro [78,79]. Moreover, microfluidic systems are ideal for high throughput testing, which simplifies the reproducibility and scalability of the study. The concept of an organ-on-a-chip, which has been popularized in the past decade, has revolutionized the study of drug delivery and discovery systems. Organ-on-a-chip systems are a conjunction of microfluidics, tissue engineering, and stem cell technology. They create a 3D environment that allows for testing and simulation of key functions of a living physiological system, including organ-specific vascular and epithelial interface, tissue organization of parenchymal cells, as well as interaction between multiple organ systems [80]. The main components of an organ-on-a-chip system consist of materials needed to fabricate the chip, a type of biomaterial that will serve as the backbone for the tissue matrix, and mammalian or preferably human cells that will form the parenchyma. Generally, the design of the chip is fabricated using soft lithography, with a replica consisting of PDMS [81]. The choice of biomaterial is influenced by the target tissue or organ system. There exist a variety of biomaterials that can serve as an artificial ECM, which include but are not limited to gelatin, collagen, alginate, fibrin, and hyaluronic acid [81]. Cell lines can consist of primary, immortalized, or pluripotent stem cells (PSCs). Primary cells are ideal for drug-delivery applications as they provide similar physiological function; however, they are more difficult to extract and culture. Organ or multiorgan on chip models are capable of providing different dynamic on-chip 3D cell culture methods for predicting the metabolic activity of drugs upon their interaction with multiple organs in several diseases [82,83]. Multiorgan simulation in a microfluidic platform requires precise recapitulation of parameters such as vasculature, immune-compartments, gradients of growth factors, and cytokines [84]. Moreover, biomimetic disease exposure to the modeled organ should be easily applicable through this platform (Fig. 25.2). Applications of organ-on-a-chip systems now include models of several physiological systems and organs. Cardiac, liver, and vascular platforms are some of the systems more widely studied [85]. Cardiac and liver on-a-chip systems are critical to the drug industry, when considering that cardiotoxicity and hepatotoxicity are common and potentially lethal medication side effects. Reproducing the physiological function of cardiac tissue is challenging; however, culturing cardiac cells in ona-chip platforms has become advantageous with the integration of electrical forces that can induce a uniform contractile function. In a study conducted by Agarwal et al. micro-contact printing developed a PDMS-based cantilever for the development of heart muscle-like tissue by which deflection of these cantilevers, termed muscular thin films, during muscle contraction, imitates the diastolic and systolic stresses generated by the engineered tissues [86]. When screening most cardiovascular drugs, it is important to consider the role that both the liver and the heart play. As such, a liver
heart on-a-chip model was developed by Vunjak-Novakovic et al. [87]. They used PSCs to form liver
heart microtissue and applied physical and chemical cues to differentiate them to cardiomyocytes, endothelial cells, and hepatocytes as a complex multiple organ model.

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Figure 25.2 Liver on-a-chip microfluidic platform that recapitulates a liver lobule-like microtissue. Highlighted in this figure are the (A) schematic design, (B) superior, and (C) lateral view of the microfluidic platform. Source: Reprinted with permission Ma C, Zhao L, Zhou EM, et al. On-chip construction of liver lobule-like microtissue and its application for adverse drug reaction assay. Anal Chem 2016;88(3):1719
27 [191].

Liver physiology is also challenging to replicate, because proteins produced by hepatocytes and other endogenous hepatic cells will rapidly degrade if the physiological environment is not appropriate. However, recent advances have led to liver platforms, such as LiverChip, that can maintain hepatic cells for over 1 month [88]. Brain is a complex organ that contains numerous cell types, including neurons

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(primary effectors of synapses), oligodendrocytes, astrocytes, and microglia, which are critical to consider when developing microfluidic brain modeling systems [89,90]. To this regard, in order to investigate the role of microglia in the progression of Alzheimer’s disease, a novel microfludic chip was developed to investigate the effect of microglia chemotaxis on Aβ plaque formation [91]. Vascular platforms are important for the study of hemodynamics and particle uptake. Several models have been explored that study bifurcations [92], shear stress [93], and particle functionalization [94]. Finally, several studies have shown the possibility of a “humanon-a-chip,” which aims to recreate the physiologic organ
organ interaction [95,96]. Such platforms can provide for more accurate drug testing and cancer screening studies and are the next step to the future of microfluidic designs.

25.1.3 Emerging biofabrication technologies 3D bioprinting is one of the recent emerging and most promising biomedical technologies. Its premise lies in the ability to construct 3D models of target tissue constructs, which can be cellularized during the printing process, and yield functional living tissue that can be further used for in vivo testing or implantation. The main formats of printing can be classified as direct or additive manufacturing, which involves printing in a layer-by-layer format, and indirect or subtractive manufacturing, which involves printing an initial mold that will be successively removed to generate the final construct of interest [97,98]. There are three overall steps to a 3D bioprinting process: generation of a 3D computer
assisted model that will guide the printing process, printing using the biomaterial and cells of choice, and maturation of the cells within the construct [97]. There are several types of biomaterials that can be used as bioinks, such as alginate, gelatin, and collagen; the ideal biomaterial must provide long-term stability while also providing the appropriate environment for cell viability. The bioprinting technologies are generally categorized to three groups, namely, microextrusion, inkjet printing, and laser guided printing. Microextrusion uses various mechanisms of pneumatic [99], piston [100], or screw [101] for flexible handling and dispensing of biopolymers with different viscosities in a final 3D tissue. This method benefits from different material characteristics, including photocurable, thermoplastic, or shear-thinning ones, for continuous printing [102
104]. Grolman et al. developed a coextrusion system to create a metastatic breast cancer model using MDA-MB invasive tumor cells encapsulated in a shell of alginate fibers and a core of mouse macrophages [105]. They showed the active molecular interaction of the cancer cells with macrophage cells, especially highlighting macrophage migration toward the tumor side via an extravasation mechanism. The inkjet bioprinting method uses applied pressure pulses on the fluidic chamber, creating cell-laden droplets that are positioned and patterned on the substrate [106,107]. 3D contractile cardiac hybrids that mimicked half hearts were generated using an inkjet printer and cardiomyocyte-embedded alginate ink [108]. In another study, inkjet printing was used to create a potential model of a blood vessel with complex geometry [109]. In this model a zigzag cellular tube was formed using

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alginate droplets encapsulated with 3T3 fibroblasts cell, showing a viability of more than 82%. Laser-assisted bioprinting (LAB) enables the formation of 3D biological constructs without dependence on a hydrogel but rather operates on a laserinduced forward transferring of cells and biomaterials in a liquid or solid phase with microscale resolution [89,106]. In one approach, LAB was used in developing a skin tissue model by which collagen-based bioink including keratinocytes or fibroblasts cells were printed in a 3D layer format to create biosimilar skin tissue in vitro [110]. LAB has also been used for hard tissue modeling, such as bone tissue. In a study conducted by Catros et al., quartz ribbon coated with a thin absorbing layer of titanium and a layer of bioink containing osteo-progenitor cells were used for bioprinting a model of bone tissue [111]. Cardiovascular, hepatic, cancer, musculoskeletal, and skin models have been some of the main applications of 3D bioprinting technologies. The focus of cardiovascular applications have been engineering heart valves [112] and cardiac tissue patches [113]. While none of these technologies have been tested on humans yet, cardiac tissue patches have shown promise in experiments with rats. Studies with liver tissue bioprinting have focused on drug testing and screening. The ability to print hepatocytes with lobular units that are similar in structure to human hepatocytes has yielded 3D models that metabolize or respond to drug effects very similarly to physiologic levels [114,115]. The target of skeletal studies has been the construct of inner ear, bone, or nose for implantation in humans. Several models of complex cartilaginous structures have been demonstrated, including the production of a bioprinted ear integrated with electronic devices that was shown to capture auditory signals[116]; another model was able to integrate cartilage and adipose tissue formation in an ear model [117]. Advances have been achieved with cancer and skin models as well, as shown with the construction of tumor microenvironments or tumor invasion models [97], or vascularized skin with epidermal and dermal components [118]. Textile bioengineering is a discipline of tissue engineering that has emerged in the recent years and which includes technologies such as electrospinning, embroidering, weaving, knitting, and braiding [119]. Textile tissue engineering enables precise control over the geometry and mechanical properties of the construct, while maintaining control over cell viability and distribution. The structural unit of a biotextile is the biofunctional fiber, which can be synthetic, hydrogel, derived from natural tissue ECM, or hybrid. In the past the main applications of this technology were surgical meshes used in pelvic and abdominal surgical procedures to assist in the closure of hernias or pelvic floor defects [120]. In the recent years the technology has expanded to include studies in cardiovascular tissue engineering with biotextiles approved by the US Food and Drug Administration (FDA) such as CorCap CSD [121], a surgical mesh implanted around the cardiac muscle to provide structural support, and musculoskeletal tissue engineering, with studies showing the capability to produce constructs with tendon-like mechanical properties. Other applications of the technology include studies on neural regeneration of peripheral and sciatic nerves [120], improved wound healing through wound dressings [122], as well as studies in which biotextiles are integrated with biosensors to build

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wearable electronics that can be further utilized to measure various physical parameters, pressure points, or even provide a form of portable power system [123]. Woven looms have been used for controlling cellular patterns in some relative applications, such as muscle tissue formation [124]. Moreover, composite hydrogelbased fibers with encapsulated cells can be used in a weaving process without damaging the cells, making this suitable for 3D organ or tissue model development. Full organ weaving is an emerging and attractive technology through which artificial tissue or organs can be developed by using several cell-laden biothreads assembled in a 3D hierarchical manner [64].

25.1.4 Stem cell technology—biomaterial interface Stem cell technology has revolutionized clinical medicine in the past decade. The ability to manipulate ESCs, MSCs, and more recently induced-PSCs (iPSCs) has created an opportunity to study the pathophysiology of many inflammatory, oncological and congenital disorders, as well as explore a myriad of therapeutic avenues. Each type of stem cell line provides their own advantage, and the choice will be dependent on the nature of the study. For example, MSCs are easily accessible, self-renewable, and can grow in vitro for several generations with high genomic stability, which makes them ideal of tissue regeneration applications. iPSCs do not have the genomic stability of MSCs; however, the ease of generation of these cells from patient-derived somatic cells makes them ideal for personalized stem cell therapies. In vitro manipulation of growth and induction of the different types of stem cells has been successful due to the presence of robust protocols. However, differentiation of stem cells, especially iPSCs, is often incomplete and results in a mixed cell population in vitro [125]. The differentiation capabilities of these cells could potentially be limited by the 2D environment of the culture dish. This issue can be solved by biomaterial engineering, which can provide the stem cell with a 3D extracellular environment and tunable mechanical properties. Several studies have shown that biomaterial stiffness can influence stem cell differentiation. For example, in murine ESCs, low stiffness will favor differentiation toward endoderm, while high stiffness will favor differentiation toward mesoderm [126]. Moreover, MSCs tend to differentiate at environments of higher stiffness, while ESCs prefer environments of very low stiffness [127]. The interaction between biomaterials and stem cells has been explored for studies in several organ systems, including cardiovascular, neural, and musculoskeletal. Cardiac engineering studies have shown that when combined with certain biomaterials, such as ECM protein
based matrices or decellularized matrices, stem cells showed specific differentiation as measured by increased expression of cardiacrelated genes [128], as well as better organization into sarcomeres and enhanced maturation [129]. Neural studies have focused on the enhancement of neural stem cell function with the addition of biomaterial scaffolds to induce regeneration after stroke recovery [130] or spinal cord injury [131]. Musculoskeletal approaches, mainly focused on cartilage and bone regeneration, have taken advantage of the

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mechano-responsive nature of stem cells to improve their differentiation time and specificity into osteocytes or chondrocytes [132,133]. Advances in stem cell technology can complement current biomaterial engineering progress to move tissue engineering a step closer to clinical practice.

25.1.5 Organoids Organoids are 3D structures that self-organize or are directed to organize under organogenesis cues and that resemble in vivo organs in cellular organization and composition [134]. Applications of organoids have spanned many organ and physiological systems, with neural and intestinal models being the most studied. Using organoids, studies have been able to replicate intestinal epithelium with crypts that closely resembled human physiology [135] and cerebral organoids that were maintained in culture for several months [136]. Other examples of the myriad of models developed through this technology include functional gastric fundus organoids [137], thymus organoids that showed thymopoiesis and humoral responses [138], and vascularized cardiac human organoids that were capable of beating [139], amongst others. In generating multicellular organoids (spheroids) as a 3D tissue model, a three-step spontaneous self-assembly cell aggregation process occurs: (1) fibronectin ECM secretion and interaction between the ECM and cell membrane integrin to stimulate primary cell attachment, (2) cadherin expression upon cell fusion, and (3) interaction of type-1 transmembrane proteins (E-cadherins) to initiate strong cell compaction [140,141]. When developing a specific organ model using multicellular organoids, metabolic activity of the encompassed cells should be considered, as reports have shown that organoids with a diameter of more than 250 μm suffer from lack of mass and oxygen transport, leading to decreased metabolic activity. The method of organoid production can also influence high throughput drugtesting assays. When compared to conventional organoid forming methods, including nonadherent surfaces (liquid overlay and hanging drop) [142,143], spinner flask [144], and external forces [145], microtechnology-based methods such as microwell or microfluidic platforms have the advantage of higher cell viability and better programmability [146,147]. Another advantage of the microwell-based organoid development method is their ease of manipulation after spheroid formation, such as encapsulation in ECM-related hydrogels for further model development [148,149]. In addition, for clusters that require ECM adhesion for further condensation, microwells can be directly replicated into biopolymers, such as collagen, and subsequently overlaid with additional collagen to create fully embedded spheroids [150]. Microwell technology in the form of parallel arrays can be also used for highthroughput drug screening and real-time live imaging of generated organoids [151]. In a study conducted by Akbari et al. a self-filling microwell array system was developed using 3D-printed molds and a nonadherent agarose hydrogel. They utilized this platform for tumor spheroid generation from MCF-7 and U-87 cell lines as a representative of breast cancer and brain tumor models [152].

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The ability to integrate stem cells, genetic modifications, and pathogens into organoids has helped model several diseases for drug development including bacterial infections and cancer. For example, organoids have been developed to study models of salmonella infection [153], Alzheimer’s disease [154], as well as Zika virus
induced microcephaly [155]. Recently, tumor models are being explored using organoids, because of the advantage this technology provides not only in the recapitulation of the spatial geometry and microenvironment necessary for tumor development but also because integration of patient specific tumor cells or tissue can lead the way toward personalized medicine. To this regard, advanced tumor organoid models have been developed, especially for colorectal [156] and prostate cancer [157]. Multicellular tumor spheroids with diameter of more than 400 μm are a powerful model with similar native tumor microenvironment, comprising a necrotic core and proliferating outer shell, which can be used for high-throughput chemotherapeutic drug testing. Other organ specific organoids made of either PSCs or organ-specific adult stem cells have been extensively used for different types of organ modeling [158,159]. In one study, salivary gland
derived stem cells (SGSCs) were cultured in polycaprolactone nanofibrous microwells and showed that SGSCs organize into salivary structures under biomimetic 3D culture conditions without niche factors [160]. To model the effect of nicotine as a neurotoxin on neural cell dysfunction, prenatal nicotine exposure was applied using human brain organoid-on-a-chip [161]. Research studies have demonstrated that integrating biomaterials into organoids can improve cellular organization by providing the appropriate microenvironment based on the tissue of interest, which can help further approximate the function and physiology of an organoid to that of human physiology. The use of biomaterials can recreate the physiological architecture and mechanical properties that are critical for embryological development, stem cell differentiation, and cellular organization within the tissue. For example, a study by DiMarco et al. observed that mouse intestinal organoids engineered in collagen I matrix showed the ability to contract, a vital gastrointestinal function which had not been reported on previous intestinal organoid models [162]. Biomechanical stiffness and elasticity, porosity, as well as type of biomaterial are some of the parameters that have proved critical in organoid research. In neural organoid studies, softer hydrogels promoted a neuronal lineage, while stiffer hydrogels promoted differentiation toward glial cells [134]. Meanwhile, another study in lymphoid organoids demonstrated that T cell migration was dependent on connecting pore size. Although Matrigel has been the biomaterial of choice for organoid models, different types of biomaterials have been explored based on the specific application. For example, hyaluronan and laminin are known to promote neural cell growth and differentiation. Advances in organoid cultures, combined with advances in biomaterial and genetic engineering, are rapidly shifting organoid technology toward personalized and precision medicine.

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25.1.6 Rationale design of biomaterials for disease modeling The popularization of human ESCs (hESCs) and the emergence of human iPSCs have opened up new avenues for native-like disease modeling [163,164]. Being able to differentiate into various types of cells, iPSCs and hESCs have been widely used to study monogenic and polygenic diseases in various 2D monolayer models in a dish [165]. However, unfortunately, the “disease in a dish” models do not fully recapitulate the native-like functions of natural tissues. For instance, one of the biggest issues with these simplified disease models is intimately linked to the fact that native-like cellular functions deviate substantially in 2D compared to 3D microenvironments [166]. Indeed, one of the unique features of human organs is their complex and multicellular 3D architecture, and accordingly, a 3D matrix is essential to enable the manufacture of fully functional tissue models. Importantly, a great variety of disease phenotypes are not limited to the cellular level and appear in the tissue and organ levels instead. Therefore the simulation of these disorders is impossible in simple 2D tissue models. In addition, previous studies have failed to find a proper in vitro method to study epigenetic- and age-related diseases [167]. Even though such limitations can be overcome by using various animal-based disease models, the clinical approaches are haunted by some ethical, cost-effectiveness, and cost-efficiency challenges. Moreover, most of the results gained from such animal models are not competent enough to go all the way into clinical trials because of species-specific differences [168]. To conquer some of these challenges, “humanized” mice have recently been employed through genetic modifications; yet, they do not reliably reproduce the human phenotype. To remedy the abovementioned challenges, numerous biomaterials science strategies have been used to develop complex 3D tissue models for in vitro disease modeling [169]. Among the forerunners in this regard are 3D organoids generated through the magical union of various stem cells and 3D scaffolding materials. Notably, some recent studies show that hydrogel-based materials can drive such stem cells into nativelike tissues [170,171]. Hydrogels are water-swollen structures that mimic the ECM of native tissues. The advantages of using hydrogels stem from their dynamic and modifiable structure. These characteristics give rise to promising opportunities such as supporting cell
cell and cell
ECM signals, and efficient nutrient and metabolic waste material exchange with the surrounding environment [172]. Besides utilizing hydrogelsbased structures to encapsulate cells and organoids, employing microfluidic on-chip technology and 3D bioprinting can boost the maturation of organoids even more. In parallel, applying a coculture technology in each of the abovementioned systems may result in even more native-like tissue models capable of also recapitulating different polygenic diseases (Fig. 25.3). The current major challenge is related to how one can incorporate such seemingly disconnected technologies—as the abovementioned ones—into a potent toolbox for carving out complex tissue models in the laboratory [98,175]. This chapter provides an overview of different biomaterials, which could be utilized to give rise to such systems. We have attempted to especially highlight their relative physical characteristics, including their

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Figure 25.3 Applications and functionalization of conventional organoids using different bioengineering strategies such as 3D printing technology, coculture system, on-chip technology, and functional biomaterials. 3D, Three-dimensional. Source: Reproduced with permission from Kim S, Cho AN, Min S, Kim S, Cho SW. Organoids for advanced therapeutics and disease models. Adv Ther 2018:1800087 [195].

biocompatibility, biodegradability, electrical properties, mechanical properties, as well as their vasculature inducing capacity.

25.1.7 Biocompatibility Over the years, natural and synthetic biomaterials have been used to drive stem cells into semimature tissues. On the positive side, it is easier to modify synthetic biomaterials as compared to naturally derived biomaterials, whilst on the negative side cell adhesion is generally weak and immune responses are often inevitable with such material types [176]. To this end, some exciting studies have shown that synthetic materials can easily be modified with various proteins and peptides to improve their biocompatibility and remedy possible foreign body responses [177]. Along these lines an interesting study was undertaken by Gjorevski et al. on a hydrogel-based composite consisting of an FDA approved polymer, PEG, and four different peptide sequences capable of improving bioactivity by triggering cellular contractility, adhesion, and cell colony formation [170]. Small peptides consisting of the RGD sequence have also been widely employed to enhance the bioactivity of different types of biomaterials, as this sequence can specifically bind to integrins α5β1, αiibβ3, and αvβ3, thereby facilitating cell spreading, differentiation, and proliferation [170,178]. In a likewise manner, natural biomaterials such as Matrigel [179], collagen [180], gelatin [181], and elastin [182] are commonly used as a native-like matrix or coating to improve biocompatibility. Although natural biomaterials are less toxic

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Figure 25.4 Different types and characteristics of biomaterials that have been used in 3D composites for disease modeling applications. 3D, Three-dimensional. Source: Reproduced with permission from Dye BR, Kasputis T, Spence JR, Shea LD. Take a deep breath and digest the material: organoids and biomaterials of the respiratory and digestive systems. MRS Commun 2017;7(3):502
14 [196].

than synthetic biomaterials, they can activate immunogenic reactions due to batchto-batch varieties [183]. An example of this is Matrigel (natural basement membrane matrix), which contains different types of critical growth factors, as well as different structural and binding proteins. For these reasons, Matrigel has been used to make a native-like environment for a number of different tumor models, some of which include breast [184], prostate [185], colorectal [186], and lung [187] tumors. However, the drawback is that variations in the concentration of proteins and other components that may exist in Matrigel-based hydrogels of different sources can significantly compromise the reproducibility of such systems [188,189]. Semisynthetic microenvironments are a useful alternative that can be applied to reduce the abovementioned obstacles typically associated with natural-based polymers. A notable example in this regard is GelMA, which brings in the best advantages of the realm of natural and synthetic materials. Indeed, GelMA has been utilized as a 3D matrix for developing different disease models (e.g., cardiac and neurological disorders) and some cancer models (e.g., ovarian and breast cancer) [177]. Similarly, the combination of collagen type I and GelMA not only promoted biocompatibility and angiogenesis but also enabled 3D printing of highly functional tissue models [190]. Up to now, different types of biomaterials have been used to build native-like microenvironments, each of which has their own unique

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characteristics (Fig. 25.4). Generally speaking, hybrid and chemically modified composite biomaterials are an interesting class of materials, as they are capable of increasing biocompatibility, bioactivity, and ultimately the complexity of the finalized tissue constructs.

25.1.8 Biodegradability When designing a proper 3D matrix for in vitro disease modeling, it is important to recapitulate the complex and highly combinatorial microenvironment of native tissues to enable disease models that cover all the necessary biological aspects. The native ECM microenvironment is also a dynamic one, in which new tissue matrices constantly replace old ones through various ECM remodeling pathways. Therefore one of the essential requirements of ECM-like biomaterials made in the laboratory is biodegradability, as this enables a fast ECM turnover, something that is needed for both optimal tissue development and disease modeling [192]. Moreover, the lack of degradability can be a limitation when it comes down to tissue vascularization [193]. However, the rapid degradation rate of some synthetic-based scaffolds can promote inflammatory reactions by creating a local acidic environment, which can be lethal for some cell types [194]. Therefore the engineering of biomaterials with a tunable degradation profile is a vital issue in the field of tissue engineering and disease modeling. As Gjorevski et al. suggested, the desired degree of degradability is intimately linked to the targeted biological process, such as cell maintenance, proliferation, or differentiation. Specifically, they figured out that the growth and differentiation of intestinal stem cells were significantly improved within softer and more degradable hydrogels, while nondegradable and stiffer structures were required to maintain the cells in an undifferentiated state. Surprisingly, using a degradable and stiff matrix enhanced colony formation and cell expansion but caused the cells to become irregular in shape and shifted them into a more inflammatory-like state. To remedy this issue a mechanically stable PEG and a hydrolytically degradable PEG was partially used. It was shown that the different ratio of these two combinations along with laminin-111 could be applied to manipulate both the matrix stiffness and its longterm degradation profile (Fig. 25.5) [170]. Nondegradable and stiff matrixes also typically compromise the vascularization process, which is not desirable, as tissue vascularization is one of the fundamental goals of 3D disease modeling (Figs. 25.4
25.6).

25.1.9 Vascularity The absence of vasculature in disease models such as intestine, kidney, liver, and lung is a critical challenge that needs to be addressed. The reason is that such tissues and organs provide complex functions wherein many different cells are linked with other cells in the surrounding environment through chemokine/ cytokine/growth factor signaling. As such, the presence of vessels as signaling connectors between cells is one of the important hallmarks of native-like tissue models

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Figure 25.5 Defined hydrogels as a 3D matrix for intestinal organoid formation. (A) A schematic presentation of the chemical mechanism of hydrogel formation. (B) Mechanical characterization of hydrogels. (C) and (D) Organoid formed within hydrogels. (E) A schematic illustration of the suggested mechanisms affecting different stages of organoid formation within the native-like 3D microenvironment. 3D, Three-dimensional. Source: Reproduced with permission from Gjorevski N, Sachs N, Manfrin A, et al. Designer matrices for intestinal stem cell and organoid culture. Nature 2016;539(7630):560, 170.

[177]. On the other hand, the presence of neovascularization in cancer models can trigger uncontrolled cancer growth [198]. Also, hypervascularization in some chronic conditions such as skin- and ocular diseases can give rise to some dysfunctionalities. Therefore vascularization and angiogenesis are essential features that need to be factored in when studying such disease models [199]. Generally speaking, for proper vascularization, it is important to tap into various coculturing methodologies to obtain the appropriate cell-to-cell communication and coordination to give rise to native-like vasculature [177]. As mentioned before, various 3D biomaterials can entrap cells and facilitate cell
cell connections by the formation of different types of junctions [200]. Accordingly, different materials

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Figure 25.6 The effect of the size of pores on cell attachment and contractility in 3D Scaffold. (A) Confocal image of MSCs shows differences between the spreading behavior of the cells within a porous scaffold with large and small pores. (B) Regulation of cell attachment and forces by changing the scaffold porosity. 3D, Three-dimensional; MSCs, mesenchymal stem cells. Source: Reproduced with permission from Reilly GC, Engler AJ. Intrinsic extracellular matrix properties regulate stem cell differentiation. J Biomech 2010;43(1):55
62, 197.

have been used to manufacture functional and vascularized 3D model organs. These materials have typically been based on natural biomaterials (e.g., gelatin [201], agarose [202], and Matrigel [202,203]) and synthetic materials [e.g., poly (glycolic acid) and poly-L-lactic acid [204]] in different forms. Moreover, hybrid scaffolds (composed of natural and synthetic materials) play a considerable role in this direction, as they can be more tunable in terms of their degradability and bioactivity (e.g., a combination of PLGA and silk) [205]. Since one of the important bottlenecks of cancer treatment is the ability to inhibit cancer-induced blood vessels, a number of 3D tumor models have also been applied over the years to study the angiogenesis inhibitory effect of new drugs. Because of their interconnected pores and “easy-to-functionalize” nature, hydrogel-based scaffolds can facilitate vascularized cancer microenvironments. For this reason, they are increasingly being used in the field to deliver native-like tumor models for screening the antiangiogenic effect of new drugs. In this regard the range of parameters, including hydrogel stiffness, degradability, porosity, and permeability, are important and need to be fine-tuned accordingly to develop the needed vascularized structures [206].

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25.1.10 Mechanical properties One of the important hallmarks of cellular function is intimately linked to the mechanical properties of the ECM, since the rigidity and stiffness of the surrounding environment can influence the fate, maintenance, and migration of cells [170,207]. Indeed, it is well-established that different tissues in different organs present variable stiffness ranging from B1 kPa to B20 GPa [208]. Interestingly, experiments have shown that multicellular differentiation in intestinal organoids requires softer gels (190 Pa), while this stiffness fails to support cells for a long-term period; instead, 1.9 kPa stiffness is necessary for cell maintenance [170]. Moreover, as previously reported, the specific phenotype of human MSCs (hMSCs) is affected by the rigidity of the surrounding environment, with highly stiff and covalently cross-linked hydrogels giving rise to more rounded cells [209]. In addition, variation in the local hydrogel stiffness can contribute to regulation of hMSC fate and affect their migration. Thus the mechanobiological property of 3D biomaterial matrices is a critical concept that needs to be evaluated carefully to yield the desired tissue models [197,210]. During tissue development, cells pass through different stages, including cell proliferation, differentiation, migration, and self-assembly. Therefore it is essential to have tunable mechanical properties over time, as matrix stiffness is a determining factor in these biological processes. There are several methods to modulate the mechanical properties of hydrogels in a static manner, such as controlling light irradiation for photo-crosslinkable materials or using various degrees of hydrogel cross-linkers [208]. Although these methods are useful, they are not dynamic and do not provide a tunable mechanical property over time, which is why further studies are required in this direction.

25.1.11 Electrical conductivity Electrical excitability and aligned cell morphology are two essential properties of cardiac tissue modeling. Indeed, to provide a coordinated and synchronized beating of cardiac tissue, mixing and coculturing of different types of cells are necessary, but not sufficient, since a stable cell
cell interaction and appropriate cell density are required as well. Coculturing and localized 3D printings of different cells are some of the methods that can provide the needed cell
cell interactions to deliver spontaneous contraction and uniform cardiac tissue beating [177]. To this end a number of recent studies have shown that the inclusion of conductive materials such as gold nanorods, graphene sheets, and carbon nanotubes within hydrogels can stabilize the important electrical signaling between the cells to deliver the promise of coherent and native-like cardiac tissue beating [177]. Some studies have also demonstrated that one can use aligned nano- and microfibers to speed up the maturation and alignment of cardiac cells, in order to generate more native-like models [177]. Several biomaterial-based platforms have been explored for cardiac tissue modeling, including self-assembled PDMS-based biowires surrounded by a collagen gel, which were subjected to electrical stimulation. This system contained human

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PSCs, fibroblasts, endothelial cells, and smooth muscle cells. Although electrophysiological maturation was observed, lack of vascularization in this study was a challenge that needs to be addressed to sufficiently bridge the gap between laboratory and human biology [211]. Differentiation, cell
cell interactions and electrophysiological maturation of neural cells in the brain are also complicated and difficult to reproduce in the laboratory. To address this problem the possibility of layer-by-layer 3D printing of fibrin sheets was investigated, and it was shown that this method enhanced cell affinity to the native-like ECM and led to maturation of primary neurons. Furthermore, thermal inkjet printing of the primary neural cells was suggested in the study to improve the electrophysiological properties of the cell [212]. However, current models for delivering electroactive tissues in the laboratory still suffer from nonnative-like maturation—especially cardiac and neural cells—due to the presence of undifferentiated cells and lack of sufficient complexity. As such, further research is required to explore other types of bioengineering approaches. Specifically, it is important to look into biomaterial systems that can yield much higher bioactivity through a more sustained delivery of growth factors and various differentiation inducers. Progress is also required for the fine-tuning of various coculturing protocols and 3D bioprinting approaches. Finally, efforts need to be placed in optimizing some of the existing protocols that enable the integration of electrical signals, which is a critical component of neural and cardiac tissue physiology. We anticipate that further explorations along such avenues could assist the tissue engineer to remedy the current shortcomings in this vibrant research theme.

25.2

Emerging applications and clinical considerations

25.2.1 Inflammatory response and cancer modeling The tumor microenvironment consists of various cellular and subcellular interactions. Parameters such as complex supportive blood vessels, secreted growth factors and exosomes, oxygen, and other chemical gradients play a major role in tuning the growth, progression, invasion, and metastatic stages of cancer cells [213
215]. Moreover, the role of the specific tumor microenvironment is altered in different stages of tumor growth and progression [216,217]. As such, tumor cells in communication with stromal cells and ECM exhibit self-tuning characteristics [218,219]. During the disease-spreading phase, new signaling pathways are activated in which cancer cells tune their microenvironment to facilitate critical events that include neovascularization and autocrine/paracrine activation of the tumor invasion state [220
222]. Herein, the development of a comprehensive model that recapitulates the major aspects of the tumor microenvironment seems to be inevitable. To this regard, different kinds of microfluidic and tissue engineering approaches have evolved for tumor microenvironment modeling, with the goal of broad spectrum chemotherapeutic drug screening [223
225]. In a study by Liu et al. [226] an integrative microfluidic design was utilized to model a liver tumor on chip by

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investigating the interaction of HepG2 cells with their relevant stromal cells and ECM using an on-a-chip mono and coculture method. Results showed that in the presence of NIH 3T3 fibroblasts cells, carcinoma cells tended to form aggregates toward the position of stromal fibroblasts, while motility and migration of fibroblasts due to soluble signaling interactions was greater than the monoculture mode. In addition, a 3D dense tissue-like model of liver carcinoma was fabricated in a fibronectin-coated PDMS-based microfluidic biochip [227]. HepG2/C3a cells grew in a multilayer structure so that collagen deposition and cytoskeleton formation were similar to the original tumor environment. Fang et al [228] also reported a new hydrogel for pancreatic ECM tumor modeling composed of CFPAC-1 cells encapsulated within transglutaminase-cross-linked collagen, which proved to be advantageous in presenting cell
matrix interaction sites from collagen-derived peptides, geometry-initiated multicellular tumor spheroids, and metabolic gradients in the tumor microenvironment. This model can mimic the natural ECM of pancreatic tumors, including the point of stiffness and other biophysical aspects [229]. Scaffold-based methods have also been used for cancer modeling. To this aim a primary brain tumor modeled by Keivit et al. used a 3D porous chitosan
alginate (CA) scaffold for culturing glioblastoma cell lines, specifically human U-87 MG and U-118 MG glioma cells and rat C6 glioma cells. Their results indicated that U-87 MG and U-118 MG cells exhibited notably higher malignancy when cultured in CA scaffolds. CA scaffolds can thus serve as a more effective platform for development and study of anticancer therapeutics [230]. Wang et al. introduced a bioengineered 3D brain tumor model to help elucidate the effects of matrix stiffness on U87 cells using PEG-based hydrogels with brain-mimicking biochemical and mechanical properties. Their results showed that changes in matrix stiffness induced differential GBM cell proliferation, morphology, and migration in the 3D hydrogel. Increasing matrix stiffness led to delayed U87 cell proliferation inside hydrogels, but cells formed denser spheroids with extended cell protrusions. They suggested that varying matrix stiffness can induce differential ECM deposition and remodeling by employing different HA synthases or MMPs. Furthermore, increasing matrix stiffness led to simultaneous upregulation of Hras, RhoA, and ROCK1, suggesting a potential link between mechano-sensing pathways and the observed differential cell responses to the changes in matrix stiffness [231]. To screen chemotherapeutic drugs on the tumor models, Zhang et al. developed a microdevice to perform chemotherapy resistance analysis in SPCA1, a lung cancer cell line. Their device includes a PDMS chip with a simple external small clip that served as a microvalve to control the fluid flow, enabling a parallel control experiment to be carried out simultaneously, and a syringe pump, which supplied the cells with fresh medium, mimicking in vivo microenvironment. Their results showed that the cells could grow and spread well for at least 3 days. The expression of P-gp and GST-p was downregulated by adding their corresponding inhibitors. The percentage of apoptotic cells for P-gp inhibition group increased 2.9-fold compared with that of the control group [232].

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25.2.2 Cardiovascular diseases Cardiovascular disease remains one of the main causes of death in the United States and worldwide. The study and treatment of conditions such as myocardial ischemia, cardiomyopathies, myocardial fibrosis, and drug-induced cardiotoxicity has been limited due to the difficulty of recapitulation of the precise cardiac physiology and muscle architecture both in vitro and in vivo. Despite the challenges, several technologies have taken aim at recreating cardiac anatomy and physiology through microfluidics, organoids, 3D bioprinting, and stem cell technology. Microvascularization of a functional myocardial unit is an issue that has been tackled by microfluidics approaches. Zhang et al., by using a 3D microporous scaffold, were able to build an on-a-chip platform that consisted of vascularized myocardial tissue capable of contraction and peripheral perfusion [233]. These results hold premise especially for the study of drug-induced cardiotoxicity. Simulation of electrophysiology is another challenge that has been addressed through on-a-chip models. Electric signals have been integrated in microfluidic platforms by the use of 3D electrodes, and, more recently, increased conductivity through tissue has been achieved through nanotechnology, including carbon nanotubes [234] and gold nanoparticles [235]. Adding an electric stimulus to these platforms has shown to increase the quality of engineered cardiac tissue, as shown by improved control over cellular orientation, growth, and maturation [233]. The intricate anatomy of the cardiac tissue is perhaps the main challenge of cardiac engineering research. When recreating the architecture of the heart, one must account for all the critical components of this organ—muscle wall (composed of epicardium, myocardium, and endocardium), the atria and ventricles, the atrioventricular (AV) and semilunar valves, and the sinoatrial (SA) and AV nodes. This can be achieved by integrating several cell types, including cardiomyocytes, smooth muscle cells, endothelial cells, and fibroblasts. Several technologies have taken aim at recreating functional cardiac tissue for study or treatment purposes. Stem cells have been used with biomaterial scaffolds to create cardiac patches, which, in a study by Kraehenbuehl et al., improved the function of infarcted myocardium [236]. As previously mentioned, using biomaterial scaffolds also improves differentiation of stem cells to cardiac progenitors. Moreover, a recent study integrated electronics within these cardiac patches, with the goal of monitoring and regulating its function [237]. 3D bioprinting has made progress especially in the development of heart valves, as several studies have reported the synthesis of trileaflet valves that appear structurally similar to that of humans [97]. Although these valves have not been tested on humans yet, the results are encouraging, considering the severity and high prevalence of valvular disease. Moreover, a functional 3D bioprinted valve, potentially using a patient’s own cells, would eliminate the need for anticoagulation that is required with prosthetic valve replacements. Reconstructing a whole heart is proving challenging; however, the technology is rapidly advancing. Cardiac patches are being printed with stem cells that show vascularization and can be implanted in rats [238,239]. Moreover, a study from Hinton et al. was able to reconstruct the

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trabecular structure of a whole heart through 3D bioprinting [240]. Finally, cardiac organoids are another technology that has attempted to mimic cardiac physiology and embryology. Organoids have been used to study cardiotoxicity [241] and metabolic pathways [242].

25.2.3 Skin diseases The principal functions of the skin are protection and insulation. Skin is the main barrier that protects the human body from the environment, including pathogens and UV rays. Skin conditions affect a majority of the population, whether it is acne, infections, wounds, burns, or skin cancer. Wounds and burns are of interest to biomaterial engineering. The majority of skin wounds are generally appropriately addressed through the body’s own healing process. However, severe wounds, such as those that are deeply penetrated, infected, ulcerated, or are due to burns, will require more careful medical attention. Chronic wounds are of particular interest, as rising numbers of obesity, type 2 diabetes, and peripheral vascular disease have led to increased prevalence of chronic and complicated wounds. Burns are also of interest for tissue engineering applications, as there is yet a treatment option that can prevent scarring and significantly restore skin architecture following second- or third-degree burns. Moreover, infection with antibiotic-resistant organisms, such as MRSA, or difficult to treat organisms, such as Pseudomonas aeruginosa, are a significant problem with burns and chronic wounds. There are several tissue engineering developments that are addressing these challenges. Significant progress has been achieved in skin engineering using engineered scaffolds that serve as skin substitutes. Bioengineered scaffolds can enhance regeneration of the skin tissue by providing a compatible environment for the native cells to thrive, decrease scarring, and provide temporary protection to an otherwise exposed tissue. Several commercial scaffolds have been developed that serve as dermal substitutes, such as Integra, or as dermal
epidermal substitutes, such as Apligraf, both of which are FDA approved [243]. 3D bioprinting is another technology that is being utilized to engineer skin tissue with a dermal and epidermal layer [244], as well as sweat glands [245]. The 3D architecture and feasibility of combining several types of cell lines can make it possible to integrate important structures such as hair follicles and melanocytes. Moreover, this technology also allows for making the bioprinted structures color compatible to the patient, a characteristic that could not be addressed by other platforms of tissue engineering. In addition, 3D bioprinting is also being used by cosmetic giants such as L’Oreal and Procter & Gamble to develop 3D models of skin tissue for ex vivo testing of cosmetic products [246]. Wound dressings are another platform for wound management that has shown to enhance the healing process. Different strategies have been tested, including hydrogels with encapsulated stem cells, growth factors, nucleic acids, or even antibiotics. Generally, these techniques have proven successful for wound-healing applications, as shown by improved skin closure, nerve regeneration, and vascularization of the healing tissue [122]. For example, Shen et al. used a

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dextran-based hydrogel in a porcine third degree burn model to show improved reinnervation [247]. Furthermore, the utility of hydrogel-based dressings can transcend treatment, to also focus on diagnosis and monitoring of wounds. To this regard, Mirani et al. demonstrated that GelDerm, a 3D bioprinted scaffold, enables pH wound monitoring and antibiotic delivery that can also interface with a smartphone application [248]. These developments can potentially allow the clinician to virtually monitor the patient’s wound status and obtain diagnosis of wound complications in real time, avoiding the development of severe infections or gangrene.

25.2.4 Gastrointestinal diseases Gastrointestinal disease is a major cause of morbidity and mortality worldwide. Liver cirrhosis due to alcohol abuse or nonalcohol fatty liver disease, chronic pancreatitis, autoimmune disease of the intestinal tract (Crohn’s disease and ulcerative colitis) and of the biliary tract (primary biliary cirrhosis and primary sclerosing cholangitis), or disease induced by pathogens (Clostridioides difficile
induced pseudomembranous colitis), are fatal conditions that at times cannot be cured. Moreover, cancers of the GI system, such as esophageal, gastric, pancreatic, and colorectal are also rapidly fatal and with high morbidity, since the removal of the affected organ is required in most cases. Thus addressing the gastrointestinal tract in vitro and in vivo is very important, not only to study the pathophysiology of the system and the conditions addressed above but also to address treatment options, including the potential for regenerating or replacing the diseased organs. Study of the hepatobiliary system from a biomaterials perspective has focused mainly on addressing liver toxicity. Many of the drugs that are available in the market, including antibiotics and antivirals, immune suppressors, acetaminophen, and even a majority of herbal supplements, are metabolized by the liver. As such, hepatotoxicity is a very common side effect, and one of the main profiles that is investigated before a drug is approved by the FDA. In order to build a platform for hepatotoxicity testing that would closely represent that of human physiology, one must recreate the hepatic lobule, the functional unit of the liver that is composed not only of hepatocytes but also includes Kupffer and sinusoidal cells. Approaches to recreate this physiology have been attempted using liver on-a-chip technologies, with models that have tested hepatotoxicity in cells for more than 30 days [88]. 3D spheroids models, which are also available commercially as RegeneTOX and GravityTRAP, have shown in studies the capability to metabolize and respond to hepatotoxic drugs in a dose-dependent manner [249]. Advances in recapitulating liver anatomy have been achieved using 3D bioprinting. Nguyen et al. was able to bioprint primary human hepatocytes and nonparenchymal cells in a defined architecture and showed distinct organization of the hepatocytes, production of albumin and ATP for over 4 weeks, as well as successfully compared hepatotoxicity profiles between structurally similar toxic and nontoxic medications [250]. The intestinal tract spans over 8 m and is composed of the small and large intestines, two structures continuous of each other, however with distinct function and

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physiology. The small intestine is important for nutrient absorption, metabolism of oral drugs, and is commonly affected by pathogens and drug toxicity. As such, study of the small intestine has focused on the physiology of the crypts, which are invaginations of the epithelial layer that contain stem cells, and thus are the keys to the regenerative capabilities of the small intestine. Functional crypt models have been developed using organoid technology that closely resemble normal physiology [135], while studies using murine models have assessed nutrient transport and protein secretion [251]. A study by Wang et al also developed a crypt-villus architecture in vitro using patient-isolated jejunum cells from gastric bypass surgery in a micropatterned collagen scaffold, which showed a distinct stem cell layer [252]. Large intestine studies have focused on colon cancer, as the third most common cancer diagnosed in the United States alone. For example, an organoid biobank was developed from 20 patients with colorectal cancer, to assess for genetical differences between these colorectal patients [253]. Finally, construction of intestine has been attempted for regenerative purposes, such as to develop patches to repair injured small intestine. However, the main challenge remains the replication of the alignment and contractile function of the muscularis propria, the smooth muscle layer of the intestinal wall [254]. Despite the challenges, research advances in gut engineering are promising for future clinical applications.

25.2.5 Neurological disorders The nervous system is the most complex system of the human body. Neurological disorders, including cerebrovascular disease and Alzheimer’s disease, affect a significant part of the population. These conditions are generally of very high morbidity as well, mainly because regeneration of the nervous system, whether central nervous system (CNS) or peripheral nervous system, is very slow and often complete recovery is not possible. Study of the CNS has focused mainly on cerebrovascular disease and traumatic spinal cord injuries. Stroke, both hemorrhagic and ischemic, is a leading cause of death worldwide. Traumatic spinal cord injury, while not as prevalent, is a devastating condition that greatly impairs the quality of life. The treatment options for both stroke and spinal cord injury are minimal. For stroke patients, treatment focuses on physical therapy and prevention of another episode. Patients with traumatic spinal cord injury are treated with steroids, as decreased inflammatory responses will allow for better healing. However, the effectiveness of these treatment options is low, and most of the patients will have minimal to no recovery. One of the main obstacles of the CNS’s regenerative capabilities is the inflammatory reaction that occurs during brain or spinal cord injury, eventually leading to the formation of the glial scar, which consists mainly of astrocytes. Transplantation of neural stem cells in patients with stroke or spinal cord injury has been attempted to improve regeneration of damaged neural tissue, however low cell viability and migration of transplanted cells away from the injury site has been an issue [255]. Biomaterials have been utilized to improve the technology. For example, Ballios et al. integrated neural progenitor stem cells within a

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hyaluronan-based hydrogel for transplantation in mice models of stroke, showing improved viability and better tissue penetration of stem cells [256]. Biomaterial scaffolds have been used along with stem cells for regenerating spinal cord injuries [257], the study which showed sustained stem cell viability for over 8 weeks within the scaffold and increased axonal regeneration. Manipulating the inflammatory process at the injury site is another strategy that can be considered in order to prevent the formation of a glial scar. To this regard, Li et al. used nanoparticles to induce conversion of astrocytes into neural cells and oligodendrocytes in a spheroid model [258]. Various strategies with biomaterial scaffolds have been studied for constructing artificial nerve conduits for enhanced axonal regeneration in the peripheral nervous system. Parameters such as luminal cues, type of biomaterial, fibers, and electrospun tubes have been investigated, and several studies have shown success in promoting healing and axonal regeneration [259]. Advances in the treatment of neurological disease are promising. Yet, it is of utmost importance to continue investigating brain physiology, as our knowledge of this organ is limited. Constructing 3D models of the brain is challenging, due to the complexity of this organ. Progress to this regard has been achieved using organoid technology. Neural organoid models have been developed that closely mimic human brain development. Such studies have demonstrated distinct organization of progenitor zones that are similar to human physiology [136], cortical neurons expressing markers from all six layers, and recapitulation of the outer subventricular zone [260]. Moreover, organoids have been used to investigate and demonstrate important pathophysiology of neurological disease, especially neurodevelopmental disorders such as microcephaly and Miller
Dieker syndrome. For example, studies using iPSCs-derived from a patient with microcephaly demonstrated premature neuronal cell differentiation as compared to normal patients [136]. It is of note that characterization of biomaterials used in organoids has proven critical for manipulating differentiation of stem cells to lineages of interest, improving the balance between different neural cells and enhancing the architecture within the organoid [134]. The current progress in neural tissue engineering is exciting; however, it is clear that the key to understanding such a complicated pathophysiology is continued interdisciplinary research between biomaterial, stem cell, and genetic engineering.

25.3

Conclusion

Understanding of tissue architecture and organ modeling is of utmost importance for the study and treatment of disease. To this regard, research advances in biomaterial engineering, in combination with other technologies, are leading the path in simulating organ level physiology using in vitro, in vivo and 3D modeling methods. In this chapter, we explored how recent advances in microengineered tissues, microfluidics, 3D bioprinting, stem cell, and organoid technologies have enhanced our understanding of tissue pathophysiology, improved recapitulation of 3D organ

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modeling, and organ
organ interactions and influenced new avenues of drug testing and treatment options. Furthermore, we discussed the strategies the bioengineer must take into consideration when constructing tissue, organ, or disease models by analyzing the importance of biocompatibility, biodegradability, vascularity, and electrical conduction in the design process. Finally, we took a system-based approach to highlight the progress made in the study of inflammation and cancer, skin conditions, cardiovascular, gastrointestinal, and neurological disorders. Although further research is required, recent trends in studying organ modeling and in vitro disease hold promise for improving the understanding, prevention, diagnosis, and treatment of a multitude of conditions.

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49 e4.

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Shabir Hassan1,*, Marcel Heinrich2,*, Berivan Cecen1, Jai Prakash2 and Yu Shrike Zhang1 1 Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, United States, 2Section Targeted Therapeutics, Department of Biomaterials Science and Technology, Technical Medical Centre, University of Twente, Enschede, The Netherlands

26.1

Introduction

Over the last decades the domains of life sciences, basic sciences, and engineering converged, enabling the fabrication of structured or three-dimensional (3D) multicellular architectures subjected to dynamic cultures, and lead to the development of microphysiological systems that are capable to replicate human physiology or pathology with great accuracy. In comparison to conventional two-dimensional (2D) monolayer cultures or in vivo models, these systems allow for controlled in vitro replication of organ- or tissue-level functions including cell-to-cell interfaces, multicellular constructs, peripheral tissues, as well as the vascular circulation of the human system [1,2]. Recently, lab-grown human organoids have drawn increasing attention among scientists for their ability to recapitulate human organ functions in an in vitro setting. Organoids are systems that emerge from omnipotent or pluripotent stem cells that are grown in controlled microenvironments to achieve biomimetic 3D architectures [3]. By controlling the differentiation of the stem cells using growth factors, different organoids can be obtained such as but not limited to intestinal or brain tissues [3]. As these tissue analogs are based on the self-assembly of the given cells, these models are excellent tools to study developmental morphogenesis and biomimetic higher order human organ functions. As the stem cells can be obtained from patients themselves, disease models based on organoids are a promising tool for patient-specific drug-screening applications in the future [4]. However, the fact that some organoids such as cerebral organoids fully rely on self-assembly and lack an external guidance resulting in a high level of heterogeneity within different organoids, limits the reproducibility of the system [5,6]. Furthermore, the generation of organoids that display organ-specific properties and are therefore suitable for drugscreening applications can take up to several weeks to months depending on the 

These authors contributed equally.

Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00019-2 © 2020 Elsevier Ltd. All rights reserved.

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desired organ type [4,6
8]. These limitations render organoids less suitable for the rapid high-throughput screening of drugs at this moment. Microfluidic chip-based platforms, or so-called organ-on-chip (OOC) platforms, can overcome some of these limitations of organoids. These systems emulate organ-level physiological or pathphysiological functions in vitro by seamlessly integrating both 3D biomimetic tissue structures and the dynamic microenvironments into chip-based devices (Scheme 26.1). As such, they are effectively used to facilitate drug development and personalized medicine at improved accuracy than possible with conventional in vitro models. Since one of the pioneering OOC platforms replicating alveoli in the lung [9], many platforms have been developed over the years that replicate the smallest functional units of different human tissues and organs, including OOC platforms of the lung [10,11], brain [12,13], heart [14,15], kidney [16,17], liver [18
20], gut [21,22], muscle [23,24], and bone [25
27] up to multiorgan platforms covering more than one tissue in a single system. Such models are often based on cells that are grown in a well-known and defined manner resulting in a high reproducibility of the systems. Furthermore, such models allow for simpler introduction of a pathophysiological situation, making them highly suitable for the screening of drugs on specific cells and interaction within the human body in a higher throughput manner whiling maintaining their high-content characteristics [2]. For instance, lung-on-chip models have been developed to investigate early markers for chronic obstructive pulmonary disease (COPD) [10], brainon-chip models mimicking the development of Alzheimer’s disease [13,28], or liver- and kidney-on-chip models to investigate drug-induced toxicity of different compounds [17,29,30,31].

Scheme 26.1 The OOC platform for emulating human organ-level physiology and pathophysiology in vitro. OOC, Organ-on-chip.

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Design and biomaterial considerations for the development of specialized microphysiological systems

Biomaterials are materials of natural or synthetic origin intended to come in contact with any tissue, blood, or biological fluids for use in diagnostic, therapeutic, or research applications without adversely affecting the tissues or organs. Based on the evolution in the development of biomaterials, they have been divided into four generations. First-generation biomaterials are also referred to as “ad hoc’ or “unplanned,” because most successes were accidental rather than by design. They were introduced in the 1960s and 1970s [32]. Examples included gold fillings, wooden teeth, poly(methyl methacrylate) (PMMA)—dental prostheses, cemented joint replacements, glass eyes, vascular grafts, among others [33]. The highlight of these materials was their bioinertness [34]. Post-1980s, collaborations between physicians and engineers led to the introduction of second-generation biomaterials in the form of engineered implants. These implants included titanium alloy dental and orthopedic implants, heart valves, and pacemakers, among many others [33]. The end of the 20th century and beginning of the 21st century marked the introduction of the third or the “actual” generation of biomaterials with the use of polymers as new and promising materials added to the field of materials that could be transformed for stimulating cellular responses at a molecular level, for example, bioactive glasses and macroporous foams designed to activate genes that stimulate regeneration of living tissues [35]. The last 20 years have led to a tremendous progress in the disciplines involved in biomaterials science. It is very difficult to estimate the duration of the last generation, as well as imagining what the next one will be [32,36]. Regular developments and strides are constantly being made in this field as a growing understanding of cellular and molecular biocompatibility enables different unique advances such as biomimetics for the different applications of biomaterials [37
40]. Today, the applications of biomaterials have traversed beyond different fields of medicine and are extensively used for the development of cardiovascular devices, orthopedic and dental applications, ophthalmologic applications and devices, bioelectrodes and biosensors, burn dressings and skin substitutes, sutures, and drug-delivery applications, etc. With such a vast footprint in medical applications, there are certain selection parameters for these materials to be considered before their respective use [41].

26.3

Selection parameters for biomedical applications

According to their applications, biomaterials may be required to maintain their structural and functional properties over a prolonged period of time [42]. Most of the biodegradable polymers contain unstable, hydrolyzed functional groups (e.g., ester, amide). When the structures containing the biodegradable polymer enter a physiological environment, water and other active agents from the tissues diffuse

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into the matrix and break up the polymer chains by hydrolysis or enzymatic degradation. An ideal biomaterial should be biocompatible, bioactive and should not cause inflammation, toxic reactions, or any allergic symptoms in the body [43]. These properties are dictated by physicochemical properties, such as mechanical strength, rheological, chemical, and biological characteristics [41]. Such properties in turn enhance the generation of tissue constructs with adequate mechanical strength and robustness, while retaining the tissue-matching mechanics, preferably in a tunable manner. In addition, these properties should affect polymerization or gelation and stabilization of the material to aid the fabrication of structures with high shape fidelity. Also, the mechanical properties of the biomaterial should be appropriate so that any scaffold developed from it should not collapse during normal wear and tear conditions of its application. The biomaterials should be easily sterilizable to prevent infection [44]. Especially in bone tissue engineering, there is a tremendous need for porous biomaterial scaffolds to guide vascularization of the developing tissue [45]. For a biomaterial used for a given mechanical application, some of the requirements to be examined include Young’s modulus, ductility, tensile strength, yield strength, compression strength, as well as fatigue and wear properties. These parameters are evaluated by the fact that they have different properties for each tissue in the human body. For example, elasticity is B0.1
1 kPa in brain tissues, but it varies between B30 kPa and above in extremely hard tissues such as bones and teeth [46]. High values for yield strength and compression strength properties prevent breakage and improve functional stability. Ductility is important for modeling biomaterial formation and dental biomaterials. The increase in hardness reduces wear rate and an increase in strength makes fracture more difficult [44]. These properties are used to define the biomaterial processing capability as well as to assess the success and biocompatibility. In addition, biomaterials should possess biocompatibility and, if necessary, biodegradability mimicking the natural microenvironment of the tissues and have the suitability for chemical modifications to meet tissue-specific needs. Since biomaterials are directly exposed to the living tissues, biocompatibility plays a critical role in the success of its proper application. Primarily, it is believed that a material is biologically compatible when its presence does not elicit any response from the biological environment. The properties should be tunable adding to the potential for large-scale production with minimum batchto-batch variations [41]. Similarly, the design and material choice for microphysiological systems are of utmost importance and highly depend on the aimed tissue or organ part. As different organs in the human body consist of a variety of cells, materials, and architectures, every organ replicate requires a unique choice of biomaterial and design. In this chapter, we will specifically discuss the progresses that have been made in biomaterials for the development of on-chip organ models in two sections. In the first section the focus will be on the development of OOC platforms with an emphasis on some of the important platforms that have been introduced in the recent past. We will discuss the biomaterial considerations that need to be kept in mind while developing these OOC platforms and deliberate upon several important on-chip organ systems. The second section will be dedicated to the different biomaterials

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that are being used not only in the development of chips for some of these platforms but also in the encapsulation of cells from different tissues that form the biological component of these systems.

26.3.1 Lung In its full complexity the lung consists of over 40 different cell types and multiple tissue components such as the vascular, muscular, and immune systems as well as the cartilage [47]. However, microphysiological systems seeking to replicate the lung often focus on the smallest functional unit of the lung—the alveoli. Characteristic for alveoli is an air
liquid interface where the alveolar epithelium is exposed to air, and underlying endothelial cells stay in contact with the blood, both separated by a basement membrane [48]. Replicating this air
liquid interface is a crucial design aspect for lung-on-a-chip. In fact, the first lung-on-a-chip platform was designed to simply replicate this interface on a chip platform. Back in 2007 Takayama et al. designed a microfluidic chip based on polydimethylsiloxane (PDMS) that consisted of two channels separated by a porous membrane that could be perfused with medium in one of the channels using an inlet and an outlet [9]. Primary human small airway epithelial cells seeded on the membrane mimicked the functional part of the alveoli. Air flowing through the upper channel and culture medium through the lower channel replicated the characteristic air
liquid interface. The choice of PDMS in this system was based on the optimal mechanical and chemical properties of PDMS and the ease to introduce defined architectures at high resolution in a well-controlled manner. For this reason, PDMS became one of the most used materials in the fabrication of OOC platforms, and PDMS will be mentioned as a base for many other platforms discussed throughout this chapter. Moreover, the elastic properties of PDMS also make it highly suitable to create mechanical movement on the cells mimicking the action of human breathing. Ingber et al. designed a PDMS-based chip with two lateral channels, which could be vacuum-actuated to stretch a semipermeable PDMS membrane that has been coated with extracellular matrix (ECM) proteins such as collagen and fibronectin to allow cell attachment (Fig. 26.1A) [11]. In such a way, they were able to mimic the mechanical stress cells experience in the alveoli while breathing. Using this model, the response of the immune system toward inflammatory stimuli on the alveolar epithelium could be evaluated. More recently, the same group developed a lung-onchip platform replicating the bronchial airways for studying early biomarkers of COPD in response to cigarette smoking [10]. A novel approach to mimic a functional lung was taken by Young et al., who used a hydrogel microlayer consisting of type I collagen and matrigel in a lung-onchip-device to introduce smooth muscle cells combined with epithelial cells and study chronic lung diseases [51]. Although their model did not feature the characteristic mechanical actuation in the lung, the introduction of the hydrogel, replacing the semipermeable PDMS membrane, allowed for the culture of different cell components of the lung, eventually resulting in a more complex lung-on-chip model.

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Figure 26.1 Microphysiological systems to replicate lung, brain, kidney, and liver. (A) Schematic representation and photographs of a lung-on-chip device made of PDMS. (B) Concept of brain-on-chip platform culturing neurospheroids under constant flow. (C) (Top) Schematic of nephron highlighting the convoluted proximal tubule and fabrication process of a 3D proximal tubule model and (bottom) photographs of the bioprinting process of proximal tubules and a mature (fully confluent) tubule. (D) (Top) Schematic representation of the glomerular capillary wall in the kidney and (bottom) schematic representation of the kidneyon-chip device. (E) Schematic representation of a liver-on-chip model containing hepatocytes and hepatic stellate cells. (F) Schematic representation of a liver sinusoid-on-chip for highthroughput drug screening. (Continued)

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26.3.2 Brain

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The brain parenchyma consists of a highly interconnected network of neurons, astrocytes, microglia, and oligodendrocytes embedded into a dense ECM [47]. As the human brain differs immensely from the brain of rodents, the approach to develop a biologically relevant human brain-on-a-chip is of high interest to study biological processes as well as the effects of diseases on the brain [52]. Nevertheless, the fabrication of a brain-on-chip is particularly challenging due to the fact that the cells from the brain are not able to display their characteristic phenotypes in monolayers [53]. Neurospheroids, self-assembled clusters of neuronal cells forming a 3D structure, have recently been proven a promising strategy to provide cells with the necessary 3D architecture to display in vivo
like phenotypes [13,28]. The implementation of cultured neurospheroids into PDMS-based microfluidic devices has thus become a common strategy to fabricate brain-on-chip platforms. Additional application of constant flow to the spheroids providing fresh nutrients and removing metabolites has been demonstrated to enhance spheroid stability and growth as shown by Lee et al. using a single-chamber PDMS-chip including oval microwells that held the spheroids in place when exposed to medium flow (Fig. 26.1B) [13]. By introducing a porous PDMS-membrane that hosted cultured brain microvascular endothelial cells separated from below cultured neurospheroids, Guerrero-Cazares, Levchenko, et al. proved the suitability for such models to be used in mimicking blood
brain barrier
like properties [12]. A different approach to provide a 3D culture environment is the use of an ECMlike hydrogel in which cultured cells are embedded. In combination with a PDMSbased chip, that separates cultures of different cell types, Kim, Tanzi, Cho, et al. recently developed a brain-on-a-chip model that included neurons, astrocytes, and microglia embedded in matrigel to provide the required 3D environment [54]. As using ECM-like hydrogels with embedded cells is a key characteristic for recent

Source: (A) Reproduced with permission Huh D, Hamilton GA, Ingber DE. From 3D cell culture to organs-on-chips. Trends Cell Biol 2011;21(12):745
54 [49]. ©2011, Cell Press. (B) Reproduced from Park J, Lee B, Jeong G, et al. Three-dimensional brain-on-a-chip with an interstitial level of flow and its application as an in vitro model of Alzheimer’s disease. Lab Chip 2015;15(1)141
50 [13]. ©2014, Royal Society of Chemistry. (C) Reproduced with permission Homan K, Kolesky D, Skylar-Scott M, et al., Bioprinting of 3D convoluted renal proximal tubules on perfusable chips, Sci Rep 2016:6:34845 [70]. ©2016, the authors, published by Nature Publishing Group. (D) Reproduced with permission Musah S, Mammoto A, Ferrante T, et al. Mature induced-pluripotent-stem-cell-derived human podocytes reconstitute kidney glomerular-capillary-wall function on a chip. Nat Biomed Eng 2017;1:0069 [16]. ©2017, Nature Publishing Group. (E) Reproduced with permission Lee S, No da Y, Kang E, et al. Spheroid-based three-dimensional liver-on-a-chip to investigate hepatocyte-hepatic stellate cell interactions and flow effects. Lab Chip 2013;13(18):3529
37 [20]. ©2013, Royal Society of Chemistry. (F) Reproduced with permission Gori M, Simonelli M, Giannitelli S, et al. Investigating nonalcoholic fatty liver disease in a liver-ona-chip microfluidic device. PLoS One 2016;11(7):e0159729 [19]. ©2017, the authors, published by the Public Library of Science.

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bioprinting approaches, 3D bioprinting of cerebral structures represents a promising approach to create controlled 3D architectures. Biomaterials that have been proven to support cerebral tissue function and maturation in 3D bioprinting are in particular gelatin methacryloyl (GelMA), a combination of polycaprolactone (PCL) and gelatin [55], or RGD-modified gellan gum [56], among others.

26.3.3 Heart The heart is the first organ to form during fetal development and while growing, when cardiomyocytes become differentiated and aligned cardiac muscular cells that are connected by gap junctions [47,57]. This permits electrical stimuli to be transmitted between cells, eventually resulting in contraction of the tissue [14]. Replication of this defined cell alignment as well as the contractile properties is crucial to create a functional heart-on-a-chip system. To achieve these key properties, human cardiomyocytes, in OOC platforms often derived from induced pluripotent stem cells (iPSCs), require culturing in a non-monolayer or 3D fashion [58]. Biomaterials suitable for such 3D culturing are usually based on fibrin [15], GelMA, matrigel [59], or similar polymeric formulations, allowing the cells to proliferate and mature toward fully functional cardiomyocytes. Besides the biomaterial, another aspect in the culture of functional cardiomyocytes is the proper alignment of these cells. Over the years, different methods have been developed to promote cardiomyocyte alignment. One way is the use of a scaffold that allows cells to attach and guides the alignment to achieve 3D cardiac tissue constructs as shown by VunjakNovakoic et al. using cardiomyocytes embedded in matrigel on top a collagen stamp as scaffold [59]. In addition, it has been shown that applying a constant flow to the cardiomyocytes further promotes cell alignment toward a cardiac-like tissue construct [59,60]. Moreover, 3D bioprinting offers a novel and promising technique to create volumetric constructs that guide cell alignment [61,62]. Recently, Zhang, Khademhosseini, et al. fabricated a vascularized cardiac tissue model based on bioprinted GelMA and alginate microfibers, which allowed controlled anisotropy [62]. Implementing this bioprinted construct into a microfluidic platform facilitated the application of a controlled flow as well as demonstrated its utility as a drugscreening platform. A different way to enhance cardiomyocyte alignment and self-contraction is the exposure to an external electrical stimulus [63] or by enhancing the conductive properties of the used biomaterial matrix, which allows for an increase in cell
cell communication [64,65]. The addition of conductive materials such as gold nanorods or carbon nanotubes has been proven to enhance the electrical properties of used biomaterials such as GelMA, eventually being beneficial for cardiomyocytes maturation and alignment [64,65]. Shin, Khademhosseini, et al. have demonstrated how gold nanorods in a bioprinted GelMA construct improved cell-to-cell coupling of cardiomyocytes as well as promoted synchronized contractions [65]. The third technique to achieve cardiomyocyte alignment is the application of constant strains to the 3D tissue constructs. For instance, Marsano, Rasponi, et al.

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used the elastic properties of PDMS to design a microfluidic platform that entrapped a 3D tissue construct of cardiomyocytes embedded into a fibrin gel [15]. Upon the application of pressurized air into the chamber at the bottom, a controlled mechanical strain could be applied to the cardiac-like tissue construct, resulting in improved alignment of the cells.

26.3.4 Kidney The kidney includes several functional units consisting of different tissue types and cells such as glomerular cells, proximal tubule cells, loop of Henle cells, thick ascending limb cells, distal tubule cells, collecting duct cells, interstitial kidney cells, and renal endothelial cells [52]. All together these tissues play an important role in removing waste and drugs from the body, balancing the body’s fluid, regulate the blood pressure by hormone production, as well as controlling the production of red blood cells [47]. However, recent studies have shown that to replicate the function of the kidney in a realistic manner, only four of these tissue types need to be cultured in a biologically relevant setting: glomerular, proximal and distal tubule, and collecting duct cells [52]. Nevertheless, so far no kidney-on-chip device is available that includes all of these tissue types. Microfluidic platforms that mimic functional units of the kidney do not significantly differ regarding the design of the chip or the biomaterials used. In the most common form, these devices are based on a two-channel microfluidic chip made from PDMS, where the two channels are separated by a (ECM protein-coated) porous membrane on which cells are seeded. Ingber et al., for instance, fabricated a kidney proximal tubule-on-a-chip by seeding primary human proximal tubular epithelial cells on top of a collagen type IV-coated porous polyester membrane while being exposed to a constant flow of medium [66]. In the bottom channel, interstitial fluid was flushed to mimic the function of the proximal tubule found in vivo. A more recent approach to mimic the biologically relevant architecture of the kidney was the fabrication of a hollow tubular fiber membrane based on polyethersulfone as shown by Stamatialis et al., instead of using a flat membrane [67
69]. The use of a hollow tubular structure including proximal tubule epithelial cells might form a better model for the proximal tubule due to increased surface area; however, on-chip applications of this fiber remain to be investigated. 3D bioprinting forms a promising strategy to create such hollow tubular structures for the fabrication of a proximal tubule model using sacrificial bioprinting (Fig. 26.1C) [70]. Lewis et al. recently further demonstrated the use of same bioprinting strategy to create a vascularized proximal tubule model [50]. By using a mixture of pluronic and high
molecular weight poly(ethylene oxide) as sacrificial bioink, they were able to bioprint tubular structures within a fibrinogen/ gelatin hydrogel. After chemical crosslinking of the fibrinogen/gelatin hydrogel using thrombin, the pluronic was thermally liquefied at 4 C and removed from the construct, leaving hollow tubes. Seeding of proximal tubule epithelial cells in the proximal tubule channels and glomerular microvascular endothelial cells in the vascular channels displayed in vivo
like properties such as the active

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reabsorption of albumin and glucose via tubular
vascular exchange as well as endothelial dysfunction based on hyperglycemia induced in the system. The biologically relevant architecture and function of their system make it a promising in vitro model to study kidney function as well as the response to certain diseases. Besides mimicking the proximal tubule in the kidney, focus was also given on a glomerular-capillary-wall model on-chip as demonstrated by Ingber et al. [16]. Using a microfluidic device similar to a lung-on-chip platform, where two channels in PDMS are separated by a porous membrane that can be stretched by applied vacuum in two lateral channels, they were able to replicate the interface between glomerular cells and capillary channels (Fig. 26.1D). Seeding iPSC-derived podocytes on top of a laminin-coated PDMS membrane while seeding endothelial cells on the other side of the membrane allowed them to mimic the function and architecture found in the glomerular capsule of the kidneys. Although current platforms are still juvenile, the impact of high throughput kidney-on-chip platforms for rapid drug screening could be of high impact on the society, in particular that the development of new drugs can be significantly optimized as current 2D culture systems or in vivo animal tests fail to replicate the human functions of the kidney and do not allow to properly predict nephrotoxicity of pharmaceutical compounds [17].

26.3.5 Liver Due to its high metabolic activity, the liver is a critical organ to human life, whose major functions include the diffusion of chemicals in and out of the blood, liposomal transport, chemical detoxification, and drug metabolism [47,71]. Characteristic for the liver’s architecture is the hexagonally shaped hepatic lobules. The smallest functional unit of these lobules is the so-called sinusoids, which are repetitive microvascular units that connect the portal vein and hepatic artery with the central vein [47]. Kupffer cells are one of the most abundant cell types in the sinusoids forming a selective barrier together with endothelial cells between the blood and the underlying hepatocytes and hepatic stellate cells [72]. Current liver-on-chip platforms in particular focus on proper in vitro cultures of primary hepatocytes and hepatic stellate cells as these cells are difficult to culture in a conventional 2D fashion and lose their characteristic hepatic functions rapidly [31,73]. Similar to previously mentioned 3D neurosperoids, these hepatocytes and hepatic stellate cells can also self-assemble to create 3D spheroids that can eventually be cultured in microfluidic devices. Using an interconnected PDMS device that consisted of defined microwells to culture spheroids, Lee et al. were able to investigate the interactions between hepatic stellate cells and hepatocytes in a dynamic 3D microenvironment (Fig. 26.1E) [20]. A different liver-on-a-chip model consisted of a welldefined array of silicon microtrenches of 4 μm in width, which were coated with heparin to allow hepatocyte attachment as well as prolonged culture duration [74]. Combined with a dynamic flow over these microtrenches, these hepatocytes showed long-term viability and stability, making this model suitable for drug-screening studies. Similar to the fabricated microtrenches, several liver-on-chip devices based on

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PDMS contained well-defined micropillars that separated cultured cells from a constant fluid flow [31,75]. In such a way, it was possible to replicate the endothelial layer in sinusoids as shown by Rainer et al., where different cells could be cultured in defined areas and exposed to drugs in a high-throughput manner (Fig. 26.1F) [19]. Besides different OOC approaches, bioprinting functional liver tissues are also of particular interest. Bioprinted vascularized liver tissues based on a GelMAhyaluronic acid (HA) bioink have shown great performance in vivo animal studies [76]. Furthermore, Khademhosseini et al. demonstrated the combination of bioprinted liver spheroids with the OOC technology to achieve a functional liver-onchip model for drug testing [18].

26.3.6 Gut Current in vitro culture models for the gut or intestine fail to mimic the human gut physiology because of the complex natural architecture found in the gut as well as the mechanical motion that is present in the gut during food digestion and nutrient absorption and the difficulties in incorporating gut microbiota, which is essential for gut function [47]. In addition, the wall of the intestine includes millions of small 3D structures termed the “villi,” which themselves consist of different types of epithelial cells, such as absorptive, mucus-secretory, enteroendocrine, and Paneth cells [22]. Although targeted for different applications, the general designs of gut-onchip platforms are similar to previously mentioned models based on the twochannel PDMS chip with a porous membrane separating the two channels [21,77]. Lateral channels with applied vacuum allow for mechanical stretching of the porous membrane mimicking the peristaltic motion of the intestine (Fig. 26.2A). A key characteristic of recent gut-on-chip models is the use of Caco-2 cells to be seeded on top of a collagen type I and matrigel-coated porous PDMS membrane [22]. Upon liquid flow and peristaltic motion, these cells underwent spontaneous morphogenesis of 3D villi, displaying gut-specific characteristics such as tight junctions, coverage by brush borders as well as mucus [77,79]. In addition, depending on the 3D architecture, these cells display the differentiation toward absorptive, mucus-secretory, enteroendocrine, as well as Paneth cells in the characteristic location within the villi similar to the in vivo situation [22]. Besides the in vivo-like properties of the cell culture, the formation of villi also significantly increases the surface area, making such gut-on-chip models highly promising for drug screening.

26.3.7 Muscle In the human body, three different types of muscle tissues can be found: visceral, cardiac, and skeletal muscles [80]. Visceral muscle forms the weakest of muscle tissues and can be found in the stomach, intestines, and blood vessels, where it is particularly involved in the movement of substances based on contractile movements. Cardiac muscle is especially found in the heart and has been introduced and explained in the subsection describing “heart-on-chip platforms”. Both visceral and cardiac muscles are so-called involuntary muscles, meaning that they cannot be

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Figure 26.2 Microphysiological systems to replicate gut, muscle, bone, and multiorgan platforms. (A) Schematic representation of a gut-on-chip. (B) Photograph and schematic of contractible muscle strip on chip (MTF). (C) Schematic displaying the process of in vivo engineered bone marrow and further culture on a bone-on-chip platform. (D and E) Concepts of multiorgan platforms containing 10 different organ types on a single chip. MTF, Muscle thin film. Source: (A) Reproduced with permission Kim H, Huh D, Hamilton G, et al. Human gut-ona-chip inhabited by microbial flora that experiences intestinal peristalsis-like motions and flow. Lab Chip 2012;12(12):2165
74 [21]. Copyright 2012, Royal Society of Chemistry. (B) (Continued)

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controlled consciously. Skeletal muscle, which is in the focus of most muscle-onchip platforms, is the only muscle that can be contracted consciously and, as the name implies, is always connected to the skeleton [80]. The smallest functional unit in such muscle tissues is the myofibrils inside of myocytes (muscle cells), which themselves are assemblies of repeated sarcomeres [47]. An important aspect for the fabrication of (skeletal) muscle is a strong alignment of muscle cells as well as the capability to contract upon an external stimulus. To achieve muscle alignment in microfluidic platforms, different strategies have been developed that allow for the controlled culturing of muscle cells to form wellaligned muscle patches. One approach is to seed cells in or on top of a well-defined pattern, which can be created by for instance microcontact printing or stamping [24,81]. Parker et al. used microcontact printing of fibronectin on top of PDMS [24]. Seeded skeletal muscle cells on top of the defined pattern self-organized in the direction of the printed pattern. In such a way, they achieved a well-aligned muscle-like film that could be placed in a microfluidic chip based on PMMA that was laser-cut to form a single microfluidic channel with inlet and outlet (Fig. 26.2B). Upon applying an external electric stimulus, these thin films were able to contact in the chip, replicating the natural contraction of skeletal muscle tissue. The approach of seeding cells in a predefined manner to guide their orientation has also been applied in bioprinted skeletal muscle tissues, where muscle cells were bioprinted in repeating parallel lines to achieve well-aligned muscle constructs [82,83]. In a different approach, such muscle networks could be obtained using the self-orienting properties of muscle cells around defined anchoring points in a 3D environment such as a hydrogel [23,84]. Varghese et al. applied this approach, using mouse myoblasts embedded in GelMA [23]. The GelMA hydrogel was placed in a polyacrylamide (PAm)-based microfluidic chip containing two PAm pillars in such a way that it would form an elongated oval tissue around the two pillars. Varghese et al. found that, using the pillars as anchoring points, the embedded muscle cells aligned themselves within the gel. Furthermore, they were able to measure the displacement of the pillars due to the exerted force by the formed muscle tissue onto these pillars, indicating the passive tension created by the muscle tissue during formation.

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The bone tissue is arguably one of the toughest tissues in the human body. Nevertheless, depending on the location, the bone can have different structures and Reproduced with permission Grosberg A, Nesmith A, Goss J, et al. Muscle on a chip: In vitro contractility assays for smooth and striated muscle. J Pharmacol Toxicol Methods 2012;65(3):126
35 [24]. Copyright 2012, Elsevier. (C) Reproduced with permission Torisawa Y, Spina C, Mammoto T, et al. Bone marrow
on
a
chip replicates hematopoietic niche physiology in vitro. Nat Methods 2014;11:663
9 [27]. Copyright 2014, Nature Publishing Group. (D and E) Reproduced with permission Wang Y, Oleaga C, Long C, et al. Self-contained, low-cost body-on-a-chip systems for drug development. Exp Biol Med 2017;242(17):1701
13 [78]. Copyright 2017, SAGE Publishing.

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functions. The tubular shaft of the femur (diaphysis) on the one hand consists of dense compact bone surrounding the yellow bone marrow in the middle [47,85]. The ends of the bone (epiphysis) on the other hand form the spongy bone, where some parts might be covered with articular cartilage [86]. Red bone marrow can be found throughout the spongy bone [86]. Though structural differences, both types of bones consist of mainly osteocytes, that maintain the bone, osteoblasts, that form the bone matrix, osteoclasts, that resorb bone matrix, as well as stem cell
like osteogenic cells, that eventually become osteoblasts [47,85]. All of these cells are embedded in a dense matrix composed of collagens and hydroxyapatite crystals [85]. Providing such a matrix in a biologically relevant manner is a critical step in the engineering of a bone-on-a-chip platform. Although different strategies have been developed over the years to provide this microenvironment, the design of bone-on-chip platforms remains rather simple, often including a main cylindrical chamber holding a bone tissue construct and a channel/chambers to provide perfusion of medium and nutrients located above of the construct or combined above and below the construct [26,27,87]. Such platforms are usually based on PDMS. A particular challenge, however, in bone-onchip systems is the development of a 3D bone construct for culture in the chip device. Zheng et al. used a heavily mineralized collagen matrix to induce bone formation based on encapsulated osteoblasts after culture for 720 hours [26]. In a different approach, Ingber et al. used mice as the culture environment to form biologically relevant bone including bone marrow in vivo for a bone marrow
onchip device (Fig. 26.2C) [27,87]. They provided a bone growth-stimulating collagen gel including demineralized bone powder and bone morphogenetic proteins (BMP2 and BMP4) in a tubular PDMS structure. After implantation into mice for 8 weeks, they observed the formation of dense bone surrounding a core of bone marrow, which was further cultured in a PDMS platform containing a main cylindrical chamber and two medium perfusion channels on top and below as previously described. A different approach for a bone-on-chip platform was recently taken by George, Truesdell, York, Saunders et al., who developed a multichamber microfluidic PDMS chip to study the interactions between osteocytes, osteoclasts, and osteoblasts [25]. A main chamber holding osteocytes, which became activated by mechanical load onto the chamber, was connected to chambers holding either osteoblasts, osteoclasts, or a coculture of both. As such, they could study the crosstalk of these cells in a controlled manner. The fabrication of functional bone tissues is also in the interest of several groups focusing on bioprinting [83,88
91]. Although these bone constructs are designed for the use in regenerative medicine, the functionality as proven by in vivo studies makes them also highly attractive to be used in on-chip platforms for rapid screening of drugs and therapeutics.

26.3.9 Multiorgan A multiorgan-on-a-chip platform, as the name implies, consists of multiple organs that are interconnected on a single chip device. Such a platform is of particular

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interest to study the interactions between different organs in the case of disease or when exposed to certain drugs, but also to get a better insight in the fundamental crosstalk between them. The simplest version of a multiorgan-on-a-chip is a platform that interconnects two organs. To this end, different dual-organ platforms have been developed such as a liver
gut
chip [29,30,92] or a liver
pancreas
chip [93], chips that mimicked the interactions between organs and surrounding vasculature [94], and a combined brain
muscle
chip that replicated muscle activation by electrical stimulation from brain neurospheroids [95]. The general design of such platforms is often similar to that for single-organ platforms, where the functional parts of the different organs are cultured in their respective confined compartments. Similar to single-organ chips, cells in these compartments are often cultured in monolayer on membranes, 3D in hydrogels, or 3D in the form of spheroids. Microchannels between the different compartments mimic the vasculature that connects different organs in the human body. Such microchannels can either promote a unidirectional flow with an inlet and outlet [95] or a circulating flow between the organs [93]. A more sophisticated approach is the body- or human-on-a-chip platform, where up to 14 organs have been interconnected to mimic the human body in a more realistic way (Fig. 26.2D and E) [78,96]. Human-on-chip platforms have a high potential to partially replace or reduce animal studies as these platforms show higher reproducibility, overall lower costs, higher work efficacy, and often a well-defined readout, which is difficult to achieve in animals [97]. Human-on-chip systems are in general multicompartment chips, where every compartment contains a microphysiological model of a certain organ similar to the single-organ chips before. These compartments are then interconnected in a biologically relevant sequence using microchannels, mimicking the circulatory system in the human body. In addition, these chips often contain a circulating flow. In such a way the blood flow from, to, and through different organs can be replicated in a biologically relevant manner, which makes these chips an ideal model to study the effects of drugs or pathophysiological conditions on multiple interconnected organs. Nevertheless, the design of such systems includes multiple challenges with regard to biological and technical fabrications [97,98]. Biological challenges are in particular the correct scaling of organ in relation to each other, proper vascularization of thick 3D tissue constructs in the compartments, the development of the specific immune response of the human body toward drugs/pathogens, or the proper culturing of multiple different types of cells in a single circulating system. As every organ-like compartment might require specific growth requirements to fully form and maintain a functional tissue-like construct, the implementation of all these factors in a single universal growth medium for all organs is highly challenging. Likewise, technical challenges that need to be addressed in the system are for instance the maintenance of a well-controlled flow between the organ compartments, proper nutrition and oxygenation of the different organs, as well as the material choice to create such human-on-chip platforms. PDMS, which is used in most OOC platforms, allows to create highly defined architectures in a chip. However, drugs and growth factors might bind or be absorbed to the material. Indeed, it was

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found that with a circulating flow in these systems, the adsorption of drugs will be significantly higher compared to platforms with an inlet and outlet [99]. Despite these challenges, the development of a human-on-a-chip platform remains one of the most promising models to replace/reduce testing of animals and provide insights into the working of drugs or the pathophysiological processes in a biologically relevant environment. Furthermore, recent efforts in the development of human-on-chip platforms have generated promising results. For example, Maschmeyer et al. developed a four-organ-model connecting human intestine, liver, skin, and kidney on a single chip, which showed human-like homeostasis between the organs sustainable for approximately 28 days [100].

26.4

Organ-on-chip platforms to mimic human pathophysiology

As previously mentioned, OOC platforms are interesting not only to study the biology of complex organs and cell
cell or cell
environment interactions but also to investigate pathophysiological conditions in the human body [2]. Such conditions are often defined by a loss of the natural homeostasis of the body based on external stimuli, such as chemicals, or by mutation in the body [2,47]. Although all the abovementioned models have the potential to function as disease models by introducing a certain stimulus that causes a disease state, some of them are particularly designed to study pathophysiological processes. For instance, Ingber et al. combined their lung-on-chip platform with a cigarette-smoking device to study the effects of smoking on the bronchial epithelium [10]. Using this platform, they were able to identify ciliary micropathologies as well as novel COPD-specific molecular signatures, demonstrating the capacity of the model to study the development of pathophysiological conditions in healthy individuals. Similarly, the aforementioned brain-on-chip model by Lee et al. combining the culture of neurospheroids with a constant intestinal flow was evaluated for the investigation of Alzheimer’s disease [13]. They introduced amyloid-β into the system, which is one of the major contributors in Alzheimer’s disease, and studied the toxic effects on neurons. They especially found that the intestinal flow worsened the effects of amyloid-β on neuronal damage. Different from the previously described models mimicking the functional part of a full organ, standalone models for vascular structures are also of interest to study pathophysiological conditions. In the context of OOC platforms, two types of disease states have been studied extensively: atherosclerosis [101
103] and thrombosis [104]. Atherosclerosis is a chronic inflammatory disorder that is accompanied by endothelial dysfunction and stenosis, a narrowing of the blood vessels. Similar to previously mentioned models, PDMS-based microfluidic chips allow for the reconstruction of perfusable channels to study atherosclerosis. For example, Hou et al. developed a PDMS-based atherosclerosis chip including a vascular channel separated by a flexible PDMS membrane from an orthogonal air channel [103].

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Introducing air into the lower channel enables the deformation of the PDMS membrane, mimicking stenosis in the upper vascular channel. The chronic inflammation of the endothelial layer was induced by exposing human umbilical vein endothelial cells (HUVECs) to tumor necrosis factor-α. They found that monocytes that were flushed through the channel showed increased binding to inflamed endothelium as well as a further increase once stenosis was modeled in the platform. They could further validate their system using samples of whole blood, demonstrating the applicability of their system as a point-of-care blood profiling device. Thrombosis is the formation of a blood clot inside a vessel, which obstructs the blood flow from the circulatory system [47]. Such a blood clot is based on an injury of the endothelial layer resulting in the formation of a clot to prevent blood loss. A clot that breaks free from its location and traveling through the body is known as an embolus, which, depending on its origin and location, can cause myocardial infarction, stroke, or pulmonary embolism [47]. It has been shown that fibroblasts surrounding the endothelial layer play a critical role in the formation of a blood clot resulting in thrombosis. Bioprinting offers a strategy to manufacture well-defined 3D architectures, including perfusable vascular channels, surrounded by tissues, including other cell types. For instance, Khademhosseini et al. fabricated a bifurcated channel including a HUVEC layer in a GelMA block consisting of human dermal fibroblasts using sacrificial bioprinting [104]. They induced a blot clot into the system using calcium chloride mixed with human whole blood. Treatment of this model with tissue plasmin activator led to dissolution of the clot, showing the clinical relevance of the model. Furthermore, they found that fibroblasts surrounding the blood vessels resulted in a fibrotic response leading to aging of the clot, similar to the clinical situation. Also, common is the introduction of a diseased state in an otherwise healthy tissue by exposure to an external stimulus. However, models that only mimic the diseased state are also used to study pathophysiological conditions. Of particular interest in this field are models that resemble cancer. Over the recent years, different 3D models have been fabricated to mimic cancer in a biologically relevant manner. Such models include spheroids or organoids [4,7,8], cancer-on-chip models [2], or bioprinted models [105,106]. Although the use of 3D models in cancer research and drug development has been known for decades, recently the tumor microenvironment (TME), cells, and ECM surrounding the tumor cells have drawn increasing attention due to its importance in tumor progression, aggressiveness, and invasiveness [107,108]. Recent 3D models often focus on the interaction of the TME and cancer cells as well as investigate the response of these cells toward anticancer therapeutics. Although different in cellular and ECM composition, such models are often based on the simple coculture of cancer cells and in the TME components in a 3D manner. Hereby, cancer cells and TME-associated cells are placed in a 3D hydrogel next to each other, allowing to study the invasion of cancer cells into the TME surrounding. Such models can be either fabricated using microfluidic platforms [109] or 3D bioprinting [105,110]. For instance, Heinrich, Prakash, et al. recently created a 3D-bioprinted model for glioblastoma multiforme, the most malignant type of brain cancer, to study the interaction of glioblastoma cells and

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cells form the immune system [105]. They found that the coculture of these cells in a relevant 3D environment displayed the upregulation of gene markers that are also found to cancer patients. Furthermore, they were able to evaluate the clinical relevance of their model by the use of chemo- and immunotherapeutic drugs that have been previously used in a clinical setting. Similarly, Belanger and Marois recently created bioprinted models of breast and pancreatic cancer using 3D bioprinting to investigate the interactions of patient-derived cancer cells with surrounding fibroblasts [110]. Further evaluation of their model using cancer-specific therapeutics could demonstrate the clinical relevance of the model. The use of microfluidic platforms to mimic pathophysiological conditions allows for the study of such conditions in a controlled and well-defined environment. Especially, the application of such on-chip devices as higher-throughput screening platforms allows for the rapid study of stimuli causing a pathophysiological state as well as for the testing of various drugs and therapeutics. In such a way, drugs can be developed in a more relevant in vitro setting, especially reducing the amount animals needed for drug evaluation as well as reducing the time to develop new drugs.

26.5

Applications beyond conventional research

The use of microphysiological platforms is not solely limited to the benches of scientific research but might also play significant roles in other fields due to their small sizes, comparably low costs, high reproducibility, and promising capability to mimic functional units of human organs in a biologically relevant manner as discussed previously. This subsection will take a closer look on how microphysiological systems can be used to gain insight on biological and physiological processes that astronauts undergo in space or how these systems can save the life of millions here on Earth.

26.5.1 Space Over the recent years, many programs have been developed with the aim of utilizing the microphysiological systems for studying human physiology and pathology in space [111]. Most OOC platforms are comparably small, lightweight, and low in costs while still providing a way to culture cells in a biologically relevant manner to mimic human tissue and organ functions. These characteristics make them ideal models to test the effects of microgravity and radiation on tissues and organs and to investigate the body’s response to stress, drugs, and genetic changes [112]. For instance, the effects of microgravity on muscles can be investigated to mimic muscle atrophy that astronauts experience in space [113]. In addition, space might also affect the brain, especially the blood
brain barrier [112]. Brain-on-chip devices could give insight on these effects in an effective way. Of particular interest to test the effects of the space environment, however, is formed around multiorgan- or

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human-on-chip models, as they combine and interconnect different organs and represent the human body in a more realistic way. Furthermore, the use of OOC platforms can accelerate the research on ageassociated diseases such as bone and muscle loss [114,115]. In space, these diseases occur much faster compared to the environment on Earth. Using respective organ models, researchers can investigate the progress of bone and muscle loss, which on Earth occurs in months or years, in a significantly reduced amount of time. Recently, SpaceX sent the first OOC platforms to the international space station to test the effects of microgravity on different organs and study the effects of drugs, stress, and genetic changes as well as pathophysiological processes [114]. A different approach to use lab-on-chip technologies in the space form the socalled Lab-on-a-Chip Application Development-Portable Test System [116]. This handheld device uses lab-on-chip technologies to analyze biological and chemical substances on surfaces within the spacecraft and in such a way helps to improve the biological and chemical cleanliness in the cabin environment of the international space station. Similar to this handheld device, several studies have focused on the fabrication of stand-alone OOC devices [117,118]. Current platforms often require larger computers that actuate perfusion flows and negative pressures, among others. Stand-alone systems would form a huge advantage for the use in space as they require significant less space and operational costs.

26.5.2 Military The US military and Department of Defense have shown significant interest in microphysiological systems as these platforms allow to test the effects of biological, chemical, or radiological weapons under biologically relevant conditions [119]. As currently the testing of such weapons is highly limited due to ethical and other reasons, such platforms can offer an effective alternative to screen the response of the human body toward such weapons. Of particular interest for military applications are multiorgan-platforms as they allow to study the response of a human-like system consisting of multiple interconnected organs toward natural or man-made biological threats [119,120]. For example, in a program supported by the Defense Threat Reduction Agency, Wake Forest Institute of Regenerative Medicine has partnered with Harvard Medical School, University of Michigan, the Edgewood Chemical Biological Center, and Morgan State University to investigate such threats using a four-OOC platform [119]. In case of an attack, OOC platforms can give rapid information about the type of attack and the possible effects on the human body, shortening the development of an effective treatment. In addition, knowing the realistic effects of such biological, chemical, or radiological weapons can be a significant advantage in battle strategy and save the lives of soldiers in battle. The development of OOC platforms that replicate the functional units of different tissues or diseases has showed tremendous capability to mimic organ functions and physiology in vitro. Interconnecting different single-organ units to a multiorgan- or human-on-chip platform is a promising approach toward replacing/reducing

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animal studies for drug screening or to obtain insights on the effects of pathophysiological processes on the full body. Although these platforms offer a great range of applications, the fabrication however remains difficult with regard to proper organ scaling and the culture of multiple tissues in a single circulating flow that all require different factors to grow and maintain their functions. A carefully designed universal medium that fulfills the needs of all organs in a given platform could solve this problem but remains to be investigated. Furthermore, proper vascularization of thicker 3D tissues in the chips, allowing sustained culture for longer periods of time, needs to be addressed. Bioprinting might offer a promising solution for the fabrication of vascularized tissue constructs as it allows for the simple implementation of multiple cells into a single construct with well-defined architecture. In fact, several bioprinted models focus on the proper vascularization of thicker tissue constructs for regenerative medicine [76,105,121]. As important their applications, the building constituents, design, and biomaterials used in the microphysiological systems are equally important since they are the defining factors for the success of these platforms. In the next section, we discuss some of the important chip components and different biomaterials that are used to encapsulate cells in these systems.

26.6

Biomaterials for chip fabrication

Different biomaterials are used to fabricate chips or bioreactors during the development of different OOCs. Since determining the optimal physicochemical properties is the vital step toward successful biomedical application of these devices, to date, various natural and synthetic biomaterials with specific features have been utilized [122]. We discuss some of them here.

26.6.1 Elastomers Elastomeric polymers have been playing an important role in the development of OOCs in general and more recently for flexible and stretchable microfluidic devices [123
125]. Utilizing different elastomeric polymers such as double-sided acrylic tape made from photocured aliphatic acrylate, complex geometries, and hybrid structures involving microfluidics and electronics with robust interface and functional microstructures have been achieved [126]. Recently, the development of elastomer-based microchannels with excellent elasticity and stability was reported by Wu et al. [127]. Using acrylic tape together with another elastomer, PDMS, the authors synthesized a flexible microchannel that could be directly assembled into various substrates of different origins such as organic, inorganic, and metallic for use in different OOCs. Among the different elastomers, PDMS has been extensively used for the microfluidic device development. Fabrication strategies utilizing materials such as PDMS polymers have long been established [128,129]. PDMS forms the basis for some of

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the cheap and fast prototyping techniques. 3D printing also allows for the creation of complex 3D structures that may not be possible with other fabrication techniques [130]. When selecting a microfabrication strategy for the development of engineered living systems, it is important to consider the effect of the construction material on the performance of these platforms or devices. PDMS offers an excellent set of properties for the development of OOC devices. Without the need for a clean room, PDMS is extensively used for its biocompatibility, low-cost, high transparency (240
1100 nm) that allow easy microscopic observations, low autofluorescence, and very good deformability that enables easy and leak-proof microfluidic connection OOCs [110,131
133]. When bioreactors consist of barrier tissues such as the lung epithelium, gastrointestinal tract epithelium, or renal epithelium, the devices may also need to comprise a porous membrane so that transport of the drug or small metabolites through these tissues can occur. Recently, thin membranes made of PDMS have been used in the development of lung-on-a-chip devices [11]. These membranes have the advantage of being ultrathin for mechanical maneuver and porous for better nutrient exchange as demanded by these special devices to replicate, for example, human breathing physiology [134]. Although PDMS-based OOC devices are biologically compatible, due to their hydrophobicity they typically have to be coated with ECM components to make cells stick to their surfaces. While these coatings provide favorable microenvironments for the cells to grow and proliferate on, the drawback, however, is that these surfaces sometimes adsorb and/or absorb drugs and metabolites [135]. Therefore it is important to consider drug
material interactions and avoid unnecessary complications to obtain accurate results. Some recent advancements have led to the introduction of more favorable biomaterials that allow better cell adhesion and proliferation, for example, hydrogels, as discussed later.

26.7

Thermoplastics

Thermoplastics are a class of materials that, by reaching a glass transition temperature, can be remolded repeatedly [135]. They can maintain their shape and are suitable for microwork processes [136]. Thermoplastics, usually optically clear, are resistant to small-molecule permeation and are more rigid than elastomers [137]. Depending on their applications, their surfaces may be modified by means of dynamic coating or surface grafting techniques [138,139]. In general, covalently modified surfaces are more stable than PDMS [139]. For instance, flexible circuit electrodes can be easily integrated, and thermoplastic surfaces can maintain their hydrophilicity for up to a few years after oxygen plasma treatment [135]. We briefly discuss some important ones. Polystyrene (PS) is an optically transparent, biologically compatible, and inert polymer, whose surface properties can be easily modified for the fabrication of biochips. PS surfaces can be modified to allow cell growth and adhesion with proper

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prior chemical modification [140]. Another polymer, polycarbonate (PC), is a durable material that is formed by polymerization of bisphenol A and phosgene [141]. This polymer offers a convenient solution for the production of multilayer devices [142,143]. Hot-embossing and subsequent gluing of two layers by means of thermal bonding are used to produce microfeatures in PC. A versatile thermoplastic that has commonly been used for microfluidic chip fabrication is PMMA. Because of its low cost, rigid mechanical property, excellent optical transparency, and electrophoresis compatibility, PMMA is particularly suitable for disposable microfluidic chips [144]. PMMA is an ideal material for “green microchips” as it can be easily decomposed into methyl methacrylate at high temperatures [145]. Other interesting thermoplastics used for chip design and manufacturing are poly (ethylene glycol) diacrylate (PEGDA), Teflon, and polyurethane (PU). PEGDA has been used to construct robust microfluid valves and pumps in different shapes [146,147]. Teflon chips on the other hand are used for their outstanding solvent stability [148]. Teflon has high chemical resistance, high gas permeability, and low toxicity [148]. In comparison to PDMS and PS, they also have low nonspecific protein adsorption [136]. PU elastomers are known for high mechanical strength, resiliency, and abrasion resistance [149,150]. Because of the hydrophobicity of PU, it has mostly been used in the manufacture of medical devices such as the artificial heart, intraaortic balloons, pacemaker leads, heart valves, or hemodialysis membranes [151]. A few recent reports have also demonstrated their applications in serving as the flexible membranes in whole-thermoplastic microfluidic chips [152,153]. Thermoplastics offer an excellent choice for the type of materials that are excellent for the fabrication of microfluidic chips. However, permeability to gas, resistance to swelling during molding, and overall hydrophobicity are some disadvantages that need more attention for their effective use in microfluidic chip fabrication in OOC applications.

26.8

Hydrogels

Hydrogels have several advantages over, for example, silicone-based materials for the design and applications of various chip components for OOC systems. They are preferred for their properties allowing diffusion of small molecules, biocompatibility with most cells, temperature sensitivity, relatively low cost, and ease of production [135]. A majority of hydrogels possess specific cell-binding sites that are desirable for cell attachment, spreading, growth, and differentiation. With the development of microfluidic applications, hydrogels can be integrated into microfluidic systems using different methods such as soft lithography, flowsolidification processes, or UV-curing [154,155]. Owing to their special properties, hydrogels are widely used as different modules from fluid flow control to biochemical reactions or biological applications. Natural and synthetic hydrogels have been used in the OOC devices [156]. Recently, hydrogels have been used for the

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development of biomimetic valves that could be opened or closed by a change of pH, similar to the venous valves [157]. The hydrogel valve not only combined sensing and actuating abilities but also required fewer fabrication steps compared to traditional microfluidic valves. In yet another study, hydrogels were used to develop 3D physio-mimetic ECMs that provided a supportive microenvironment for rodent and human islet cultures [158]. Combining casting and bonding processes, Nie et al. fabricated a hydrogel-based microfluidic chip of gelatin and GelMA that were used to establish a vessel-on-a-chip with vascular function in both physiological and pathological situations [159]. Hydrogel-based microfluidics are gaining momentum and we expect more novel materials to join the list for the development of robust, yet, physiologically more relevant OOCs with biologic and more biocompatible materials in the coming years [159,160]. We have discussed some important materials that are used in the chip fabrication for the OOC platforms. The next section details some biomaterials that are used in tissue fabrication for the biological components of the OOC platforms.

26.9

Biomaterials for tissue fabrication for organ-onchip platforms

Natural polymers: Natural biomaterials are polymers obtained from a natural source [161]. These biomaterials include ECM components such as collagens, elastin, proteoglycans, and HA. Other natural sources are plants and insects. Some of the biomaterials obtained from these sources including chitosan, gelatin, dextran, glucose, fibrin, fibroin, collagen, and alginate are actively used for in vitro applications. These and other natural polymers (e.g., chitosan and hydroxyapatite when mixed with PU, etc.) are typically elastic [162], displaying excellent biocompatibility and good oxygen permeability [163]. As such, these polymers are used as important biomaterials in tissue engineering, biosensor development, and drug-delivery systems. Apart from being highly biocompatible, their availability makes them materials of choice for biomedical applications including those in OOCs [164
166]. We discuss some of these natural biomaterials and also provide a brief overview of the synthetic ones. Collagens are biomaterials widely used in engineering tissue models in the development of microphysiological systems. Collagen types I
IV form major components of the native ECM and are the main proteins responsible for the structures of many connective tissues [102,167]. Collagen, therefore, has tissue-matching properties, combined with superior biocompatibility for in vitro and in vivo studies, and is used in the tissue fabrication for OOCs [168]. The other properties include low immunogenicity, and enzymatic degradability. Depending upon the purpose and the target OOC, different methods have been used to tune its mechanical properties and stability; for example, it was used for free-form scaffold printing [169]. A convenient property of collagens is that they have cell-adhesion sites. This makes

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it particularly useful for their incorporation into OOC platforms to provide basic adhesion sites for cells. Examples include platform models for heart [170], liver [171], skeletal muscle [170], kidney [172], neuronal network [173], and tumor spheroids [174], among other tissues. Not only as a tissue fabrication component of OOCs, collagens have also been used as a structural component for some of the microfluidic devices. These include the formation of biomimetic microvessels embedded in collagen hydrogels for microenvironmental studies of angiogenesis, a micro-placed hepatocyte spheroid culture system, the use of microcontact-printed collagen to form cell adhesion sites in 3D collagen, and the gels used to direct the formation of the neural network by controlling the cultivable gel area by means of photothermal etching method [49,175
177]. Gelatin is a single-chain protein extracted from collagens by partial hydrolysis [178]. It is similar to collagen in the composition, but its advantages over collagen are its lower cost and being less antigenic. It is biologically compatible, biodegradable, and can easily form a gel [179]. It is commonly combined with other natural biomaterials to support cell culture for OOC platforms [179,180]. McCain et al. used gelatin hydrogels as muscle thin film substrates for engineering purposes in cardiac tissues [177]. Another form of gelatin that has utility in OOC platform technology is the GelMA produced by chemical modification of gelatin with methacryloyl groups [181
183]. GelMA hydrogels are advantageous because their mechanical properties can be adjusted by changes in the degree of crosslinking by methacryloyl substitution and by UV curing [183]. GelMA has been extensively used in cardiac and vascular tissue models, among many other tissue models [159,182,184,185]. Fibrin, a fibrous protein composed of fibrinogen monomers, is involved in blood clotting caused by enzymatic polymerization of fibrinogen by thrombin [186]. It has been used as scaffolds for cell encapsulation, distribution matrices for soluble factors such as proteins or growth factors, bioadhesives, and scaffolds in the manufacture of various tissues [187,188]. For OOC platforms, fibrin has been used as a component of artificial ECM. For example, adipose-derived stem cells were used in a quasi-3D construct developed for the analysis of fluid and cell behavior and alginate/gelatin/fibrin hurdle was used to obtain cell adhesion, growth, and differentiation [189]. An alternative use of fibrin in OOC platforms is to examine the formation of clots [11]. HA or hyaluronate is a glycosaminoglycan and hydrogel constituting the ECM and present in connective, epithelial, and nerve tissues [190]. It is a structural component in the tissues and joints, and it is also found in the synovial fluid and vitreous humor, improving the biomechanical stability [190]. Moreover, it plays an important role in cell proliferation and migration, inflammation, and wound healing [191]. HA can be degraded by hyaluronidase produced by cells and chemically modified to increase its hydrophobicity and processed to fibers, microspheres, and membranes [192]. Scaffold preparation using HA is difficult due to its viscous nature. To overcome this, it is crosslinked with other synthetic polymers, such as poly(ethylene glycol) (PEG) and polylactic acid (PLA) to further enhance its gelforming properties [193]. The use of HA in OOC models is limited; however, a few

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examples included a system using HA as a biomimetic material with tunable stiffness to understand mechano-regulation associated with metastasis [194]. Chitosan, a linear polysaccharide, is an inactivated derivative of the mass in the shells of shellfish. It is structurally similar to glycosaminoglycans, which makes it an interesting candidate for engineering tissues [195]. Chitosan-based systems can also be developed to respond to light, pH, temperature, and ionic concentration [196]. Due to its inherent antimicrobial properties, it also provides a suitable environment for long term cultures with less contamination risk. It is highly crystalline and can be easily processed in a variety of forms in conjugation with other biomaterials to form hydrogels, nanofibers, nano- and microparticles, and sponges [197]. In conjugation with other materials, chitosan has been used as a tissue component in OOCs. 3D chitosan
alginate used to serve as an ECM that promoted the conversion of cultured cancer cells to a more malignant in vivo
like phenotype [198]. In another study, chitosan
glycol matrix was used to produce cardiomyocyte spheroids [101]. Alginate is a polysaccharide block copolymer composed of glucuronic acid and mannuronic acid derived from brown algae [199]. It is useful for many applications because it is low-cost, has low toxicity, and is easy to be chemically modified [195]. One of the benefits of using alginate is its immediate gelation when exposed to divalent cations such as calcium (usually in the form of calcium chloride solution) that can allow the formation of microfibers or special structures. Additional polymers can be added to enhance its performance for scaffold formation [134]. We discussed some natural biomaterials based on the polymers obtained from mammalians, plants, and other living systems. They offer many interesting properties such as biocompatibility, and good mechanical properties that can be tunable through chemical or photocrosslinking methods. However, the use of natural biomaterials has some disadvantages as well. First of all, because they come from natural sources, there are significant differences on a lot basis. They must also be sterilized and purified that add further complexities in their use. On the other hand, synthetic biomaterials offer a comparatively better control of chemical and physical properties. We next briefly discuss some of the synthetic biomaterials. Synthetic biomaterials: Some of the drawbacks of naturally obtained biomaterials can be addressed using synthetic biomaterials. New biomaterials are being developed at a rapid pace to augment OOCs while doing away with the limitations of the natural polymers in imitating both therapeutic and basic biological studies. One major advantage is that the chemical and physical properties of these polymers are more adjustable than those of the natural biomaterials. The mechanical properties and degradation rates of these polymers can be varied to suit specific application. Processes for their production can be controlled in such a way that batch-to-batch variability is minimal compared to naturally derived biomaterials. In addition, some of these materials are equipped with the features of ECMs and ECM-bound growth factors [200]. Recent development includes networks of nanofibrillar structures formed by self-assembly of small building blocks, artificial ECM networks of protein polymers or peptide synthetic polymers with bioactive ligands, responding to cell-secreted signals [200,201]. Among synthetic polymers,

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PLA, poly(lactic-co-glycolic acid)—PLGA, poly(L-lactic acid)—PLLA, poly(D-lactic acid) (PDLA), poly(glycolic acid)—PGA, and PCL are preferred for use in OOC platforms [40,202]. We discuss some of the important ones here. PLA is a thermoplastic aliphatic biodegradable polyester derived from renewable resources such as starch, cassava, sugarcane, and tapioca roots [203]. Although not transparent, owing to its biocompatibility and biodegradation, PLA has found its main use as a support structure in OOC devices, medical implants and scaffolds, and as polymeric scaffold for drug-delivery purposes [204]. Due to the chiral nature of the lactic acid in PLA, several distinct optical isomers of it exist. Most famous of them is PLLA. PLLA has a low glass transition temperature, which is undesirable [205]. Progress in chemistry and biotechnology has led to the development and commercial production of the D-enantiomer form [206]. A stereocomplex mixture of PLLA and PDLA has a higher glass transition temperature and better visual transparency, which are expected to increase their use in the OOC platform technology [207]. PGA and PLGA are biodegradable, thermoplastic polymers and PGA is the simplest linear, aliphatic polyester. It can be prepared starting from glycolic acid by means of polycondensation or ring-opening polymerization. Because of its hydrolytic instability, however, its use has initially been limited [208]. With time, it has found its use in tissue engineering and OOC development. PGA is generally manufactured into a mesh network and used as a scaffold for bone [209,210], cartilage [211,212], and tendon [213], etc. However, for OOC platforms PGA mainly finds its use through its copolymers such as PLGA. These materials are widely used for the synthesis of absorbable sutures and are being evaluated in the biomedical field [214]. PLGA was used to coat microwells inside PDMS films to study salivary gland cell differentiation in a 3D matrigel construct [215]. PCL is a biodegradable polymer and is mainly synthesized from ε-caprolactone through ring-opening polymerization [216]. For its peculiar mechanical properties, miscibility with a large number of other polymers, and biodegradability, it has been thoroughly used for drug delivery and of bone tissue scaffolds [217]. In addition, it is preferred in surgical yarn production and biomedical applications such as drug release systems [40]. Approved by the US Food and Drug Administration, PCL has been extensively used as a bioinspired material for applications in tissue engineering, medical devices, and green chemistry, among others. Synthetic biomaterials have been studied extensively in tissue engineering applications and reviewed in detail elsewhere [122,143]. Synthetic hydrogels provide many advantageous properties related to natural materials such as that they allow the adjustment of mechanical properties according to the tissue properties. However, a major disadvantage of the use of synthetic biomaterials is the absence of cell adhesion ligands on the surface. This challenge can be addressed by chemical modification of the surface by adhesion molecules such as laminin, fibronectin, and other peptide sequences [161,218,219]. In addition, bioactive molecules with other suitable properties may also be incorporated by covalent attachment, adsorption, or electrostatic interactions to provide desired properties such as prosurvival or angiogenic properties.

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Hybrid: These biomaterials have advantages of both natural and synthetic biomaterials. For example, these biomaterials offer a wide range of chemical and mechanical products. They can be synthesized with a controllable and reproducible approach. The degradation rate of these materials can also be adjusted according to use and application. More importantly, the natural component of hybrid natural
synthetic biomaterials provides cell affinity for hybrid biomaterials [53]. Some examples of hybrid biomaterials used in engineering tissues include PEG
fibrinogen [34], PLA
chitosan
gelatin [166], and chitosan
siloxane [150]. There is an increasing interest in developing degradable hybrid polymer biomaterials with controlled properties for highly efficient biomedical applications. While there have been efforts to mimic the extracellular protein structure such as nanofibrous and composite scaffolds for bone tissue engineering and other applications [220], to functionalize scaffold surface for improved cellular interaction, to incorporate controlled biomolecule release capacity to impart biological signaling, and to vary physical properties of scaffolds to regulate cellular behavior, more research is needed in this direction before such biomaterials can go mainstream. In the development of the OOC platforms, biomaterials have an important role right from the fabrication of the chips to the biological components encapsulating cells within. Elastomers and thermoplastics have been used extensively used as building constituents for majority of the OOC platforms. Owing to their excellent transparency and biocompatibility, they are materials of choice for their ease of synthesis and noninvasive microscopic analysis of these chips. Their low hydrophilicity has led to the introduction of hydrogels as building blocks for OOCs. For hosting the biological component of OOCs, the cell-encapsulating materials should be highly biocompatible and perhaps bioactive. Natural polymers form an obvious choice. Collagens, gelatin, fibrin, HA, chitosan, and alginate are some of the polymers that have been used extensively for 3D cell culture and microfluidics in OOC devices. They have been used either in their pristine form or in chemical conjugation with other chemicals and polymers to fine-tune their crosslinking, gelation times, mechanical properties, and degradation rates depending upon the choice of the tissue to be studied or application of the OOCs. Their batch-to-batch variation and complex synthesis methods have led to the introduction of synthetic polymers. A hybrid of the natural and synthetic polymers has led to the merger of cytocompatible and OOC-directed properties in the form of hybrid polymers. Constant strides are being made in the development of these materials. In the next decade, one would expect the introduction of biomaterials that can host a wide variety of cell types and offer better vascularization and control over mechanical properties and degradation profiles. Envisioning a multiorgan OOC platform would get a boost with the development of such biomaterials in the near future.

26.10

Challenges and outlook

Over the last many years, numerous biomaterials have been adopted for the development of microphysiological systems that are capable to assist and mimic the

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function of human tissues and organs to a great extent. Such innovations display great applications in investigation of pathophysiological processes due to their high reproducibility, low costs, and strong biological relevance. The development of lab-on-a-chip devices has brought a paradigm shift in studying different diseases from both pathophysiological and drug discovery perspectives. Although further research is needed to develop fully functional humanon-chip platforms that interconnect different organs in a biologically relevant manner, current systems already facilitate better predictions of the behaviors and responses of the human body than conventional 2D models. The combination of these platforms in the future with bioprinted tissues or lab-grown organoids could form a promising strategy to overcome current limitations and possibly improve the capacities of our models in vitro. Furthermore, such systems are still evolving and can help to investigate certain research questions that were not possible to be addressed until very recently. For example, the effects of space travel (microgravity, radiation) on humans and the progress of diseases in certain geographically and environmentally rare conditions. Such research will not only save the lives of millions by offering rapid analysis in the case of an attack with chemical weapons, but also give us a sneak peak of the dangers of space when humans are planning to venture and inhabit different planets. Based on the application, OOC platforms have been fabricated from components that have expedited their use in the biomedical field. Synthetic and natural materials have been used as building components for OOC platforms. Depending upon the tissue to be studied, natural, synthetic, or hybrid biomaterials have taken a major leap in the last couple of decades for the development of the biological components of the OOC devices, leading research and applications to an unprecedented accuracy similar to the tissues and organs in their natural states. While recent advancements in the building components and cell-encapsulating materials in OOCs have provided some important insights into the development of in vitro microphysiological systems that were previously otherwise difficult to achieve, there is still a dearth in the form of biomaterials that act as native tissue microenvironments. We expect some major breakthroughs in the field in the next couple of decades in the form of development of versatile biomaterials leading to the innovations in the development of thick vascularized multiorgan-on-a-chip platforms.

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Further reading Guarino V, Gentile G, Sorrentino L, et al. Polycaprolactone: synthesis, properties, and applications. In: Encyclopedia of polymer science and technology; 2017.

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Ece Bayir1, Mert Sahinler 2, M. Mert Celtikoglu 2 and Aylin Sendemir 2,3 1 Ege University Central Research Test and Analysis Laboratories Research and Application Center (EGE-MATAL), Izmir, Turkey, 2Department of Bioengineering, Ege University, Izmir, Turkey, 3Department of Biomedical Technologies, Ege University, Izmir, Turkey

27.1

The role of bioreactors in tissue engineering

The ultimate goal of tissue engineering (TE) studies is to construct fully functional three-dimensional (3D) artificial tissues and organs by using a combination of cells, biomaterials, and signaling factors [1,2]. These signaling factors include biochemical ones, such as serum proteins, growth factors, and cytokines; physicochemical ones such as O2 tension, pH, and CO2 concentration; as well as physical ones such as mechanical cues, electromagnetic environment, and temperature. The main challenge that has kept tissue-engineered constructs (TECs) from being widely adopted to clinics is poor cell survival due to lack of proper vascularization and resulting poor integration in vivo for clinically relevant, thick constructs [3,4]. In static culture, mass transport of oxygen and nutrients, as well as removal of waste and metabolites, depends solely on passive diffusion. Depending on the cell type and the diffusivity properties of the scaffold, the diffusion for cell survival is limited to 100
200 μm distance [5], which is approximately the same size of the in vivo vascular network mesh that ensures the viability of our tissues [6]. Therefore maintenance of 3D, clinically relevant sizes of TECs, requires dynamic culture that introduces convection and perfusion, in addition to diffusion into the culture system. TE bioreactors are the keys to translate lab-grown constructs to clinically relevant, large scale, viable, and financially plausible tissues and organs by providing proper mass transport, as well as a strictly controlled culture environment for several important variables such as culture media contents, temperature, oxygen tension, and pH. A more recent, and possibly shorter term attainable role attributed to TE is the production of in vitro disease models that can serve to further improve our understanding of the cellular mechanisms of several diseases, as well as the effects of certain parameters (such as stress factors or environmental variables) on disease progression [7
9]. These models are extremely useful for developing novel therapeutic approaches, performing high-throughput preclinical drug screening, as well as personalizing treatment options by precision medicine approach [10,11]. These in vitro disease models also serve the purpose of reducing animal testing in consideration of the three Rs (reduction, refinement, and replacement) [12,13]. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00018-0 © 2020 Elsevier Ltd. All rights reserved.

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In order to obtain functional tissues/organs in vitro, native micro-environments of the tissues and organs should be mimicked, since the living cells and tissues experience variety of physical stimuli along with the chemical ones in physiological conditions through other cells and extracellular matrix (ECM). Tissue development in vivo as well as in vitro relies on spatial and temporal organization [14,15]. Conversion of these physical stimuli into biochemical activities through strainresponsive integrins at cell
ECM interface, ionic channels and cell
cell adhesion proteins are defined as “mechanotransduction” [16
19]. Integrins, which are thought to be mechanosensors, can go through conformational changes that regulate affinity for the ECM proteins [20]. Also, change in ECM properties such as stiffness and adhesiveness are known to affect the regulation of Yes-associated protein/transcriptional coactivator with PDZ-binding motif, transcriptional regulators in Hippo pathway activity, in turn affecting multiple aspects of cell metabolism [21]. Cell
cell mediated mechanotransduction can occur through various mechanosensing structures, such as a-catenin, an intracellular protein, desmosomes, and transmembrane receptor Notch—transmembrane ligand Delta interaction [22]. Furthermore, mechanical signals originated from outside the cells can navigate itself to the chromatin complex through the actin, tubulin and intermediate fibers, and lamins [23
25]. The term “mechanoepigenetics” has been proposed recently to describe the effects of mechanical signals on epigenetic activity and chromatin [24]. These biochemical activities originated by mechanical signals through various pathways control critical aspects of cell behavior, such as morphogenesis, proliferation, migration, differentiation, matrix production, and protein expression in general [21,26,27]. TE bioreactors can provide a great level of control over cellular mechanotransduction by utilizing various mechanical forces in order to achieve a well-mimicked, tissue-specific environment for cell survival, proliferation, migration, and differentiation [28,29]. A bioreactor is simply defined as a “device” or “system” that maintains biologically active conditions for cultivating an organism and/or carrying out a biological reaction. Besides TE, bioreactors have been also utilized for the expansion of stem cells prior to seeding on scaffolds. Embryonic stem cells and mesenchymal stem cells (MSCs) that are particularly important for TE due to their renewal and differentiation potentials have been cultivated and proliferated effectively in bioreactors, with or without the utilization of microcarriers [30
32]. It is important to reach large cell numbers without losing stemness characteristics over long culture periods for successful TE applications, and this is practically impossible in static culture systems [33,34]. This approach has been borrowed from initial production purpose of bioreactors for classical industrial processes, such as fermentation, and the very few studies available in the literature have used stirred bioreactors and roller bottles, without paying particular attention to culture conditions. Further studies are needed for optimization of culture conditions such as hydrodynamic properties and oxygen concentration. Since stem cells form cellular aggregates in suspension cultures that, in turn, lead to differentiation, utilization of suitable microcarriers of 100
250 μm size to avoid agglomeration have been preferred [33,35,36]. For some particular TE applications, such as heart valves and vascular tissues, that have to bear significantly high mechanical loading immediately after

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implantation, bioreactors also provide a means to validate the performance of the TECs under physiologically relevant loading before implantation and their preconditioning and maturation [37]. Since TE bioreactors enable the control of environmental conditions, such as pressure, oxygen tension, temperature, nutrient supply, waste removal and pH, as well as mechanical and physical cues, while allowing aseptic operation (e.g., mounting, feeding, and sampling) [38,39], for TE applications in a realistic and larger scale, bioreactors are a must. This chapter will provide an overview of various kinds of bioreactors that is specifically developed for TE applications, particularly emphasizing the design parameters for improving mass transport and for application of tissue specific mechanical cues.

27.2

Bioreactor configurations

27.2.1 Stirred bioreactors Stirred bioreactors are systems that generally consist of a glass or biocompatible transparent polymeric material body, which contain an impeller (stirrer), side arms for cell and medium addition and aeration, and a number of probes that can be added to the interior of the bioreactor if desired (Fig. 27.1). Generally, different types of stirred bioreactors are used in cell culture to achieve high cell density in suspended cultures and to obtain high amounts of cellular products in general [40,41]; however, these bioreactors are also frequently used in TE studies [42
44]. Spinner flask, which is one of the most commonly used types of bioreactors in TE, is a simple version of the stirred bioreactors, where the stirrer is driven magnetically. Spinner flasks are more economical and easier to use than other types of bioreactors.

Figure 27.1 A typical stirred bioreactor.

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In stirred bioreactors the culture medium is homogenized by hydrodynamic forces, and the concentration gradient within the medium is deteriorated by mixing [45]. The mass transfer takes place by both diffusion and convection through mixing [46]. Therefore the nutrient and oxygen transfer to the cells at the inner part of the tissue is increased. The most important variable in stirred bioreactors is the mixing speed. Although mass transfer is enhanced by turbulence and the eddies caused by high mixing speeds, high shear stress caused by high speeds can also cause damage to the cells [47]. In addition, high mixing speed may result in loss of ECM in the tissue [48]. For this reason the optimum mixing speed should be determined for each specific culture/scaffold combination. Generally, spinner flasks are 120 mL in volume and are mixed between 10 and 80 rpm [49]. Stirred bioreactors have a magnetic or motor driven stirrer and needle or wire structures, where tissue scaffolds are hung [50] (Fig. 27.1). The cells can be seeded by dynamic seeding method after the tissue scaffolds are submerged in the media. At this point the mixing speed is again crucial, because the cells can be collapsed and the cell attachment to the scaffold can be reduced at very low speeds. In addition to the dynamic seeding method the cells may also be preseeded under static conditions and transferred with the scaffold into the spinner flask [51]. Half of the used culture medium in spinner flasks is changed every 2 days, in general [52]. The oxygenation of the medium and the gas exchange takes place from the top of the bioreactor. Spinner flasks are mainly used in bone and cartilage TE but have recently begun to be used in new fields, such as cardiac TE [28]. In bone TE studies using spinner flasks, it has been observed that the proliferation of human MSCs (hMSCs), alkaline phosphatase (ALP) production, and cell differentiation increase by mechanical stimulation induced by the reactor [53]. It is also shown that cell distribution, matrix deposition, and mineralization within the scaffold increase [54
56]. In cartilage TE studies, cartilaginous matrix formation is enhanced under higher shear stress using chondrocytes in spinner flasks. Moreover, mechanical stimulation induces glycosaminoglycan (GAG) and collagen II production from cells [57,58]. Similarly, in the studies of cardiovascular TE, increase in the proliferation rate of cardiac myocytes was observed due to the shear stress applied in spinner flasks [59]. The main limitations of stirred bioreactors are high shear stress and limited mass transfer. In spite of the increase of cell proliferation and differentiation rate in shear stress, cellular damage, such as apoptosis or lysis, also increases at higher speeds [60]. Mass transfer has been shown to be limited to ,1 mm in spinner flasks as bone formation by rat and hMSCs was observed only on the outer layer of the scaffolds, while observing a necrotic core within the scaffolds [61,62]. Therefore oxygen and nutrient transfer to the inner parts of the scaffold cannot be adequately enhanced in large scaffolds by stirred bioreactors.

27.2.2 Wave bioreactors Wave bioreactors are disposable bag bioreactors that are generally produced using transparent flexible polymers. The bag bioreactor is placed on a rocking base for

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Figure 27.2 Schematic drawing of a wave bioreactor.

wave-induced agitation (Fig. 27.2). This base may be a closed system, which has controlled environment, or it may be an open system and placed in an incubator. The velocity, angle, and rate of the rocking have to be adjusted according to the needs and the size of the cultured constructs. These types of bioreactors have been used in the field of TE only in the last decades, although they have been used in other culture systems with the aims of reaching a high number of cells using a minimum amount of medium, increasing the production efficiency of antibodies, vaccines, and therapeutic proteins [63]. In this type of bioreactors, micro-carriers are usually used to produce microscale tissues. Wave bioreactors can be used as a batch, fed-batch, or continuous systems. Feed bag and harvest bag are connected to provide continuous feeding of fresh medium and removal of the used medium when the bioreactor is used continuously [64]. The most preferred wave bioreactor process is a fed-batch system in order to prevent the potential for substrate/product inhibition, which can be observed in batch systems, and ease of use. The waves generated by rocking increase the mass transfer and keep the cells/aggregates suspended [65]. Wave bioreactor’s geometry, filling level, rocking angle, velocity and rate, and the viscosity of the medium change the shear stress applied to the cells. Mass transfer is increased by low shear stress, which is obtained by rocking without mechanical mixing. It is also a bubble-free system and, thus, it is suitable for sensitive cells, such as stem cells. Because they are disposable bioreactors, the risk of contamination is low in these systems. In a recent study the effect of the wave bioreactor on the self-assembled 3D aggregate formation of hMSCs and its effect on the size distribution were investigated. The kinetic energy of the cells increased by the wave generated as a result of rocking, and increased collision frequency supported the aggregate formation [66]. In systems with high mixing speed, cells are separated from aggregates when shear stress is higher than the cell adhesion forces. Since very low shear stress is generated in the wave bioreactors, larger aggregates can be obtained compared to the systems that have high mixing speeds. Preservation of stemness characteristics, migration, and secretion rates of hMSCs in rocking culture were determined to be higher than static culture [66]. In another study performed on ECM-coated polycaprolactone

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(PCL) microcarriers using hMSCs, the bone formation was investigated. Since cell aggregates were .2.7 mm, they used a wave bioreactor instead of static culture in order to avoid necrotic areas. Wave bioreactor was chosen instead of stirred bioreactor for this research, due to the fact that PCL microcarriers, which have a high density of 1.09 g/cm3, require high mixing speed. It was determined that the cell viability and mineralization rate were higher in dynamic culture compared to the static culture [67]. In another study performed on electrospun poly(lactic acid) (PLA) scaffolds using CP5 chondrocytes, the amount of lactate dehydrogenase (LDH) released by the cells and cell viability in static culture and wave-induced agitated culture were compared. It was stated that the cell viability increased with wave-induced agitation, since the transition of the medium through the porous scaffolds in the wave bioreactor was higher than the static culture. In addition, it was also shown that the amount of LDH released by the cells in the wave agitated culture was three times less than the static culture, which indicates lower membrane damage. The researchers concluded that the rocking system protect cells from negative conditions of static culture [68]. The biggest disadvantages of wave systems are the need of optimization for the rocking rate, velocity, angle, and bag filling level for each specific study, and during the study as the cell density and matrix deposition increases in time, and relatively high cost due to the fact that these are disposable and some systems contain sensors [65].

27.2.3 Parallel-plate bioreactors and parallel-plate flow chamber bioreactors Parallel-plate flow chambers (PPFCs) are constructed by two parallel plates with an opening on each side of the plates to create pressure difference to form laminar flow of the culture media between the plates [69] (Fig. 27.3). The top plate is

Figure 27.3 Schematic drawing of a parallel-plate flow chamber bioreactor.

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usually a gas permeable, liquid impermeable membrane, and the bottom plate is glass or a biocompatible polymer. Well-defined shear stress originated from the laminar flow that mimics in vivo conditions is extensively utilized in investigation of cell mechanotransduction studies using PPFCs [70
72]. Laminar flow is characterized by low Reynold’s number, which is directly proportional to velocity and channel thickness; thus it is harder to maintain laminar flow as the thickness of the chamber increases. Therefore it is hard to utilize parallel flow bioreactors for thicker TE constructs. Besides, as the tortuosity of the flow path changes while the flow passes through the porous scaffold, the laminar character of the flow that the cells within the scaffold are subject to is questionable. Therefore although they are important tools for the investigation of cellular mechanotransduction, utilization of parallel-plate bioreactors for TE is limited to membranes or very thin constructs. In one of the few examples of the use of PPFCs for TECs, effects of 29.5 N/m2 of shear stress on mouse 3T3 fibroblasts seeded on plasma treated poly-(L-lactide) (PLLA) membranes were tested in order to evaluate the efficiency of plasma treatment in terms of cell retention, morphology, and shape factor [73]. In another study a parallel-plate bioreactor was used to apply fluid flow-induced shear stress to tissue-engineered articular cartilage. Primary bovine articular cartilage cells showed increased collagen type II production and matrix stiffness under the effect of flowinduced shear [74]. Effects of shear were also investigated on endothelial cells seeded on microgrooved polymer surfaces to get a good understanding of proper surface microstructure requirements for endothelial cell alignment [75].

27.2.4 Rotating wall vessel (reduced gravity) bioreactors Rotating-wall vessel (RWV) bioreactors were developed by the National Aeronautics and Space Administration (NASA)—Johnson Space Center, United States in order to create microgravity to protect cell cultures from high mechanical forces that occur during takeoff and landing of spacecraft [76]. There are three different types of RWV bioreactors. The original one is the slow turning lateral vessel (STLV) bioreactor that consists of two concentric cylindrical structures. Since the vessels are filled completely with culture media, there is no gas
liquid interphase. The oxygen transfer takes place through silicone membranes in the inner cylindrical structure (Fig. 27.4). The outer vessel is generally made of transparent biocompatible polymeric materials. The space between the two concentric cylinders includes microtissues, scaffolds, or microcarriers in the culture medium [64]. STLV bioreactors are fed-batch systems; however, rotating-wall perfused-vessel bioreactors, which are similar to STLVs, have a circulation system that allows periodic replacement of the culture medium [77]. Another type of RWV bioreactor is high aspect rotating-vessel bioreactors. In this system the gas transfer is increased by reducing the rotational speed in the shorter and larger bioreactor vessel and by a wider membrane [78]. In RWV bioreactors the vessel is rotated at certain speed to balance the gravitational force, the hydrodynamic drag force, and the centrifugal force [79]

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Figure 27.4 RWV bioreactors: (A) STLV bioreactor, (B) HARV bioreactor, and (C) RWPV bioreactor. HARV, High aspect rotating-vessel; RWPV, rotating-wall perfused-vessel; RWV, rotating-wall vessel; STLV, slow turning lateral vessel.

Figure 27.5 The gravity force (Fg), centrifugal force (Fc), and the hydrodynamic drag force (Fd) can be balanced by the rotational speed of the RWV bioreactors; thus the tissueengineered constructs remain in suspension. RWV, Rotating-wall vessel.

(Fig. 27.5). In this way the engineered tissues in the bioreactor vessel can remain suspended in a state of “free fall” [47]. The rotation speed should be increased as the tissue grows, in order to maintain the balance between the forces. By the effect of rotation, sedimentation of the tissue fragments is prevented, while the concentration gradients are disrupted by the dynamic laminar flow. Moreover, since the inner structure of RWV bioreactor does not include a stirrer, the cells are protected from excessive shear stress and turbulence [80]. RWV bioreactors also allow 3D organotypic cultures and engineered tissues to have more physiologically in vivo
like microenvironments. In RWV bioreactors, long-term cartilage and bone TE have been studied frequently. In a chondrogenic differentiation study with hMSCs, microgravity was used to reduce hypertrophy [81]. In another chondrogenic differentiation study with rabbit bone marrow
derived MSCs (BMSCs) in polyglycolic acid (PGA) scaffolds, it was shown that simulated microgravity induce chondrogenic differentiation and increase the production of collagen type II and aggrecan by cells; thus a more effective tissueengineered cartilage is obtained compared to in vitro static culture [82]. In bone TE studies using isolated osteoblasts from Sprague-Dawley rats in bio-derived bone scaffolds, it was shown that the cultivation of cells under RWV bioreactor conditions enhanced the mineralized nodule formation and ALP expression [83]. A similar

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result was obtained when rabbit BMSCs and MSC-derived endothelial cells were cocultured aiming to obtain vascular bone tissue in a RWV bioreactor. Enhanced cell distribution, proliferation, and differentiation rates were obtained by the mechanical stimulation created by the RWV bioreactor and the interaction of the two types of cells in the porous PLA-based scaffold, and vascular structures were observed in engineered bone tissue [84]. In a bone TE study with human fetal MSCs the cells were incubated in a biaxial RWV bioreactor after the cells were seeded on the PCL
tricalcium phosphate (TCP) scaffolds. It was shown that cAMP, TGF-β1, and nitric oxide productions, which are related to bone tissue repair mechanism, were upregulated compared to static culture in biaxial RWV bioreactor, and high cell viability, ALP production, and mineralization rate were obtained [85]. It has also been shown in several studies that simulated microgravity promote multipotential differentiation capacity of MSCs [86,87]. In addition to these studies, different tissues such as kidney, liver, lung, connective tissue, pancreas, bladder, eye, blood vessel, small/large intestine, tonsil, vagina, and adipose were also studied [88,89]. It was observed that simulated microgravity increased insulin release in the research performed on PGA scaffold with pancreatic islet cells in RWV bioreactors [90]. In a recent study on differentiation of pluripotent stem cells to retinal organoids, differentiation of the cells was induced as a result of cultivation in RWV bioreactors by mass transfer enhanced microenvironment and biophysical stimulation, and larger organoid sizes were obtained compared to static culture [91]. RWV bioreactors are also used frequently in cancer models, such as colon, prostate, ovarian, and breast, since they allow forming tumor-like 3D cell aggregates with low shear stress [92]. Since RWV bioreactors induce self-organized spheroid formation within a few hours, different tumor models can be produced using these bioreactors [93]. It is possible to use these models in prescreening pharmaceutical agents and drug testing studies and use them as reference specimens for immunostaining in the surgical pathology laboratory, because the tumor models generated are more in vivo
like than static models [94]. In general, rotating-wall bioreactors are suitable for the culture of fragile tissues, due to their low shear levels, compared to fixed wall bioreactors such as the stirred or parallel-plate bioreactors. The disadvantages of the RWV bioreactors are that the optimization of the culture conditions can take a long time [88]. Another deficiency reported with the RWV bioreactors was the nonuniform growth of the tissue within the construct [95]. Moreover, cells can be damaged due to possibility of scaffold collision to the bioreactor wall, and due to insufficient mass transfer that is limited to the scaffold surface [96]. Osteocalcin and ALP expressions in rat and human MSCs have been shown to be lower in rotating wall bioreactors compared to spinner flask cultures, either due to lower shear stress and/or scaffold collision [62,97].

27.2.5 Strain bioreactors Strain bioreactors are developed to apply direct mechanical strain to cultures of mechanically responsive tissues, such as bone, tendon, and cardiac tissues [44,98
101].

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The direct mechanical strain can be applied in forms of stretching, compression, and bending through a strain bioreactor [55]. In stretch bioreactors (Fig. 27.7A) the tensile force, for 3D constructs, is usually transferred through clamps that are attached to the scaffold, and computer controlled linear actuators are used for utilization of stretch [102]. However, design of these clamps can vary greatly depending on intended application, since clamping should induce minimal damage to the construct during loading. Several different types of clamps were designed to optimize scaffold mounting, such as spiral grips [103] and attachment hooks [101], which are particularly used for thicker constructs, and grip pins [104] and standard clamps [105] for thinner constructs, as it is crucial that the clamps do not induce cracking or tearing in the scaffold where it is mounted (Fig. 27.6). The effect of cyclic stretching is known to induce cell alignment perpendicular to the direction of stretch on the construct, which may provide similarity to the native tissue of interests [101,106]. As reported by Burk et al., short-term (15
60 minutes intervals) uniaxial cyclic stretching of 2% strain at a 1 Hz frequency induced the tenogenic differentiation of mesenchymal stromal cells cultured on decellularized tendon matrix scaffolds [105]. In another study, tenocytes seeded onto acellularized rapid flexor tendons were subjected to cyclic stretch force of 1.25 N for 5 days. The ultimate tensile strength of the loaded constructs was reported to significantly increase as opposed to the unloaded group. Furthermore, the elastic modulus of loaded constructs was also found to be significantly higher compared to the control group [107]. In a study regarding myocardial TE, it was shown that 4 hours 10% uniaxial stretch at 1 Hz protocol applied on myocardial constructs induced an increase in contractile force along with the myofibril alignment [101]. Number of studies have reported that degree of alignment depends on the stretch magnitude, frequency, and waveform [108
110]. Stretching can also be biaxial or equiaxial (Fig. 27.7A), which represent the physiological conditions more in tissues such as peritoneum, skin, and aortic valves

Figure 27.6 Illustration of (A) spiral grips, (B) attachment hooks, and (C) standard clamps.

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Figure 27.7 Examples of strain bioreactors: (A) stretch bioreactors, (B) compression bioreactors, and (C) bending bioreactors.

[111
113]. In a study where fibroblasts seeded on collagen gels were used to investigate the effect of dynamic stretch, three different cyclic equiaxial stretching conditions, such as constant 7% and 20% strain and increasing strain starting with 7% followed by 15% and 20% strain at 1 Hz each for 2 days were applied to the gels. mRNA levels of collagen type I and collagen type III were found to be higher in increasing strain group compared to static control group and constant 7% strain group. However, no difference in terms of fibronectin expression level between the groups was found [114]. In another study, Lei et al. utilized a novel biaxial stretch bioreactor to investigate the extracellular remodeling in aortic valves. Isolated porcine aortic valve cusps were subjected to 15% and 10% radial and circumferential, respectively, cyclic biaxial stretch for 14 days. The biaxial stretch increased the collagen synthesis, which results in improved maintenance of native ECM structure [115]. Studies regarding tissues that experience compression in the native environment, such as cartilage in the knee joints and bones, can be stimulated by compression strain bioreactors (Fig. 27.7B). A simple compression bioreactor generally utilizes a motor that can apply linear motion and a control mechanism that provides an ability to choose different magnitudes and frequencies [116]. These types of bioreactors can be developed to provide both dynamic and static loading, so they can be manipulated for different types of applications [29]. Studies regarding cartilage generally utilizes a dynamic loading, since static loading may have negative effects on the cartilage formation mostly due to limited mass transport [117
119]. The effects of dynamic compression with insulin-like growth factor I (300 ng/mL) on bovine patellofemoral cartilage explants were investigated through low amplitude, sinusoidal (2% strain, 0.1 Hz) compression. The combined effect of compression and growth factor increased the 3H-proline incorporation by 180% and 35S-sulfate incorporation by 290% compared to the cartilage under basal conditions. Dynamic

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compression alone increased the 3H-proline and 35S-sulfate incorporation by 30% and 120%, respectively [120]. There are several other studies on cartilage TE showing the stimulating effects of compressive strain on scaffold elastic modulus [121], sulfated GAG, and hydroxyproline content [122]. A study utilizing another compression strain bioreactor revealed that 25% strain at 0.05 Hz enhanced the expression levels of genes related to osteogenesis on bone marrow
derived stem cells [123]. The effect of cyclic compression on MSCs seeded on PCL/β TCP scaffolds was investigated in terms of osteogenic differentiation. The constructs were cultured on static conditions for 24 hours; then were divided into two groups as static and cyclic compression (0.22% strain at 1 Hz). Osteogenic differentiation markers, osteonectin, COL1A1, and osteocalcin were found to be upregulated in cyclic compression group compared to the static group. Also, an increase in ALP activity in cyclic compression group was reported [124]. In another study regarding osteogenic differentiation, BMSCs were seeded onto bovine decalcified bone matrix, and the constructs were subjected to cyclic compression (10% compression at 0.5 Hz triangular wave) for 2 hours/day. The bioreactors used to dynamically culture the constructs also utilized a substance exchanger system with feedback control for pH and pO2. The bioreactor system’s ability to keep the pH, pO2, and nutrient level in balance, as well as removal of metabolic waste were verified. The proliferation of cells was increased in the dynamic group as opposed to the static control group. Also, the calcium levels of constructs cultured under compression were increased starting at day 10 as opposed to static control group [125]. In a unique study conducted by Lujan et al. a novel bioreactor system that can apply consistent and independent 0.1
10 N compression at 1
10 Hz to six constructs was used. The bioreactor also had the ability to detect changes in mechanical properties of the constructs during the incubation period, which allowed researchers to efficiently evaluate the biomechanical development of the constructs. The system successfully detected the 2 kPa reduction in equilibrium modulus of collagenase sensitive polymers that were subjected to collagenase [126]. More studies utilizing perfusion and flexure on top of compression bioreactors are discussed in combined bioreactors section of this chapter [103,127
130]. Bending bioreactors have been developed to provide more similar environment to the tissues that are usually found under more complex strain conditions in the native environment, such as long bones and heart valves [131,132]. Bending causes stretch on one side of the neutral axis, while inducing compression on the other side, imposing a pressure gradient that reaches maximum tensile and compressive values at the two surfaces of the scaffold [133]. These bioreactors can be designed to apply three-point or four-point bending depending on the application (Fig. 27.7C) [133,134]. As in other types of strain bioreactors, these also utilize computer-controlled linear actuators that can modulate the vertical position and velocity of the load cell, which can transfer the bending force through a shaft [133]. In a study that utilized a four-point bending bioreactor used rat MSCs seeded on plastic strips to achieve increased cell proliferation and expressions of Ets-1, Cbfa1, and ALP genes related to osteogenic differentiation. Cells on the plastic strip were subjected to 2000 microstrains (0.2%) at 0.5 Hz. The mean cell number in stretched

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group was found to be higher than the unstretched control group. ALP activity in stretched group was detected to be threefold higher than the control group; and Ets1 and Cbfa1 mRNA expression levels were also increased in stretched group as opposed to the control groups [135]. In another study a three-point bending bioreactor was utilized on ovine smooth muscle cells (SMCs) seeded on PGA/PLLA scaffolds to apply cyclic bending with 6.35 mm displacement at 1 Hz for 3 weeks to develop tissue-engineered heart valves (TEHVs). The collagen amount in wet weight increased 63% in bending group compared to the static group [128]. Another study regarding heart valves were conducted with porcine aortic valvular interstitial cells seeded on poly(glycerol sebacate) (PGS) scaffolds cultured under cyclic stretch (10% strain,1 Hz) and flexure (0.15 mm21 change in curvature). Significant increase in DNA (191%) and collagen (423%) content were reported in the cyclic stretch group compared to the static group; however, no significant difference, in terms of DNA and collagen content, were reported between the cyclic stretch and flexure group [104]. Utilization of hydrostatic pressure (HP) bioreactors is another way to apply mechanical stimulation to the TECs. These types of bioreactors generally involve pressure pistons or pumps, filters for aeration, one-way valves, and chambers that can withstand the applied pressure [4] (Fig. 27.8). Many different types of cells in tissues and organs, such as osteocytes and chondrocytes, are subjected to HP in vivo [57,136]. The effects of cyclic HP on growth of chick femur skeletal cells seeded on hydrogels were investigated, where the constructs were subjected to HP of 0
279 kPa at 1 Hz for 1 hour. The stimulated gel constructs were reported to be denser than the control group due to increased mineralization [137]. There are contradicting reports in the literature on whether cyclic or static HP is more stimulating for cartilage TE. Stimulative effects of cyclic HP on chondrogenic differentiation were demonstrated on BMSCs seeded on dehydrated collagen I scaffolds that were subjected to 1 MPa pressure at 1 Hz for 4 h/day for 10 days. Alcian blue staining was visually stronger at the pressurized group as compared to the unpressurized control group. Also, mRNA expression levels of aggrecan, collagen II, and Sox9 were significantly higher in the pressurized group compared to the control group, indicating enhanced chondrogenic differentiation [138]. Another study in favor of cyclic HP on tissue-engineered cartilage has shown increased tissue formation of human nasal chondrocytes in gellan gum hydrogels under cyclic HP loading of 0.4 MPa compared to static HP and control static cultures over 3 weeks [139], while static HP of 2.8 MPa was shown to cause higher GAG production of bovine

Figure 27.8 A simple hydrostatic pressure bioreactor system.

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chondrocytes on collagen sponges over cyclic HP, although static strain is not representative of physiological cartilage loading [140]. Although design and manufacture of strain bioreactors that can apply various forms and magnitudes of strain is possible, the optimal magnitude, frequency, and duration of application of strain bioreactors have to be determined case-specifically, taking into account the scaffold type and shape, as well as the increased cell number, decreased porosity, and increased elastic moduli due to deposited ECM during the culture time.

27.2.6 Perfusion bioreactors A perfusion bioreactor is a closed-loop system, which allows medium to transfuse through the TEC directly with modifiable pressure, flow rates, and types. Thus the system helps to create a controlled microenvironment to overcome mass and gas transport limitations, enables homogenous and dynamic cell seeding, while providing biomechanical stimulation to the TEC [141
143]. The system comprises some basic compartments: a body, where flow enters the system via inlet and passes through fixed TEC and leaves from the outlet to the pump that circulates the enriched media to feed the TEC (Fig. 27.9). Since the medium is forced to flow through the pores of the scaffold, it creates a pressure that induces shear within the internal walls of the scaffold depending on the pore size, shape, and the degree of porosity (Fig. 27.9), as well as the flow rate and regime [143
145]. Flow type and rate can functionalize and increase the viability of TEC, while they can be damaging to the cells and the scaffold on high flow rates [143,146]. Perfusion bioreactors are utilized for cell seeding into the scaffolds frequently. With adequate flow the cell suspension can be perfused through porous scaffold

Figure 27.9 The most prevalent scheme of a perfusion bioreactor. The pump circulates the medium that has been taken from the reservoir and richened with oxygen to the bioreactor body, where TEC had been housed. TEC, Tissue-engineered construct.

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and driven to migrate and attach homogenously, while efficient mass transport to the TEC is provided as compared to conventional cell-seeding techniques [141,147,148]. Although most perfusion bioreactors are unidirectional, a bidirectional perfusion bioreactor was shown to result in higher seeding efficiency (75% 6 6%) of scaffolds compared to static (57% 6 5%) and spinner flask seeding (55% 6 8%) [149]. For a functionalized TEC, cell viability and the scaffold integrity are the key factors; therefore, efficiency of the perfusion system depends on optimizing scaffold design and flow dynamics. Different flow regimes, such as unidirectional laminar, pulsatile laminar, turbulent and oscillating flow can be used to deliver shear stress similar to physiological levels [150]. The scaffolds should contain at least 70% porosity to endure the effects of flow and to increase permeability. Also, pore size and architecture are crucial for nutrient and oxygen transport, while allowing uniform spatial distribution [148,151,152]. Perfusion bioreactors are most commonly used for bone TE applications, since the mechanical loading regime most closely simulates the in vivo scenario. When physiologically loaded, fluid flow is induced within bone tissue leading to communication, differentiation, and mineralization of osteoblasts and osteocytes [153]. Increased differentiation, bone specific protein expression and mineralization were shown on various scaffolds by perfusion bioreactors [154
158]. Bone-related gene expression and biological apatite crystal formation was shown to be upregulated only by 1 mL/min flow that biomechanically induced cells, while providing nutrients to the scaffolds by a perfusion bioreactor. A homogenous spatial distribution of human osteoblasts within the Ca-alginate scaffolds was achieved by 14 days of perfusion, and due to the induced shear, the osteblasts were differentiated and functionalized into biomineralized osteoclusters that can be transferred to patients as cell therapy for bone regeneration [159]. Several other types of cells were cultured with success in perfusion systems for the production of TECs, including MSCs [160], chondrocytes [161,162], keratinocytes [163], hepatocytes [164], and cardiomyocytes [165,166]. Jung et al. cultivated a scaffold-free blood vessel for preclinical drug screening in a perfusion bioreactor. With the help of 2 mL/min flow (physiological shear stress of 6.8 dynes/cm2) for a week, human endothelial progenitor cells strongly attached to fused hMSC sheets. Mechanical stimuli differentiated hMSCs into smooth muscle periphery and firm structure of scaffold-free tissue-engineered blood vessel endured the burst pressure above 200 mmHg [167]. Dermagraft is an example of commercialized skin TE product that is produced by a perfusion system. 3D printing and bioprinting are emerging technologies that have been used to produce large living tissues. Their strongest advantage is their conformity for large sizes of TECs, and even whole-organ TE; and this technique has enabled to mimic natural microenvironments accurately in terms of physical, biochemical, and complexity levels [168
170]. The maturation of computationally designed and printed cell and scaffold constructs into a whole tissue is possible with perfusion bioreactors [171
173]. In a recent study a 3D bioprinted dm3 scaled tissue maturated with connection to macrovascularization (bioprinted internal structures) and microvascularization

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(cells’ self-organization via flow) by a perfusion system. The group has bioprinted fibroblasts and human dermal microvascular endothelial cells within a gelatin, fibrinogen, and alginate-based bioink and cultivated the TECs for 12 days in a selfprinted silicone bioreactor with continuous flow rate of 300 mL/h. They showed induced collagen production within TEC compared to static culture and the presence of lumen-based microvascular organization related with hydromechanical stress [174]. Another application of perfusion bioreactors are decellularization, which is an emerging technique to remove the cells from the ECM of native tissues to obtain 3D organ/tissue scaffolds for TE. By the help of perfusion bioreactors, various tissues and organs have been successfully decellularized [175,176]. Target organs can be placed and fixed into the chamber to effectively remove cells using chemicals, such as detergents, salts, enzymes, and/or by the help of the physical effects caused by the flow, while reducing immunogenicity, preserving the ECM composition, architecture, bioactivity, and mechanics [177]. Due to stiff structure and low permeability of the native grafts, those solvents and other agents used for decellularization cannot infuse to the organ efficiently in static conditions. Perfusion systems overcome decellularization problems both by improving the transport of chemicals and rinsing the cells away with the superior penetration virtue [141,178]. Sarig et al. managed to decellularize thick porcine heart tissue (10
15 mm) efficiently without losing mechanical properties [179]. Compared to other techniques (stirring, sonication, etc.), trypsin/triton-based perfusion provided efficient penetration to obtain thicker acellular slabs while preserving ECM components. Group also used perfusion bioreactor for functional angiogenesis in vitro with those porcine cardiac ventricular ECMs (pcECMs). With static seeding, cells penetrated only into 100 μm depth of the 1.7 mm thick pcECM scaffolds due to poor diffusion. After dynamically seeding and cultivating cells with 40 mL/min flow for 7 days, they managed to fold depth of cell clusters four times and shorten cultivation time for 2 weeks [180]. The researchers find the optimum flow conditions in perfusion systems usually by trial and error, which costs significant amounts of time and money. Computational modeling systems that can predict induced shear stress levels depending on flow conditions, cell/ECM concentration, and scaffold properties can help proceed much faster and reliably for experimental design. In a recent study using computational fluid dynamics (CFD), researchers modeled the effect of wall shear stress (WSS), which leads to bone formation, at the scaffold surface with varying flow rates (0.1
12 mL/min). Depending on the scaffold microstructure (spherical/rectangular pores, pore diameters, and porosity percentile), they calculated the optimal flow range for maximizing the mineralization in different scaffolds for perfusion systems as ranging from 0.17 to 1.66 mm/s [143]. In another study, Lembong et al. developed a pillar-based fluidic perfusion culture chamber to enhance osteogenic differentiation of hMSCs cocultured with human umbilical vein endothelial cells (HUVECs). For tuning the microenvironment precisely, 3D printed photo-reactive acrylate had been used as the scaffold. With the help of 3D printing technology, manipulation of pillar geometry caused different flow patterns inside the scaffold. Using flow-induced shear stresses that ranged from 1 to 1000 MPa,

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spatially directed flow (10 mL/min) enhanced ALP expression and cell proliferation. The two cell populations were self-assembled and separated within the scaffold by B200 μm distance, and calcium signaling between the two cell types was modulated [160].

27.2.7 Hollow-fiber bioreactors Cells present in all tissues in vivo are not more than 100
200 μm away from capillaries, so that mass transport of nutrient and oxygen, as well as waste removal can be achieved through diffusion [181]. This phenomenon cannot be easily mimicked in vitro. For this reason, bunched hollow fibers acting like artificial capillaries are placed in a cartridge so as to mimic the natural microvascular structure (Fig. 27.10). Different cell types can be seeded within the extra-capillary space, as well as outside and/or inside the hollow fibers. Hollow fibers are generally 50
3000 μm in diameter and can be produced from different materials [182]. Natural/artificial, biodegradable/nonbiodegradable polymers are generally used as fiber materials [183]. Polymer hollow-fiber production is most commonly performed by wet phase inversion spinning method. Apart from this method, extrusion, electrospinning, dip coating, and melt spinning methods can also be used [184,185]. However, melt spinning is the least preferred method due to the high-temperature requirement and the difficulty in porosity control [182]. The properties of hollow fibers are dependent on many variables, such as polymer type, porosity, pore diameter, wall thickness, molecular weight cut-off (MWCO) value, permeability, physicochemical properties, type of chemicals used in the production method, and whether a coating is applied or not [185
189]. In addition to these variables, parameters such as the number of hollow fibers, and the relative location of these fibers and the cells should be considered specifically for each study. For example, nonbiodegradable polymers are generally selected if cell expansion or an external artificial organ application is to be performed, or if the aim is to obtain durable tissues [190,191]. If it is desired to increase the mass transfer to the cells, while protecting them from the mechanical effects, the cells can also be planted in the outer walls of the hollow fibers or extra-capillary space instead of the inner wall of the hollow fibers [192,193]. In the extra-capillary space a cell embedded or pure hydrogel can be introduced to mimic the ECM, or static or

Figure 27.10 (A) Bunched hollow fibers placed in a cartridge. (B) The cross section of hollow-fiber bioreactor.

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dynamic culture media may be present in this region [194]. The MWCO ratio of the membrane can be adjusted to protect the cultured cells in the hollow fibers from the immune system cells within the host if the engineered tissue is to be implanted [195]. Furthermore, the cocultivation of two different cells without a common medium is allowed in this system easily. The polymer properties can be adjusted as desired and the media in and out of the fibers can be easily separated from one another [194]. Hollow-fiber bioreactors have a great surface area/volume ratio compared to other bioreactor types; therefore cell attachment and proliferation rate is very high because of the improved mass transfer [196]. Hollow fibers also provide compactness and high external HP resistance, and they require small amount of culture supplements [197,198]. Because hollow-fiber bioreactors mimic the natural capillary system and it is easy to obtain heterogeneous tissues with these reactors, they are used for several TE studies, such as bone, cartilage, kidney, and liver, but they are most appropriate for tissues that require tubular shapes, such as blood vessels, intestine, and urinary organs [189,193,199
202]. One of the most popular areas using these bioreactors is blood
brain barrier (BBB) research [203
205]. In 2D systems, it is difficult to form tight junctions that are necessary to perform the barrier properties between endothelial cells (for more information, please see Chapter 16: Blood
brain barrier tissue engineering, Agathe Figarol and Michiya Matsusaki). Endothelial cells seeded on the inner side of the fibers were shown to increase the tight junctions with shear stress formed by the flow of the medium [206]. In BBB studies, different types of cells, such as glial cells, neurons, or pericytes, can be implanted together or separately in the extra-capillary space to mimic the brain part of the BBB [203,207,208], so that the BBB system can be more effectively mimicked by cellular interactions. Hollow-fiber bioreactors are also suitable for the cultivation of cells with high metabolic activity, such as hepatocytes, because of the high rate of mass transfer achieved. In a study in which primary human hepatocytes were used, an in vitro model was obtained to perform pharmacological studies on human liver functions. They used hollow-fiber bioreactor technology to minimize the required media components and the number of cells to be used. In this system the stable maintenance of primary human hepatocytes was achieved, and it was stated that the obtained model could be used for long-term drug screening [209]. In another study, hollowfiber bioreactors were used in order to produce implantable size bone TE products that could not be produced by conventional methods due to lack of oxygen and nutrients. Sheep MSCs were seeded in extra-capillary space and cultured for 12 days to differentiate into osteoblasts. Hollow fibers mimic the Haversian canals in bone tissue, and the cells are protected from the mechanical forces produced by continuous flow [210]. Hollow-fiber bioreactors are not suitable for imaging. Furthermore, cells cannot be easily harvested due to the fact that they can be stuck in the fibers. This situation limits their use in studies requiring cell harvesting [194]. Other disadvantages include the need for expertise in their use and the high cost of commercially available bioreactors.

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27.2.8 Microfluidic bioreactors Microfluidic systems are relatively new types of bioreactors. They were produced in order to overcome many problems encountered in conventional systems, such as consumption of high amount of growth medium and media components, such as growth factors, noncompliance with high-throughput screening, difficulty to control parameters and microenvironment, high production cost, complications in live-cell analysis and imaging, inability to provide adequate oxygen, and nutrient to tissues. Microscale patterns can be produced precisely in desired geometry and dimensions in microfluidic bioreactors by soft lithography, photolithography, molding, etching, conventional machining, bioprinting, modular assembly, electroforming, electro-discharge machining, or laser ablation techniques [211
213]. They can particularly be used in studies requiring high number of experiments, such as optimization studies, statistical design experiments, strain selection, and drug screening, as they can be easily produced with relatively low-cost, and they consume lower amount of media constituents. Oxygen and nutrient transfer to tissues is increased in microfluidic bioreactors because of the large surface-area-to-volume ratio design in manufactured systems. They are also suitable for shear stress studies at single cell level, because flow control can be achieved accurately [214]. In pioneering microfluidic bioreactors, microscale patterns were designed on a fixed surface [215] (Fig. 27.11). However, only 2D cell culture studies can be performed in these bioreactors. In recent years, higher volume microfluidic bioreactors are produced to provide in vivo
like microenvironments to allow cells to form a 3D matrix. In a study conducted to obtain a 3D microfluidic system, PGS layers were produced using photolithography and plasma-etching techniques, and the layers were stacked by bonding layers physically. Thus a 3D polymeric structure with microfluidic systems was obtained [216]. A single microfluidic bioreactor can be used to simulate different tissues/organs by studying with different cell types simultaneously, and these systems are named “organ-on-a-chip” (OOC) [217]. In OOC systems, microarchitecture and functions of organs, such as lung, heart, intestine, liver, kidney, cartilage, skin, pancreas, bone marrow, and BBB, are attempted to be modeled on a single chip. Thus, for example, the effects of pharmaceuticals on different organs can be determined at once, or the effects of one organ’s reaction on another can be measured. In an OOC study modeled on four human tissues, lung (A549), liver (C3A), kidney (HK-2),

Figure 27.11 Microfluidic patterns on a fixed surface.

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and adipose (HPA), the effects of TGF-β1 released by gelatin microspheres on the lungs were observed, and it was stated that the function of A549 cells was enhanced, while the functions of C3A and HK-2 cells were uncompromised [218]. In another study, micro-liver and micro-intestine tissues were modeled on a microfluidic system, and also micro-breast cancer model was formed, where anticancer agents and estrogen-like substances were assayed [219]. There are different types of microfluidic bioreactors according to their substrates. Since glass-based bioreactors are optically efficient, they are often used in studies requiring live-cell imaging [220,221]. They can be used for long periods and many times. In addition, they are suitable for hypoxia studies, because glass is not permeable to oxygen [222]. Polymer-based bioreactors are the most preferred microfluidic bioreactors, because they are inexpensive and easily produced. They are generally manufactured from synthetic polymers, such as polydimethylsiloxane (PDMS), polymethylmethacrylate, polycarbonate, and polystyrene [223
225], but can also be produced from natural polymers, such as agarose, fibrin, and collagen for 3D studies [226,227]. Biodegradable or nonbiodegradable polymers can be selected according to the study. PDMS is the most popular polymer used for microfluidic bioreactors, because it is cheap, gas permeable, and can be manufactured easily by techniques such as soft lithography [217,228]. Microfluidic bioreactors are particularly suitable for microvascular TE studies, since precisely controlled mechanical stimuli (especially shear stress) can be applied [229,230]. In a study, 3D collagen microfluidic system was constructed by pouring the tumor cell suspension in the neutralized collagen solution in a vascularshaped mold and after gelation of collagen, the needle in the middle of the collagen was removed. Endothelial cells were injected into the micro-channel and the 3D structure was rotated in order to provide cell attachment [231]. The effect of shear stress on tumor vascularization was investigated in this system, and it was shown that increasing shear stress decreased endothelial permeability and tumor expressed angiogenic factors [232]. In another study performed with microfluidic systems a silk-fibroin microfluidic device was fabricated with a solvent casting method in PDMS molds. Microfluidic device was designed to produce maximum WSS, and perfusion culture was performed after HepG2 cells were seeded. Albumin secretion was increased from HepG2 cells after the increasing proliferation in microchannel according to static culture [233]. Electrical and optical stimuli can also be applied in conjunction with these bioreactors. An electrical stimulus has been given with microfluidic bioreactors that contain microarray electrodes and neural signals, and action potentials and electrophysiological changes were measured in neural TE studies [234,235]. The effect of light and laser-based therapy was also studied in optic stimulating bioreactors [236]. The biggest limitation of this type of bioreactors is that organ-size products cannot be obtained. Due to their small size, they are generally improper for the integration of scaffolds, which support 3D matrix formation. Another disadvantage of the system is that since the scale is reduced, some protocols have to be adjusted for each study.

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27.2.9 Combined systems The physiological loading conditions in the body are much more complex compared to simple loading conditions induced by the various types of bioreactors introduced in this chapter. To simulate native tissue microenvironment more precisely, combinations of the different types of bioreactors can be used in order to fulfill the loading requirements for tissue specific applications. The most common use of combined bioreactors involves perfusion loop on top of stretch, compression, or HP bioreactors [29]. In a study conducted for the simulation of blood vessels by Dermenoudis and Missirlis, a combined bioreactor system that is capable of delivering any combination of the four mechanical stimuli, (1) normal and (2) circumferential stresses due to blood pressure, (3) shear stress due to blood flow, and (4) the gravitational field induced by rotation, that can be controlled independently was designed and developed. Bovine capillary endothelial cells were cultured on ethylene vinyl acetate tubes, and their response in terms of viability and morphology to combination of different mechanical stimuli showed that shear stress alone at a rate of 103 s21 reoriented the cells parallel to the shear direction, while cyclic uniaxial stretch of 6.7% at 1 Hz caused a cell orientation vertical to the stress direction. It was shown that rotation was the most complex stimulus, continuously changing the polarity axis of the cells when applied alone. When rotation was combined with other stimuli, it suppressed the elongation, while not affecting the orientation profile [237]. Engelmayr et al. investigated cardiac muscle tissue formation of BMSCs on TEHVs through a bioreactor system that can apply cyclic flexure and laminar flow, simultaneously or separately. The constructs were cultured statically for the first 4 days and then were grouped into static, cyclic flexure (0.554 mm21 change in curvature at 1 Hz), flow (1.1505 dyn/cm2 of WSS), and combined flex-flow groups. The results indicated 75% higher collagen wet weight content and an effective stiffness value after 3 weeks on flex-flow group as opposed to other groups. These collagen and stiffness values were reported to be similar to the vascular SMC seeded constructs that were cultured under flexure stimulus only in the team’s previous study [128], that was also showing the effects of the flex-flow conditions for differentiation of BMSCs toward muscle cells [103]. In another study regarding cardiac TE a system that can apply mechanical strain and fluidic shear stress to the cellseeded scaffolds was designed and validated as an in vitro myocardium model. Isolated cardiac fibroblasts seeded onto 3D collagen constructs were subjected to 5% cyclic strain at 1 Hz with 10 μL/min of flow rate. This system’s usability regarding cell viability were validated (85%
90 % viability); and also, no significant change in terms of fluid permeability was detected between the static or strained gels, indicating no leakage occurred due to 5% stretch. [238]. A similar study related to cardiac TE investigated the effects of a system that can apply compression along with the fluid shear stress. After 4 days culture of neonatal rat cardiac cells, better preservation of cardiomyocyte markers in the compression and fluid shear stress group was reported as opposed to noncompressed group [165]. Lu et al. designed and developed a novel bioreactor system that applies cyclic stretch along with the rhythmic electrical stimulation and constant perfusion to simulate

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cardiac niche. Bioreactor’s biocompatibility was assessed through Ki-67 proliferation marker using HUVECs seeded on PDMS substrates cultured under 10% cyclic stretch at 1 Hz. No significant change in Ki-67 marker was detected between the static and stretched group. Stretched group was detected to have significantly higher proliferation rate or mitochondrial activity as compared to the static group validating the biocompatibility of the bioreactor system. The stretched cells were aligned perpendicularly to the direction of the stretch, validating the efficiency of mechanical stretch. The electrical stimulation system was characterized and validated through contractile response of rat cardiomyocytes [239]. Increased levels of connexin-43, a gap junction protein, on neonatal rat cardiac cells subjected to continuous electric stimulation and medium perfusion through a combined bioreactor system was achieved in another study [240]. Another bioreactor system that can apply both perfusion and electrical stimulation for tissue-engineered cardiac constructs was used on rat heart ventricle cells seeded onto porous PGS scaffolds. The constructs were subjected to four different conditions, such as no external stimulus, only electrical stimulus (3 V/cm at 3 Hz), only perfusion (18 mL/min) and perfusion with electrical stimulus. DNA content in the perfusion with electrical stimulus was found to be higher than any of the other groups. Also, the constructs cultured under both stimulations exhibited an increased amplitude of contraction, which is an indicator of functionality of the constructs [241]. In a study that tested the effects of perfusion and cyclic compression on scaffold-free cartilage constructs created by porcine chondrocytes, the constructs were cultured under static and compression (0.5, 10, and 20 N for first, second, and third weeks, respectively) conditions. GAG content was found to be significantly higher in mechanically loaded group as opposed to the static group and the native tissue. However, total collagen content was similar in all experimental groups [130]. Another study involving chondrocytes utilized a combined bioreactor system that can simulate rolling action of articular joints through cyclic shear and compressive forces. Chondrocytes seeded on PGA disks were subjected to compressive and shear loading of 6.5% at 1 Hz. The wet and dry weight of the constructs was increased 2.6-fold in combined stimulated group as opposed to static group. Also, the GAG, total collagen and type II collagen concentration were increased 2.9-fold, 1.7-fold, and 5.6-fold, respectively, in combined group compared to the static group [242]. In another study, it was reported that 10% cyclic strain compression at 0.5 Hz along with 10 mL/min perfusion induced the proliferation and fibrocartilaginous differentiation of human BMSCs on polyurethane-based constructs, which may be used for menisci TE applications [129]. Effects of HP along with perfusion (HPP) were investigated on human chondrocytes. The chondrocytes were cultured using the combined bioreactor system that applied 0.1 MPa HPs with 2 mL/min perfusion (HPP) to one group and only 2 mL/min perfusion (P) to the other. Results indicated increased mRNA expression level of collagen type II α1 chain, collagen type I α1 chain, cartilage specific genes (COMP and SOX9) in both HPP and P groups as opposed to the static control group. HPP group compared to P group showed higher expression level of an adhesion-modulatory molecule, Tenascin-C, due to the enhanced adhesion promoted by HP [243].

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A very specialized combined bioreactor for vocal fold TE application that can mimic the airflow-induced stimulation was designed and validated [244]. The bioreactor utilized perfusion loop and two synthetic vocal fold to achieve native phonation environment. Although the cell viability rates were around 90% for both static and phonated groups, the collagen type-I synthesis was found to be higher in the perfused-phonated group compared to the only perfused group. Jagodzinski et al. studied the effects of perfusion and cyclic compression on proliferation and differentiation of bone marrow stromal cells through a combined bioreactor. The cultures were maintained in three different conditions, such as only continuous perfusion with 10 mL/min flow rate, continuous perfusion with 10% cyclic compression at 0.5 Hz and static control. Proliferation of the cells and RUNX2 mRNA expressions were increased after 2 weeks both in only perfusion and perfusion with compression groups compared to the static control group. Also, the cells stimulated by both perfusion and compression reported to have the highest amount of osteocalcin after just 1 week [245]. Another experiment was conducted to investigate the effects dynamic conditions on human adipose-derived stem cells seeded on bone biomimetic nanocomposite, poly-lactic-co-glycolic acid and amorphous calcium phosphate nanoparticles, scaffolds. The scaffolds were cultured under static, perfusion (0.5 mL/min flow rate) and perfusion with compression (0.2 mm displacement at 1 Hz) conditions. Perfusion and perfusion with compression conditions upregulated the gene and protein expression of adipogenesis marker, PPAR-γ-2, while decreased the expression of osteogenic markers, ALP and RUNX2, as opposed to static conditions. However, perfusion with compression group showed significantly higher levels of osteogenic marker genes with lesser level of adipogenesis-related genes compared to perfusion only group, which proves the importance of adjustment of amplitude and frequency of the mechanical stimulations [127]. There is an increasing number of research groups that employ home built combined bioreactors for their TE experimental set-ups. Although combined systems have the advantage of tissue, size, scaffold specific stimulation of in vivo conditions to a better extend, they also introduce the complexity and less control over testing parameters. Usually, biological reactions to combined loading are not easy to predict, and definitely not a sum of the individual effects. There are numerous interplays between different cellular components, and as the number of stimulation variables increase, the optimization of the correct timing, amount, and frequency of the parameters becomes harder.

27.3

Cell-seeding techniques for bioreactors

Adherent cells need to attach to a support material and produce ECM in order to present a continuous tissue in vitro [246,247]. The first and the most important step for achieving a healthy construct is cell seeding for TE on 3D scaffolds. It is not possible to regulate the distribution and density of the cells without an effective

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cell-seeding procedure [248]. For this reason the choice of the cell-seeding technique is critical to control engineered tissue formation. Cell-seeding methods alter according to the type and shape of the scaffold material used, the scaffold thickness, pore diameter, amount and geometry of porosity, and the type of the tissue of interest [249]. Cell-seeding methods are divided into two types: static and dynamic methods. The most commonly used are static methods, where suspended cells can be seeded directly onto the scaffold or by injection. In the static sowing method the concentrated cell suspension is pipetted on the scaffold, and the scaffold is taken to the incubator for a few hours or days for the cells to penetrate into the 3D scaffold by gravity [250]. Static cell seeding is the most commonly used method because of its simplicity and minimal cell damage; however, it is the least effective method, especially for thick and low porosity scaffolds [251]. When short incubation period is applied before transferring to the bioreactor, the cells are removed by the culture media and the penetration rate is decreased. Although increasing of the incubation period enhances cell adherence, disparate ECM deposition is observed in this method, because the cells are not homogenously distributed throughout the scaffold [252]. A semi-dynamic method by slow rotation of the scaffold for tubular constructs during the incubation period for cell penetration by gravity was shown to improve homogeneity of cell distribution [237]. Scaffolds can also be coated by ECM proteins, such as fibrin, fibronectin, collagen, and laminin, to enhance cell adhesion. The coating process can be carried out by immersing the scaffold in the coating solution or by passing the coating protein through the scaffold within a perfusion bioreactor [253]. In dynamic cell-seeding methods an external force is applied to the cells to penetrate throughout the scaffold pores. Since dynamic methods provide homogenous distribution and high efficiency, more even ECM production and high growth rate are obtained. One of the most widely applied dynamic cell-seeding methods is carried out in spinner flasks. The scaffold is fixed into the spinner flask, which is partially filled with the cell suspension, and the cells penetrate into the scaffold at a certain mixing rate. However, this method requires long periods, such as 24 hours, for cell attachment and especially inefficient for low-cell concentrations [254]. Another common dynamic method is cell seeding by centrifugal force. 3D scaffolds and cell suspension are placed in a tube and centrifuged at specific rotational speed and time. Thus with the help of centrifugal force, the cells are forced to penetrate into the scaffold. The rotational speed and time applied in this method should be optimized according to the cell type and scaffold characteristics. Penetration rate and seeding efficiency increase at high speeds; however, the cells can be damaged by high centrifugal force. Centrifugal cell seeding is usually effective in suspensions with low cell density [250,254]. In the cell-seeding method by vacuum the scaffold and the cell suspension are placed in a closed system and vacuum is applied, leading to cell penetration into the scaffold. This method is a fast seeding method and the seeding efficiency is high [255]. One of the most efficient dynamic methods is cell seeding by perfusion bioreactors. In this method the scaffold is placed in the perfusion bioreactor, and the cell suspension is continuously passed through the scaffold. By perfusion method the cells are shown to present more uniform distribution and high

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viability compared to static and spinner flask methods [252]. The disadvantages of this method are that the seeding period can be long, and a separate bioreactor system for cell seeding, as well as expertise are required. In addition, flow rate should be optimized according to the needs of the system for efficient cell seeding. The cells can be swept away from the scaffold at high speeds, while at low speeds a homogenous distribution may not be achieved. In addition to the commonly used spinner flask, centrifugation, vacuum, and perfusion methods, compression release
induced suction (CRIS) and magnetic cell sowing methods are also used. In the CRIS method the scaffold and the cell suspension are placed in a closed system, where compression is applied to the scaffold. Repeated utilization of compressive loading and unloading induce the cells to enter the inner parts of the scaffold [248]. In the magnetic cell-seeding method the cells labeled with magnetic particles are forced to penetrate inside the scaffold by the help of externally applied magnetic force [256]. Especially for vascular TE, novel cell sheet
based and electrostatic cell-seeding methods can be used. In cell sheet
based cell-seeding method, first, the cells are allowed to form a confluent monolayer in a temperature-sensitive culture plate; then they are removed as an intact sheet of cells and wrapped around a vascular scaffold [167]. Although it is a disadvantage that cultivation may need to be performed for a long time, a very homogenous distribution can be achieved by this method. In electrostatic seeding method, it is aimed to increase the adhesion of the cells to the scaffold by changing the surface properties of the scaffold [257]. The surface charge of the scaffold is changed by electrostatic cell-seeding apparatus from negative to positive in order to enhance cell attachment. Because of the negative charge of the cell membrane, the surface of the scaffold is polarized and positively charged, and thus the cells can easily be attached to the surface. Apart from these methods the cells may also be entrapped while the scaffold is produced. Cells are mixed in the polymer solution during the scaffold production and are distributed homogeneously within the inner parts of the scaffold [258]. But this method requires that the scaffold production technique should not contain any cytotoxic components, such as toxic solvents, heating or freezing steps, or high shear stress. New bioprinting techniques are candidates for homogenous cell seeding, as well as high control over construct organization. The summary of the advantages and disadvantages of the bioreactor types used in the field of TE and the types of stimulation applied to TECs according to the bioreactor types are given in Table 27.1.

27.4

Design considerations and future outlook

Bioreactors are becoming an indispensable part of TE research, since they have the potential of improving the efficiency of the process, particularly for clinical implementation of TECs. TE bioreactors can provide G

G

improved mass transport, highly controlled culture conditions,

Table 27.1 The properties of different bioreactor types. Type of bioreactor

Type of stimulation

Advantages

Disadvantages

References

Stirred

Shear stress

Increased cell proliferation rate, matrix deposition, and expression of phenotype specific proteins induced by shear on the scaffold

[28,42
44, 53
59]

Wave

Shear stress

Low shear stress Bubble-free system is suitable for sensitive cells

Parallel-plate bioreactor

Shear stress

Easy to manufacture Inexpensive

Cellular damage and loss of ECM because of high shear stress Mass transport is limited to ,1 mm surface The need of optimization for the rocking rate, velocity, angle, and bag filling level for each specific study Limited mass transport Expensive Hard to utilize for 3D constructs

Rotating-wall vessel

Low shear stress, reduced gravity conditions

The cells are protected from excessive shear stress and turbulence Induction of self-organized spheroid formation within a few hours Microgravity is mimicked; thus studies on the development of tissues in space environment can be carried out

The optimization of the culture conditions can take a long time Cells can be damaged due to possibility of scaffold collision to the bioreactor wall Mass transport is limited to ,1 mm surface Nonuniform cell and matrix distribution

[81
91,95
97]

[66
68]

[73
75]

Strain

Compression Tension Bending Torsion Pressure

Perfusion

Shear stress

Hollow fiber

Low shear stress

Microfluidic

Shear stress Tension Compression Pressure

Ability to mimic physiological loading conditions Increased cell proliferation rate, matrix maturation, and expression of a variety of phenotype specific proteins Mimicking in vivo physiological environment of the tissue while providing uniform cell distribution, efficient mass transfer, and maturating tissue-engineered construct with high viability

Cells can be protected from mechanical forces Vascularized tissue constructs can be obtained Cocultivation of two different cells without a common medium is allowed High surface area/volume ratio Precisely controlled mechanical stimuli Decreasing the consumption of a high amount of growth medium and media components High-throughput screening Low production cost

Possibility of damage in constructs due to scaffold mounting and applied direct strains Limited mass transport

[44,57,98
101, 104
115, 117
130, 135
140]

Suitable only for mechanically durable and high porous scaffolds ($70%). Optimization of flow rates is vital High rates of flow induced shear can cause cell and membrane disruption Not suitable for cell imaging Cell harvesting is limited Commercial hollow-fiber bioreactors are expensive

[141,143, 153
167, 171
180]

Organ-size products cannot be obtained Some protocols have to be adjusted for each study due to the reduced scale

[217
219, 229
236]

[189,193, 199
210]

(Continued)

Table 27.1 (Continued) Type of bioreactor

Type of stimulation

Advantages

Disadvantages

References

Combined

Shear stress Compression Tension Bending Torsion Pressure Electromagnetic

Ability to apply different kinds of stimuli simultaneously

Requires higher level of expertise. Increased number of parameters makes optimization of culture conditions harder

[103,127,128, 130,165, 237
245]

ECM, Extracellular matrix.

Bioreactors in tissue engineering: mimicking the microenvironment

G

G

G

G

G

737

physiologically relevant stimuli, continuous medium supply, reduction of process steps, automated sampling for quality control (QC), and standardization.

Improved mass transport, by far, is the key objective of using TE bioreactors, since improper cell viability due to lack of vascularization has been the ratelimiting aspect of successful implementation of TECs in clinics. While enabling growth of the tissue in larger dimensions than statically diffusible 100
200 μm layers, bioreactors show adequate oxygen, nutrient, and biosignal supply to the interior of TECs with the combination of porous scaffolds and accurate perfusion [38,150,259,260]. Also, bioreactor-based cell seeding is more advantageous, since the cell suspension penetrates through porous scaffold, and dynamic culture drives cells to migrate and settle homogenously compared to conventional cell-seeding techniques [141,260]. Although perfused fluid flow has shown great improvement for mass transport problem, and allowed the production of clinically relevant-sized constructs, there is a drawback of increased shear stress with increased flow rate. Particularly, since the desired cell proliferation and matrix deposition change the porosity profile of the construct, particularly in long-term cultures, increased shear rates that are damaging to the tissue are observed due to increased tortuosity. Adaptation of hollow-fiber approach that uses tube or vessel-like structures within the construct to overcome the transport problem while creating low shear stress on the cells could be a target for future applications [261]. Having overcome the mass transport burden, contemporary bioreactors must be designed to further improve the culture conditions to mimic physiological tissue microenvironment, including biochemical, biophysical, as well as mechanical and electromechanical factors. In most of the cases, functionalization and maturation of the TECs is only possible with accurate mechanical stimuli (e.g., shear stress, compression, stretch, and hydrodynamic pressure) [38,150,260,262,263]. Since many physiological movements involve cyclic loading, implementation of cyclic strain in the loading scheme seems to be beneficial choice. Cyclic strain has already been shown to increase mechanical properties such as modulus or elasticity, particularly for cartilage and cardiovascular constructs [101
104,238,239]. Combined bioreactor systems are far more successful in simulating physiological tissue microenvironments, but increased number of parameters leads to higher demand for expertise and higher effort for optimization of culture conditions. Even failure of optimization of flow rates alone can lead to differentiation toward an undesired phenotype, or even cell damage [39,143,264,265]. When it comes to apply the theoretical notions onto culture parameters, most of the studies rely upon time-consuming, experimental, trial and error methods specific for the application itself [143,260,266]. Accurate approach for each parameter for different target tissues, scaffolds, cell type(s) is crucial. For example, HP applied during cartilage culture can lead to an improved mass transfer of small and large molecules into the cartilage matrix but can also induce a mechanical stimulation of embedded cells

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Biomaterials for Organ and Tissue Regeneration

[57,138
140,243]. These two effects have to be examined separately. Particularly, when a combination of forces (e.g., HP, shear stress induced by perfusion, shear stress or gliding forces induced by mechanical impact, and pulsation flow) is applied, the exact local conditions experienced by the cells have to be determined. To avoid time and money consumption upon optimizing the parameters based only on experimental inferences, manufacturing procedures should be supported by in silico data or theoretical calculations. Simulations based on CFD or finiteelement approach can give a solid opinion that can potentially be used to design and predict properties for TE bioreactors. Cell distribution, nutrient transport, mineralization, and degradation scenarios have been in silico modeled by various groups [38,143,145,146,266] (for more information, please see Chapter 28: Simulation of organ-on-a-chip systems, Filipovic et al.). Supporting in vitro experiments of those simulations and mathematical models, in most cases, have shown good correlations with data from bioreactor dynamics and show promising predictive methodology. With machine learning algorithms and adding in silico evaluation into routine procedures for bioreactor applications, required knowledge in this field will expand exponentially together with improvement in understanding the mechanism behind each phenomena. High control over culture conditions brings along repeatability, as well as standardization. Implementation of online monitoring systems, such as continuous monitoring of O2 tension, pH, essential culture constituents such as glucose, cellproduced products [267,268], and possible realization of closed loop feedback control systems that can regulate the culture conditions accordingly would highly improve tissue viability and maturation, as well as repeatable quality of the final constructs. Feedback systems can even be implemented to adjust mechanical loading parameters according to ECM deposition rate, or expression of specific cytokines. The new era, with the leap of biotechnology and nanotechnology, has changed our way of understanding of medicine and expectations from healthcare approaches. “Personalized” or “precision” medicine is becoming increasingly important, and all future medical interventions, including TE, must be tailored to the individual patient. Bioreactors designed according to the specific patient’s needs would be necessary for personalization of TECs. Patient-derived cells, cultured on specifically (bio)printed scaffolds in the size and shape of the defect of interest, and tailored by patient specific culture media with particular loading conditions is no fiction. Recent discoveries on influence of our microbiota over tissue and organ functioning, and increased awareness on the effects of immune system cells and immunoregulatory factors on modulation of tissue functions [269] calls for introducing these elements among currently involved parameters in our future bioreactor designs. For the ultimate goal of replacing damaged tissue with a TEC, controlling the environment is not enough itself. For a clinically relevant approach, QC and good manufacturing practice requirements have to be met for design and production steps. To fulfill these necessities, whole production and cultivation scheme,

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739

including biopsy, proliferation, cell seeding, tissue formation, and delivery to the site of application (e.g., hospital), must be designed under this concept, which includes automated steps for statistically reliable, reproducible, scalable, uniform, monitorable, and safe manufacturing [38,149,264].

27.5

Conclusion

Production of functional tissues in vitro requires an orchestrated combination of various biochemical, biophysical, and mechanical cues. Bioreactors provide an efficient means to introduce these factors in a highly controlled manner, but the parameters must be optimized for each patient, cell, scaffold, tissue-specific case. The complex interactions between various parameters within a bioreactor environment are yet to be discovered for a highly effective bioreactor design, and clinically relevant production of TECs, and this discovery can only be led by a collaborative interdisciplinary effort.

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3):1412. [258] Hoffman AS. Hydrogels for biomedical applications. Adv Drug Deliv Rev 2012;64:18
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9.

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[267] Lee PS, et al. Developing a customized perfusion bioreactor prototype with controlled positional variability in oxygen partial pressure for bone and cartilage tissue engineering. Tissue Eng, C: Methods 2017;23(5):286
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27.

Further reading Lanza R, Langer RS, Vacanti J. Principles of tissue engineering. 4th ed. Academic Press; 2013.

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Nenad Filipovic1,2,3, Milica Nikolic1,2,3 and Tijana Sustersic1,2,3 1 Faculty of Engineering, University of Kragujevac (FINK), Kragujevac, Serbia, 2Steinbeis Advanced Risk Technologies Institute doo Kragujevac (SARTIK), Kragujevac, Serbia, 3 Bioengineering Research and Development Center (BioIRC), Kragujevac, Serbia

28.1

Introduction

As science and technology evolves, knowledge of entire organisms fit into the in vitro models, representing a starting point in biological and medical research. The combination of traditional materials biology and chemistry offers an extraordinary opportunity for the development of new materials, devices, and processes [1]. One of the most important advances in materials research has been the introduction of modeling and simulation as an integrated part of the research and development in materials science. Numerous methods are developed to describe phenomena in materials science on several levels with respect to space and time. The importance of this kind of modeling and simulation in materials research is visible especially in the number of papers and research that are increasing. In vitro tissue engineering has become now a major field to investigate in order to provide solutions for organ transplantation, since there is a shortage of available tissue. Experiments alone cannot provide a full insight into biophysical and biochemical processes that affect tissue growth and behavior of the tissue cells [2]. Therefore mathematical modeling can provide additional information in different underlying processes that influence final tissue formation for implantation; it can be used to predict obtained material properties based on processes used during their creation (i.e., electrospinning) [3], but can also estimate how drugs behave in different devices, device efficacy [4], etc. Today, investigations in biology and medicine start with in vitro testing. Observation in behavior of the cells, isolated from the organism, is essential for the improvement of the science and increase of knowledge. Once the results from in vitro experiments are collected, they need to be validated with in vivo measurement, which is a more invasive technique. In the process of validation, researchers concluded that observed cells behave differently when they are isolated from the system and that they lack some functions and processes occurring in living organism. Therefore the need for the cells environmental representation improvement is raised, that is, microfluidic chips are made. Microfluidic chips are small in size and contain multiple microfluidic channels. Microfluidic channels are lined up with living human cells of developing organ. Biomaterials for Organ and Tissue Regeneration. DOI: https://doi.org/10.1016/B978-0-08-102906-0.00028-3 © 2020 Elsevier Ltd. All rights reserved.

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Besides these organ building cells, there are endothelial vascular cells, forming artificial vasculature. Interaction between organ building cells and endothelial vascular cells gives better environmental conditions for the observed cells in comparison to the in vitro testing. Even though in vitro testing and newly developed organ-on-a-chip devices give plenty of valuable information about monitored cells, there is a need for the development of computer in silico models, primarily to reduce experiments on animals and clinical trials. Improved technology also reduces time responses of the systems. Therefore coupled with experimental in vitro models, computer in silico models are further developed. These models can be built on different scales (macro, meso, micro, nano), which depends on the purpose of the model—macromodels use, for example, finite element method (FEM) and explain behavior of the whole system on the level of organs; mesomodels use mesoscopic methods such as dissipative particle dynamics (DPD) and explain behavior at the level of molecular clusters; micro- and nanomodels use molecular dynamics method and explain behavior at the molecular or atomistic level. Under the Horizon 2020 project PANBioRA, as part of the integrated platform for biomaterial risk assessment, one work package is dedicated to the in silico modeling. Development of workflow multiscale platform for basic in silico testing is in progress: personalized immunoprofiling-based biomaterial monitoring, organon-a-chip monitoring of the effects of the biomaterial at single organ, biomaterial
tissue interaction levels, and also parts that will be simulation or experimentally supported simulation only (biomaterial corrosion, massive release and effects on microbiota, respectively) [5]. The ultimate goal is to verify these models using experimental data and develop adequate in silico models with accurate and satisfactory results, which would reduce costs and time of repetitive experimental measurements. In order to achieve this, in vitro models are the first step. In vitro models were developed for many human organism segments—osteoarthritis [6], psoriasis [7], blood
brain barrier [8], myocardial tissue [9], and myocardial ischemic injury [10], as well as paroxysmal supraventricular tachycardia [11], Alzheimer’s disease [12], murine middle ear epithelium [13], vascular inflammation [14], thrombosis [15], models of cancer [16] etc. In this chapter, how such models can be used for tissue engineering and regenerative medicine will be covered.

28.1.1 General overview Nowadays, an area that has gained much research interest is the field of in vitro tissue engineering, mainly because new approaches, such as multiscale modeling, have provided an insight into mechanisms that drive different biological processes. Others reason for that is also the shortage of available tissue for organ transplantation [2], as well as there is a global need to replace or reduce the use of animals in tissue examinations [17
19]. Sometimes, there are some events or processes, that is, tissue growth and behavior of the tissue cells, for which experiments alone cannot provide enough information to describe them [2]. In these cases, mathematical

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modeling can provide additional descriptions and mechanisms that lay under the observable mechanisms and affect final tissue formation, used in implantation. Organ-on-a-chip systems are an attempt to miniaturize tissue level biological events (Fig. 28.1). Many researchers and institutions are currently working on development of the different organs-on-a-chip—heart [20], lungs [21], liver [22], kidney [23], brain [24], bone [25], etc. The major advantage of organ-on-a-chip technology is its capability to represent structural and functional complexity of living tissues and organs, unlike in vitro cell culture techniques, which fail in the reproduction of dynamic mechanical and biochemical microenvironment. Organon-a-chip microdevice mimics microsystems and possesses tremendous potential as an innovative and predictive screening platform [26]. Building artificial organs, which should further be validated, requires knowledge in cellular manipulation, coupled with knowledge in physiological behavior of tested organs and their response to different events. Kitsara et al. [20] wrote a stateof-the-art for the heart-on-a-chip. The main approaches of microfluidics for cardiac

Brain on-a-chip

Heart on-a-chip

Lung on-a-chip

Arteries on-a-chip

Kidney on-a-chip

Liver on-a-chip

Gut on-a-chip

Bone on-a-chip

Figure 28.1 Schematic representation of different organ-on-chips.

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culture systems and their assessment are described in the paper [20]. Huh has focused on describing microphysiological system that replicates the functional unit of the human lung—alveolar
capillary interface [21]. The model describes dynamic mechanical activity and physiological function of the breathing lung-on-achip and can be used for investigating complex human disease processes of the lungs. For that reason, he is considered as one of the founders of lung-on-a-chip model. Knowlton and Tasoglu [22] were focused on developing the liver-on-a-chip model. Because liver is responsible for detoxification of the metabolic products, it is subjected to the high frequency of drug-induced injuries. Therefore liver-on-achip tissue model is necessary to be developed for drug screening and hepatotoxicity testing. Lee and Kim [23] examined the state-of-the-art for kidney-on-a-chip models. Various drugs applied to the human body can result in kidney dysfunction; therefore drug development process needs to address this issue. The main advantage of using kidney-on-a-chip model in comparison to the in vivo animal experiments is the reduced discrepancies in drug pharmacokinetics and pharmacodynamics between humans and animals. Several kidney-on-a-chip models show the ability to reproduce microenvironment of the kidney tubule and drug nephrotoxicity. With these microfluidic chips, drug-induced biological responses can be measured. Kidney-on-a-chip models can be used in kidney disease modeling and development of novel therapies according to the identification of drug effects, interactions, and nephrotoxicity. Wheeler [24] researched brain-on-a-chip systems. These models are very complex to develop because the functioning of the brain is not understood completely. Brain-on-a-chip models mainly include nerve cells growing in the culture with the ability of signal conduction—stimulation of nerve cells. Hao et al. [25] examined bone-on-a-chip model for the purpose of investigating bone metastasis to breast cancer cells. About 70% of the metastatic breast cancer leads to metastasis of bone tissue. The mechanism of bone tissue “infection” by breast cancer in metastasis is not completely revealed. For the purpose of bone-on-a-chip model development, mature osteoblastic tissue is naturally formed. Osteoblastic tissue contains heavily mineralized collagen fibers. Coculture of metastatic breast cancer is examined with osteoblastic tissue. The described model shows an opportunity to investigate the interaction of cancer cells with bone matrix. Unique hallmarks of breast cancer bone colonization confirmed with in vivo experiments can be observed using this model. Special attention should be paid to the lung-on-a-chip devices, as they are somewhat in the center of research. Wyss Institute for Biologically Inspired Engineering at Harvard University was the first to develop lung-on-a-chip, as a microfluidic device, which mimics breathing human lungs on a microchip [27,28]. Lung-on-achip can be utilized in order to simulate breathing of healthy lung, the movement of nanoparticles, which origins from air and water pollution and cause inflammation, to test new therapies and drugs for lung diseases, to measure drug toxicity or simulate the progression of lung cancer. Microdevice has been used in order to obtain a better understanding of lung’s function and properties as well as to simulate pulmonary diseases [29]. The main processes during the breathing in lungs are alveolus and capillary gas exchanges. Besides the gas exchanges simulations, microdevices

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can be used in toxicity analysis and studies of pulmonary diseases. The aforementioned model developed by Wyss Institute is used to analyze nanoparticles’ absorption and acute toxic response to the nanoparticle absorption due to stretching of alveolar
capillary barrier [30,31]. As previously mentioned, several different organs were point of interest in developing on-chip technology. However, special interest is taken into liver, mainly because liver is involved in more than 300 vital functions [32]. However widely researched, both through modeling and experiments, an artificial organ or medical device that is capable of replacing the liver has not been made yet [33]. Some available systems have been produced to remove blood toxins accumulated during liver failure (e.g., hemofiltration, hemodialysis, and plasma exchange). However, the main drawback of these systems is the fact that the chance of patient’s survival is very low with the use of these systems alone [34], meaning organ transplantation is still the only available treatment in case of serious, end-stage liver diseases. The following chapter presents available models for lung and liver. In order to demonstrate how mathematical modeling can be incorporated into on-chip organ models, modeling of bioreactor under perfusion flow is presented, followed by modeling of liver cell aggregation [which is the main way of creating liver-like organoids (hepatocyte spheroids)], with adequate results and comparisons with the literature. Finally, the conclusion section provides a summary of the state-of-theart, obtained results of the modeling section, and gives possible directions for future improvements and trends in this field.

28.1.2 Review of the lung and liver cell line models Lung-on-a-chip model was first developed at Wyss Institute, and from that point, different model variations were investigated by different authors. Lung-on-a-chip models include two chambers—one with circulating air and the cells forming epithelial layer, representing airways, and the other with circulating blood, containing endothelial cells, representing blood vessels. Interaction between epithelial and endothelial cells is happening through the membrane separating these two chambers. This model can be used for the simulation of healthy lungs, as well as several types of lung diseases. The model of bioreactor is created with the purpose to be used as a part of lungon-a-chip device. The idea is that bioreactor represents blood vessel, through which blood flows together with immune cells, specifically monocytes, as they are the first to react among the immune cells. Immune reaction can be provoked by the presence of some bacteria in lung tissue or as a reaction and side effect of certain drugs applied to the diseased lung tissue—anticancer drugs or medicament treatment for certain lung disease type. In order to test lung cells behavior, different lung cells are used, both primary (e.g., human bronchial epithelial cells and HBEC) and nonprimary lung cells (e.g., A549, Calu-3, and BEAS-2B). Stewart et al. investigated asthma and bronchial epithelial cell culture systems as air
liquid interface
differentiated models [35]. They used primary HBEC cells

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and nonprimary Calu-3 and BEAS-2B cells. Focus of their investigation was on the application of the nonprimary cells in asthma research, since primary cells are expensive for large studies and drug screening. They found that Calu-3 cells express epithelial markers similar to the primary cells and develop a high transepithelial electrical resistance, which is an important parameter in the formation of tight junctions and can be used as a marker of epithelial disruption. Kreft et al. concluded that Calu-3 cell line cultured with air-interfaced culture produced a cell layer similar to the real airway epithelial morphology and electrical resistance in vivo. They investigated Calu-3 cell line under different culture conditions [36]. Calu-3 cell line is also used for the growth inhibition in research of anticancer therapeutics, because of its cancer origin [37]. Interaction between drugs could induce negative side effects and increase toxicity. Mamlouk et al. concluded that nonsteroidal antiinflammatory drugs (i.e., aspirin, ibuprofen) reduced the uptake of salbutamol across Calu-3 cells [38]. In the literature, several developed lung-on-chip models and many in vitro testing of lung cell lines can be found. Some of them are mentioned in this chapter. These models are formed experimentally and usually analyzed by imaging technology, which is in rapid progress these days. However, not many mathematical and computer in silico models, explaining lungs’ functionality and lungs’ disease processes, cannot be found in the literature. Two of them are explained next. Savla et al. investigated the repair of airway epithelium after injury, which is important in restoring epithelial barrier integrity [39]. The model they built, which is based on numerical solutions of mathematical equations, focuses on the differences in spreading, migration, and proliferation of the epithelial cells with cyclic strain applied. Cyclic strain is used to mimic respiration process, and their research showed that both elongation and compression slowing airway wound closure and that compression has a greater impact. Numerical solution is utilized to determine the shape of the diffusive wave equation of the cell density. For the estimation of the parameters, they used NHBE, 16HBE, Calu-3, and CET cell lines. Significance of this model is reflected in comparison to the experimental data and good matching was obtained between the experiments and computer model. This model may be utilized for testing of the epithelial layer and barrier formation of different lung cell types. Hancock and Elabbasi [40] created lung-on-a-chip model in COMSOL commercial software. They developed this model upon Huh research [21]. With COMSOL software, computational fluid dynamic for blood flow can be simulated, together with fluid
solid interaction, representing the interaction of the blood vessel and lung tissue cells. They modeled the polydimethylsiloxane membrane between the two chambers as nonlinear material model with a thickness of 10 μm. COMSOL offers particle tracing (which can be utilized for nanoparticles inhalation into lungs, drug transport through the porous membrane, etc.). Vacuum pressure waveform is applied to simulate the breathing process. This model simulates a well lung-on-achip device developed by Huh, but mathematical model is not explained and given in this paper. The fact that a small number of mathematical and computer models for organon-a-chip devices exist, specifically lung-on-a-chip, opens a great area for

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investigation and computer modeling section. There are several advantages of development of in silico models for lung-on-a-chip devices. In silico models, once validated, can provide a faster response of the system and can reduce number of experiments performed on animals and in human clinical trials. For example, on animals and humans, existing or newly developed drugs with the best achieved results can be applied. The model of bioreactor can be used to test immune response of the tissue, which is supplied by blood, to the different biomaterials that are used for implants, as it was explained in the paper by Sharifi et al. [41]. Liver tissue consists of around 80% of hepatocytes under normal physiological conditions and in vivo [42]. Because of the fact that most of the liver’s important functions are performed by these cells, the focus of research is on them [43]. Another proof that investigation of liver has been popular is the review that gives an overview of different microtechnologies that have been used for liver studies, where biomimetic technologies for constructing microscale 2D and 3D liver models are presented, as well as liver-on-a-chip systems and microscale bioreactors [44]. In addition, the same paper presents applications of liver microtechnology and predicts future trends in this area [44]. After literature review, it was concluded that the formation of spheroidal liver cell aggregates appears to be very important in growing functional liver tissue, which will remain viable for longer periods, in vitro [33]. The main aim in using mathematical models lies in the idea to understand the dynamics of cell aggregation, and particularly the roles of cell
ECM (extracellular matrix) adhesion, chemical cues that regulate cell movement (chemotaxis) and cell
cell interactions in coculture. The two-phase approach that we will later explain in detail allows to couple the motion of the cells with culture medium in which they are grown, where the two phases are treated as viscous and inviscid fluids, respectively. In addition to the previous multiphase models [45], here we couple cell movement with the deformations of the ECM (which coats the base of the culture well). Numerical simulations are then used to determine the effects of certain key parameters on the cell aggregation. Some authors have investigated and created models to predict previously unrecognized mechanisms, having already quantitatively characterized the architecture of liver lobules, the repetitive functional building blocks of liver [46]. It was suggested that hepatocyte aggregation into spheroids is the result of their increased activity, but due to the absence of a vascular network the cells in large spheroids, mass transfer limitations are experienced. Therefore a group of authors wanted to define the spheroid size that enables maximal cell viability and productivity [47] and developed a combined theoretical and experimental approach to define this optimal spheroid size. Researchers took interest in different techniques to examine reduced dedifferentiation of hepatocytes, causing them to lose their normal functions and eventually die [48,49]. Since the process of dying cells usually happens after few days, coculturing with other cell types (i.e., stellate cells), as well as the use of growth factors and combinations with polymer scaffolds have been introduced [48].

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Formation of multicellular spheroids has been suggested as a possible solution, as these spheroids have been mimicking liver tissue in vivo to some extent and improved viability in such way [50]. In that sense, culturing hepatocytes as spheroids in a normal procedure involves seeding cells in culture wells with ECM, whilst nutrients are provided to the cells by culture medium [51]. If the substrate is adequate, cells should aggregate within an estimated period of approximately 1 day [52,53]. Formed clusters (called aggregates) detach from the surface of the ECM and form spheroids with a diameter in the range from 100 to 150 μm [54]. In aggregates the cell diameter is in range of 10
30 μm [52,55]. When modeling these processes, and based on the biological context, it is necessary to consider interactions between different types of materials like cells, extracellular water [45,56], ECM [57], and cell populations such as macrophages [58]. In that sense, these different materials are modeled as different phases, forming multiphase models. These multiphase approaches take into account coupled motion of the cells and culture medium in which they are grown. A review of the in vitro and on-chip models in the sense of biomaterial interaction research, with primary concern on lung and liver cell lines, has been published by our group recently [59]. In most studies, values of the relevant parameters are based on literature data or are estimated in very high ranges. In addition, literature research has shown that studies that investigate multiphase models are rare. Several studies such as by Glicklis et al. [47], Breward et al. [45], and Green et al. [51] investigate somewhat multiphase models, otherwise there is very little literature on multiphase models regarding liver. The last-mentioned study has been the only available research found to be dealing with multiphase model of a liver from the mathematical side and modeled aggregation process. The main drawback of this research is that chemical signals are neglected by prescribing the forces generated by the cells as a function of cell density and an idealized one-dimensional geometry are assumed.

28.2

Review of numerical solutions of developed models

This chapter uses several numerical methods in the development of models. For the purposes of using them later in the chapter, we will first explain them in more detail.

28.2.1 Finite-difference method In mathematics, finite-difference methods (FDMs) are numerical methods for solving differential equations by approximating them with difference equations, in which derivatives are approximated by finite differences. The principle of FDMs is close to the numerical schemes used to solve ordinary differential equations. It consists in approximating the differential operator by replacing the derivatives in the equation using differential quotients. The domain is partitioned in space and in time, and approximations of the solution are computed at the

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space or time points. The error between the numerical solution and the exact solution is determined by the error that is committed by going from a differential operator to a difference operator. This error is called the discretization error or truncation error. The term “truncation error” reflects the fact that a finite part of a Taylor series is used in the approximation [60]. For the sake of simplicity, we shall consider the one-dimensional case only. The main concept behind any finite-difference scheme is related to the definition of the derivative of a smooth function u at a point x: uð x 1 hÞ 2 uð x Þ h!0 h

u0 ðxÞ 5 lim

(28.1)

and to the fact that when h tends to 0 (without vanishing), the quotient on the righthand side provides a “good” approximation of the derivative. In other words, h should be sufficiently small to get a good approximation. It remains to indicate what exactly is a good approximation. Actually, the approximation is good when the error committed in this approximation (i.e., when replacing the derivative by the differential quotient) tends toward zero when h tends to zero. If the function u is sufficiently smooth in the neighborhood of x, it is possible to quantify this error using a Taylor expansion [60]. PDEs (Partial differential equations) are in such way converted into a system of linear or nonlinear equations, which can then be easily solved by matrix algebra techniques, and therefore this adaptation is suitable when solving the problem on computers. Derived from FDM, Crank
Nicolson method is usually used for numerically solving the heat equation and similar PDEs. Since it is numerically stable, it is generally more suitable for solving diffusion equations (or equivalent). Problems involving time, as one independent variable, sometimes lead to parabolic PDEs, the simplest of which is the diffusion equation (also used in this chapter), derived from the theory of heat conduction [61]. The diffusion equation plays an important role in a broad range of practical applications such as fluid mechanics. Only a limited number of special types of parabolic equation have been solved analytically and the usefulness of these solutions is further restricted to problems involving shapes for which the boundary conditions can be satisfied. In such cases, numerical methods are some of the very few means of solution. The main limitation of the analytical methods is that they are only applicable to special, noncomplex geometrical shapes, with material parameters independent on the solution variables and other simplified conditions. They are not applicable to general, complex nonlinear problems, for example, airflow in alveolated structure with large motions of the boundaries, or blood flow in microvessels. When modeling field problems by the FEM, we seek the solution that satisfies the governing equations within the finite elements (subdomains) in a weighted sense (details are given in the next section). Also, the physical field within each element is approximated and expressed in terms of the nodal values, as in the case of solids where the displacements of points within a finite element are approximated from the nodal point displacements.

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28.2.2 Finite element modeling The FEM is a numerical method that is used for solving problems in different areas of engineering. Usually, a typical problem that is to be solved with FEM includes structural analysis, heat transfer, fluid flow, mass transport [62], and electromagnetic potential [63]. The analytical solution of these problems generally requires the solution to boundary value problems for PDEs. The FEM formulation of the problem results in a system of algebraic equations. The method yields approximate values of the unknowns at discrete number of points over the domain [64]. To solve the problem, it subdivides a large problem into smaller, simpler parts that are called finite elements (Fig. 28.2). The simple equations that model these finite elements are then assembled into a larger system of equations that models the entire problem. FEM then uses various methods from the calculus of variations to approximate a solution by minimizing an associated error function. Studying or analyzing a phenomenon with FEM is often referred to as finite element analysis. One of the first steps in FEM is to identify the PDE associated with the physical phenomenon. The PDE (or differential form) is known as the strong form and the integral form is known as the weak form. Once the integral or weak form has been set up, the next step is the discretization of the weak form. The integral form needs to be solved numerically, and hence, the integration is converted to a summation that can be calculated numerically. In addition, one of the primary goals of discretization is also to convert the integral form to a set of matrix equations that can be solved using well-known theories of matrix algebra. As shown in Fig. 28.2, the domain is divided into small pieces known as “elements” and the corner point of each element is known as a “node.” The unknown functional uðxÞ are calculated at the nodal points. Interpolation functions are defined for each element to interpolate, for values inside the element, using nodal values. These interpolation functions are also often referred to as shape or ansatz functions. Once the matrix equations have been established, the equations are passed on to a solver to solve the system of equations. Depending on the type of problem, direct or iterative solvers are generally used. More detailed explanation on FEM could be found in Refs. [65,66]. As previously mentioned, FEM method has been used in many applications, out of which applications in cardiovascular system are one of the widely researched problems in Bioengineering Research and Development Center (BioIRC)

Figure 28.2 Schematic representation of FEM concept. FEM, Finite element method.

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Figure 28.3 FEM method applied to the blood flow in carotid artery [left—discretization of the domain with FE, right—distribution of von Mises stress (Pa) within the artery walls of the carotid artery bifurcation due to action of blood, for systolic deceleration flow]. FEM, Finite element method.

(Fig. 28.3). FE mesh is shown on the left—the blood flow domain is modeled by 3D fluid finite elements. On the right is the example of results—distribution of von Mises stress (Pa) within the artery walls of the carotid artery bifurcation due to action of blood, for systolic deceleration flow. Other applications include muscoskeletal system, testing of cardiovascular stents, electrospinning simulations, etc.

28.3

Modeling of bioreactor for lung cells

Flow conditions and immunocompetency are important components of organ-on-achip systems. Thus there is a need for understanding and modeling of cellular movement and attachment in mini bioreactor systems for the development of such organ-on-a-chip systems. As an example, the analysis of fluid flow within the bioreactor where the fluid contains cells (such as monocytes) was given below. In this configuration, fluid flows through the bioreactor constantly, with the same input velocity, thanks to peristaltic pump, which pumps flow from the reservoir (the system is presented in the Fig. 28.4) and the cells in the fluid sediment onto the bottom surface of the bioreactor. Geometry of the bioreactor is taken from the literature. Sharifi et al. [41] analyzed bioreactor model while modeling the cascade of events during immune cell response to the implants. Geometrical parameters are chosen to be similar to the geometrical parameters of the Shafiri et al. bioreactor model [41]. Values of the geometrical parameters are given in Table 28.1 and the model is presented in Fig. 28.5. Behavior of the fluid inside the bioreactor was modeled using Navier
Stokes equation together with the continuity equation. Monocytes as particles travel being carried by the fluid flow. Their movement in the model is presented as mass transfer in fluid. Detailed mathematical model is described in the following text.

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Figure 28.4 Bioreactor fluid supply system.

Table 28.1 List of the geometrical parameters of the bioreactor with used values and units. Geometrical parameter

Value

Unit

a b c d r

4 2.5 15 8 2.5

mm mm mm mm mm

Prescribed boundary conditions included fluid inlet velocity and zero pressure at the outlet. Monocytes were introduced with the fluid at the inlet. Distribution of the monocytes along with the bioreactor is monitored, and trajectories of the monocytes are obtained based on calculated drag force, defined by Stoke’s law, which assumes movement of the spherical insoluble objects in the fluid. Model was solved in open-source solver PAK, developed at the Mechanical Faculty of University of Kragujevac and BioIRC. PAK solver uses FEM to solve a specific engineering problem. It has different modules for computation of fluid dynamics, analysis of solid bodies, and fluid
solid interactions. PAK can be used to obtain fluid and solid velocity vector and pressure, solid deformation and stress analysis, temperature field, electrostatic and electromagnetic fields, as well as other specific quantities. So far, PAK was applied in several different applications in mechanical engineering area—for testing materials, constructions, dam calculations, etc. In BioIRC, PAK solver was specifically adjusted to solve problems in bioengineering area [65,67,68,69]. PAK is used to simulate cardiovascular system—blood vessels and specifically coronary and carotid arteries [70], as well as for processing atherosclerosis and simulation of nanochips for drug supplies. It was used for simulation of stent implantation [71] and stress analysis. Further, models of the middle

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Figure 28.5 Geometrical parameters of the bioreactor model.

[72] and inner ear [73,74] are created and solved with PAK solver, many different models in dental area, several hip models [75]. The results of the performed calculation are stored in specific textual files, which can be loaded and visualized in developed CAD application. Back to the bioreactor model, the model was also simulated with commercial software COMSOL in order to verify obtained results with PAK solver. Compared results, obtained with two different finite element solvers, PAK and COMSOL, showed good matching [26]. This model is developed with the aim to become a part of the lung-on-a-chip model, where the bioreactor model will present blood vessel with circulating monocyte cells in order to analyze their response to the certain damage of the lung tissue cells—induction of the bacteria to the lung tissue or inflammation process related to the certain lungs disease or application of the certain drugs. The final model of the lung-on-a-chip system requests much more work to be performed, but at this initial stage, we have developed the mathematical model and numerical solution for it. Detailed explanation of derived mathematical model of bioreactor and performed numerical method are presented next. Behavior of the incompressible fluid is presented with Navier
Stokes equation [Eq. (28.2)], which is the mostly used equation for describing and modeling the

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behavior of the viscous fluid. It takes into account viscosity of the fluid and pressure. Solution of Navier
Stokes equation is fluid flow velocity, which means that we can know vector of fluid velocity at each point in space of modeled domain and at each time point from the chosen time interval. Navier
Stokes equation is the basic equation for modeling of pipelines, other water systems, blood flow, airflow around the airplane wing, etc. ρ

@V 2 μr2 V 1 ρðV  rÞV 1 rp 5 0 @t

(28.2)

where V represents fluid velocity vector, μ is the dynamic viscosity of the fluid, ρ is the fluid density, p is the fluid pressure, t means the time, and r stands for operator nabla. In case of the stationary flow, the first member of Eq. (28.2) is neglected (equal to zero). Expected fluid flow through the bioreactor is laminar, so profile of fluid velocity is parabolic. Medium input velocity is 4 mm=s at inlet boundary. Also, open end, meaning zero pressure, is assumed at the outlet boundary. Navier
Stokes equation is not the conservation equation. Therefore continuity equation is used as underlie, in order to achieve conservation of mass. The continuity equation for incompressible fluid has simplified form, since there is no change of fluid density in time [Eq. (28.3)]. r  V 50

(28.3)

Monocytes are modeled as a concentration inside the fluid and movements of the monocytes are presented as mass transfer in fluid. Their concentration follows Smoluchowski’s equation: @C 1 rð 2 DrC 1 CV Þ 5 0 @t

(28.4)

where D represents diffusion coefficient of the monocytes, C is the concentration (number of cells). As can be seen from Eq. (28.4), monocytes concentration changes in time but also depends on diffusion (the second term on the left-hand side, diffusive term) and fluid velocity (the third term on the left-hand side, advection term). These three equations make one coupled system, which ought to be solved. Values of the parameters used for simulation of the fluid flow and monocytes behavior inside of it are given in Table 28.2. The values of density and viscosity are chosen to be similar to the water, at the body temperature (37  C). FEM is applied to the model of the bioreactor. Eqs. (28.2)
(28.4), projected in Cartesian coordinate system, are written as follows:

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Table 28.2 Values of physical properties used for the model of the bioreactor. Parameter ρ μ V0 D C

Name Density of the fluid Dynamic viscosity Inlet velocity Diffusion coefficient Concentration of monocyte cells

Value

Unit 23

0.99  10 0.69  1023 4 4.5  1024 6  103

g/mm3 g/(mm s) mm/s mm2/s cells/mm3

 2    @Vx @ Vx @2 V x @2 V x @Vx @Vx @Vx @p 50 2μ 1 Vy 1 Vz ρ 1 1 1 ρ Vx 1 2 2 2 @x @t @x @y @z @x @y @z (28.5)  2    @Vy @ Vy @2 V y @2 V y @Vy @Vy @Vy @p 50 ρ 2μ 1 Vy 1 Vz 1 1 1 ρ Vx 1 2 2 2 @y @t @x @y @z @x @y @z (28.6)  2    @Vz @ Vz @2 V z @2 V z @Vz @Vz @Vz @p ρ 50 2μ 1 Vy 1 Vz 1 1 2 1 ρ Vx 1 2 2 @z @t @x @y @z @x @y @z (28.7) @Vx @Vy @Vz 1 1 50 @x @y @z

(28.8)

 2    @C @ C @2 C @2 C @C @C @C 2D Vx 1 Vy 1 Vz 5 0 1 2 1 2 1 @t @x2 @y @z @x @y @z

(28.9)

In order to obtain finite element matrices, Eqs. (28.5)
(28.9) are multiplied with interpolation functions hI and integrated over the volume of finite element. Quantity to be solved (V; p; andC) can be presented as a product of interpolation functions and values of the quantity at the appropriate nodes [Eq. (28.10)] Vi 5 hJ ViJ ; i 5 1; 2; 3; p 5 h^J PJ ; C 5 hJ C J

(28.10)

^ differ from the interpolation functions Interpolation functions for pressure, (h), used for velocity and monocyte concentration, (h), because for pressure we are using only corner nodes. Therefore instead of 21 nodes used for interpolation fluid velocity (3D elements), we are using 8 nodes for interpolation of pressure. First, Navier
Stokes equation is written in index notation, for example, i; j 5 1; 2; 3, meaning x; y; z directions, and then derived.

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ð hI ρ V

@Vi dV 2 @t

ð

ð

ð

hI μVi;jj dV 1 V

hI ρVj Vi;j dV 1 V

hI p;i dV 5 0

(28.11)

V


 ð ð  

  @ hJ ViJ J dV 2 hI μ hJ Vi ;jj dV 1 hI ρ hJ VjJ hJ ViJ ;j dV hI ρ @t V V V ð J 1 hI ðh^J P Þ;i dV 5 0

ð

(28.12)

V

Interpolation functions are functions of coordinates and are not time dependent. The second member of Eq. (28.12) contains the second partial derivative of velocities by coordinates, so we can apply some operation, that is, Gauss’s theorem, on that member in order to obtain simplified form. ð V


 hI μ hJ ViJ ;jj dV 5

ð

ð 
 
 μ hI  hJ ViJ ;j dV 2 μhI;j hJ ViJ ;j dV ;j V ðV ð
 J 5 μhI  ðhJ Vi Þ;j  nj dS 2 μhI;j hJ ViJ ;j dV V ð ðS 5 μqs  nj dS 2 μhI;j hJ;j ViJ dV S

(28.13)

V

Gauss’s theorem is also applied on the integral from Eq. (28.12) containing fluid pressure. ð V

hI ðh^J PJ Þ;i dV 5

ð V

ð ð ð J J J ^ ^ ^ ðhI h J P Þ;i dV 2 hI;i h J P dV 5 hI h J P ni dS 2 hI;i h^J PJ dV V

S

V

(28.14) By substituting developed form of the second velocity derivate integral and pressure integral [Eqs. (28.13) and (28.14)], we obtain the final form of Navier
Stokes equation [Eq. (28.15)] ð

ð ð ð J ρhI hJ dV V_ i 1 μhI;j hJ;j dVViJ 1 ρhI ðhJ VjJ ÞhJ;j dVViJ 2 hI;i h^J dVPJ V V V V ð ð 5 μqs nj dS 2 hI h^J PJ ni dS S

S

(28.15) Eq. (28.15) is also written in matrix form [Eq. (28.16)] M V_ i 1 ðKVMi 1 KVRo ÞViJ 1 KVP PJ 5 FS 5 0 J

Matrices from Eq. (28.16) are written next:

(28.16)

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ð M 5 ρ hI hJ dV

(28.17)

V

KVV 5 KVMi 1 KVRo

(28.18)

ð KVMi 5 μ

hI;j hJ;j dV

(28.19)

  hI hJ VjJ hJ;j dV

(28.20)

hI;i h^J dV

(28.21)

V

ð KVRo 5 ρ V

ð KVP 5 2 V

In Eq. (28.16) on the right-hand side, there is a force, FS , representing surface force, which is equal to zero due to natural boundary condition. Continuity equation is written in index notation and derived then: Vi;i 5 0 ð

(28.22)

h^I Vi;i dV 5 0

(28.23)


 h^I hJ ViJ ;i dV 5 0

(28.24)

h^I hJ;i dVViJ 5 KPV ViJ 5 0

(28.25)

V

ð V

ð V

According to obtained matrix KPV it can be concluded that it has the form of transpose matrix KVP . T KPV 5 KVP

(28.26)

Mass transfer in index notation and derivation is presented next. @C 2 DC;ii 1 C;i Vi 5 0 @t ð hI V

@ðhJ C J Þ dV 2 @t

ð V

(28.27) ð

hI DðhJ CJ Þ;ii dV 1

V

hI ðhJ CJ Þ;i ðhJ ViJ ÞdV 5 0

(28.28)

Gauss’s theorem is applied on the integral from Eq. (28.28), containing the second derivative of monocytes concentration. Due to already presented utilization of

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Gauss’s theorem, application to the diffusive term of the monocytes concentration is omitted, but developed form is applied to Eq. (28.28) in order to obtain the final form of diffusion-advection equation for modeling monocytes behavior [Eq. (28.29)]. ð ð ð ð J J J J _ hI hJ dV C 1 DhI;i hJ;i dVC 1 hI hJ;i C hJ Vi dV 5 Dqn ni dS 5 0 V

V

V

S

(28.29) Ð Surface force (FS 5 S Dqn ni dS) from Eq. (28.29) is equal to zero, since mass transfer through the walls of the bioreactor is disabled. Matrices from Eq. (28.29) are listed as follows: ð C5

hI hJ dV

(28.30)

V

ð KCV 5

hI hJ;j C J hJ dV

(28.31)

DhI;i hJ;i dV

(28.32)

V

ð KCC 5 V

Finally, obtained system of the equations in matrix form is presented with the following: 2

M 0 4 0 0 0 0

38 9 2 KVV 0 < V_ = 0 5 P_ 1 4 KPV : _; C KCV C

KVP 0 0

38 9 8 9 0