Principles of Regenerative Medicine [3rd Edition] 9780128098936

Virtually any disease that results from malfunctioning, damaged, or failing tissues may be potentially cured through reg

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Principles of Regenerative Medicine [3rd Edition]
 9780128098936

Table of contents :
Front Cover
......Page 1
PRINCIPLES OF REGENERATIVE MEDICINE......Page 2
PRINCIPLES OF REGENERATIVE MEDICINE......Page 4
Copyright......Page 5
Dedication......Page 6
Contents......Page 8
Contributors......Page 20
Preface......Page 26
MOLECULES THAT ORGANIZE CELLS......Page 28
Changes in Cell–Cell Adhesion......Page 29
Invasion of the Basal Lamina......Page 30
Transcription Factors That Regulate Epithelial–Mesenchymal Transition......Page 31
MOLECULAR CONTROL OF THE EPITHELIAL–MESENCHYMAL TRANSITION......Page 32
Wnt Pathway......Page 33
Additional Signaling Pathways......Page 34
CONCLUSION......Page 35
References......Page 36
COMPOSITION AND DIVERSITY OF THE EXTRACELLULAR MATRIX......Page 42
RECEPTORS FOR EXTRACELLULAR MATRIX MOLECULES......Page 43
SIGNAL TRANSDUCTION EVENTS DURING CELL–EXTRACELLULAR MATRIX INTERACTIONS......Page 45
Adhesion and Migration......Page 46
Proliferation and Survival......Page 49
Differentiation......Page 50
Apoptosis......Page 51
Adhesion and Migration......Page 52
Differentiation......Page 54
CELL–EXTRACELLULAR MATRIX INTERACTIONS DURING REGENERATIVE FETAL WOUND HEALING......Page 55
Proliferation......Page 56
IMPLICATIONS FOR REGENERATIVE MEDICINE......Page 57
References......Page 58
BLASTEMA FORMATION......Page 64
Hemostasis and Reepithelialization......Page 65
Mechanisms of Dedifferentiation......Page 66
Cell Cycling During Blastema Formation......Page 67
Blastema Cell Migration and Accumulation......Page 68
The Apical Epidermal Cap–Nerve Interaction......Page 69
Interaction of Cells From Opposite Sides of the Limb Circumference......Page 71
References......Page 72
GROUND STATE AND PRIMED EMBRYONIC STEM CELLS HAVE UNIQUE SIGNALING NETWORKS UNDERLYING PLURIPOTENCY......Page 76
LEUKEMIA INHIBITORY FACTOR AND BONE MORPHOGENIC PROTEIN SIGNALING PATHWAYS REGULATE MOUSE EMBRYONIC STEM CELL SELF-RENEWAL......Page 77
WNT SIGNALING CONTRIBUTES TO MAINTENANCE OF PLURIPOTENCY IN MOUSE EMBRYONIC STEM CELLS AND TO THE NAIVE HUMAN EMBRYONIC STE .........Page 79
THREE TRANSCRIPTION FACTORS, OCTAMER BINDING PROTEIN 4, SRY-BOX 2, AND NANOG, FORM THE CORE PLURIPOTENCY TRANSCRIPTIONAL NE .........Page 80
MYC LINKS CELL SIGNALING TO PLURIPOTENCY GENE REGULATION......Page 81
A SPECIFIC EPIGENETIC PROGRAM HELPS MAINTAIN PLURIPOTENCY......Page 82
MICRORNAS INTEGRATE WITH CELL SIGNALING AND TRANSCRIPTION FACTORS TO REGULATE STEM CELL PROLIFERATION AND DIFFERENTIATION......Page 84
CONCLUSIONS......Page 85
References......Page 86
INTRODUCTION......Page 92
Adult Wound Healing and Scar Formation......Page 93
Fibroproliferative Scarring......Page 95
Hypertrophic Scars......Page 97
Fetal Scarless Wound Repair......Page 99
Targeting the Inflammatory Response......Page 100
Connective Tissue Growth Factor......Page 102
Wingless Type Signaling......Page 103
5-Fluorouracil......Page 104
Bleomycin......Page 105
Cryotherapy......Page 106
Growth Factors and Cell Signaling Molecules......Page 107
Embryonic Stem Cells......Page 108
Mesenchymal Stem Cells......Page 109
Epidermal Stem Cells......Page 110
Induced Pluripotent Stem Cells......Page 111
PERSPECTIVE......Page 112
References......Page 113
INTRODUCTION......Page 120
SINGLE-CELL ISOLATION......Page 122
ACQUIRING SINGLE-CELL DATA......Page 123
Single-Cell Transcriptomics......Page 124
Single-Cell Proteomics......Page 126
Reducing Noise in Single-Cell Data......Page 127
Mathematical Identification of Cellular Subpopulations......Page 128
DETERMINING SUBPOPULATIONS......Page 130
Development of Cell-Based Therapies......Page 131
CLINICAL IMPLICATIONS OF CELLULAR HETEROGENEITY IN TISSUE REPAIR AND DISEASE......Page 132
Cellular Heterogeneity in Wound Healing......Page 133
Cellular Heterogeneity in Aging......Page 134
CONCLUSIONS......Page 135
References......Page 136
Mouse Embryonic Stem Cells......Page 140
Blastocyst......Page 141
Parthenogenesis......Page 142
Microenvironment......Page 143
Maintenance......Page 144
Evolution of Human Embryonic Stem Cell Derivation and Culture Methods......Page 145
HUMAN EMBRYONIC STEM CELL DIFFERENTIATION AND MANUFACTURING FOR CLINICAL APPLICATION......Page 146
References......Page 147
Further Reading......Page 150
INTRODUCTION......Page 152
Single Blastomere Biopsy......Page 153
Irreversibility as a Criterion for Diagnosing Embryonic Death......Page 154
Morphological Criteria for Predicting the Capacity of Irreversibly Arrested, Nonviable Human Embryos to Develop Into a Huma .........Page 155
References......Page 157
PLACENTA: FUNCTION, ORIGIN, AND COMPOSITION......Page 160
AMNIOTIC EPITHELIAL CELLS......Page 161
AMNIOTIC MESENCHYMAL STEM CELLS......Page 162
Preclinical Studies......Page 163
Characterization......Page 165
Preclinical Studies......Page 167
Heart......Page 168
Kidney......Page 169
Intestine......Page 170
References......Page 171
A BRIEF HISTORY......Page 176
PUBLIC VERSUS FAMILY (OR PRIVATE) BANKS......Page 177
Donor Recruitment and Consent......Page 178
Volume and Cell Count Considerations......Page 179
Processing and Cryopreservation......Page 180
Cord Blood Unit Characterization......Page 181
Cord Blood Transplantation for Nonmalignant Hematological Diseases......Page 183
Cord Blood Transplantation for Inherited Metabolic Disorders......Page 184
Cerebral Palsy......Page 186
Stroke......Page 187
Autism Spectrum Disorder......Page 188
References......Page 189
MECHANISMS OF REPROGRAMMING......Page 196
REPROGRAMMING TECHNIQUES......Page 197
INDUCED TRANSDIFFERENTIATION......Page 198
DISEASE MODELING......Page 199
PERSONALIZED MEDICINE......Page 201
CELL THERAPY......Page 202
CONCLUSIONS AND FUTURE DIRECTIONS......Page 203
References......Page 204
ADULT STEM CELLS......Page 208
ISOLATION OF RODENT MULTIPOTENT ADULT PROGENITOR CELL......Page 209
Hematopoietic Reconstitution With Multipotent Adult Progenitor Cells......Page 210
Effect of Multipotent Adult Progenitor Cells on Graft Versus Host Disease......Page 211
Multipotent Adult Progenitor Cell Immunodulatory and/or Trophic Effects in Ischemic Disease......Page 212
Possible Mechanisms of Trophic Effects: Secreted Proteome of Multipotent Adult Progenitor Cells......Page 213
References......Page 214
HEMATOPOIETIC STEM CELL PROPERTIES......Page 218
Fetal Liver Hematopoiesis......Page 219
In Vitro Hematopoiesis......Page 220
Phenotypic Properties of Hematopoietic Stem Cells......Page 221
Bone Marrow Transplantation......Page 222
Autologous Peripheral Blood Stem Cell Transplantation......Page 223
Hematopoietic Stem Cell Transplantation for Severe Combined Immunodeficiency......Page 224
Hematopoietic Stem Cell Transplantation for Tolerance Induction......Page 225
List of Acronyms and Abbreviations......Page 226
References......Page 227
INTRODUCTORY OVERVIEW......Page 232
THE STEM CELL NATURE OF MESENCHYMAL STEM CELLS......Page 233
WHICH TISSUES CONTAIN MESENCHYMAL STEM CELLS?......Page 234
MESENCHYMAL STEM CELL EXOSOMES......Page 235
IMMUNOMODULATORY EFFECTS OF MESENCHYMAL STEM CELLS......Page 238
INDUCED PLURIPOTENT STEM CELL–DERIVED MESENCHYMAL STEM CELLS......Page 239
Acknowledgments......Page 240
References......Page 241
INTRODUCTION AND HISTORY......Page 246
NEW INSIGHT......Page 247
Clinically Relevant Therapies Using Mesenchymal Stem Cells......Page 249
Diabetes......Page 250
THE NEW MESENCHYMAL STEM CELLS......Page 251
References......Page 252
INTRODUCTION......Page 256
Integration of Hepatocytes After Transplantation......Page 258
CLINICAL HEPATOCYTE TRANSPLANTATION......Page 259
Hepatocyte Transplantation in Acute Liver Failure......Page 260
Hepatocyte Transplantation for Metabolic Liver Disease......Page 261
Hepatocyte Transplants for Non–Organ Transplant Candidates......Page 264
Methods to Improve Engraftment and Repopulation......Page 265
Stem Cells and Alternative Cell Sources for Liver Therapy......Page 266
SUMMARY......Page 268
References......Page 269
DEVELOPMENT OF THE HEART FROM CARDIAC STEM/PROGENITOR CELLS......Page 274
c-Kit+ Cardiac Progenitor/Stem Cells......Page 275
Cardiac Neural Crest–Derived Progenitors......Page 277
Epicardial Progenitor Cells......Page 278
CELL-BASED THERAPEUTICS FOR HEART DISEASE......Page 279
MECHANISMS OF ACTION......Page 281
Pluripotent Stem Cells......Page 282
Adult Stem Cells......Page 283
Endothelial Progenitor Cells......Page 284
Mesenchymal Stem Cells......Page 285
Cardiac Stem Cells......Page 288
Other Cardiac Stem Cells......Page 290
COMBINED STEM CELL THERAPEUTICS......Page 292
References......Page 294
INTRODUCTION......Page 300
THE MOLECULAR CHARACTERISTICS OF MUSCLE STEM CELLS DURING MYOGENESIS IN REGENERATION......Page 301
FUNCTIONAL CHARACTERISTICS OF MUSCLE STEM CELLS......Page 303
ISOLATION OF MUSCLE STEM CELLS......Page 304
TRACKING MUSCLE STEM CELL BEHAVIOR THROUGH LIVE IMAGING (BIOLUMINESCENCE IMAGING AND INTRAVITAL IMAGING)......Page 305
Extracellular Matrix Components......Page 306
Biophysical Cues......Page 307
SATELLITE CELL SELF-RENEWAL MECHANISMS......Page 308
MUSCLE STEM CELL–INTRINSIC DEFECTS IN AGING AND DISEASE......Page 310
CHALLENGES IN THE USE OF SATELLITE CELLS IN REGENERATIVE MEDICINE......Page 311
OTHER STEM CELL TYPES WITHIN MUSCLE......Page 312
Induced Pluripotent Stem–Derived Muscle Stem Cells......Page 313
References......Page 314
CELLULAR FRACTIONS......Page 322
Adipose-Derived Stromal Cell......Page 323
CLINICAL DELIVERY OF ADIPOSE-DERIVED CELLS......Page 324
ENGINEERED NEO-TISSUE......Page 327
Carcinogenesis and Tumorigenesis......Page 328
References......Page 329
TYPES AND SOURCE OF STEM CELLS IN THE PERIPHERAL BLOOD......Page 334
Mobilization of Bone Marrow Cells......Page 335
Identification and Isolation of Endothelial Progenitor Cells......Page 338
In Vitro Expansion of Endothelial Progenitor Cells......Page 339
The Role of Endothelial Progenitor Cells in Physiological and Pathological Neovascularization......Page 340
Identification, Isolation, Characterization, and In Vitro Expansion......Page 342
Tissue Regeneration......Page 344
Tissue Engineering......Page 346
Mesenchymal Stem Cells......Page 347
The Use of Mesenchymal Stem/Marrow Stroma Cells for Gene Therapy......Page 349
CONCLUSIONS AND FUTURE DIRECTIONS......Page 352
References......Page 353
FROM ADULT PANCREATIC ISLETS TO STEM CELLS......Page 362
β CELLS FROM PLURIPOTENT STEM CELLS (EMBRYONIC STEM CELLS AND INDUCED PLURIPOTENT STEM CELLS)......Page 363
β CELLS FROM ADULT STEM/PROGENITOR CELLS OF THE BILIARY TREE AND PANCREAS......Page 368
MESENCHYMAL STEM CELLS TO MODULATE IMMUNITY AND PROMOTE TISSUE REPAIR IN DIABETES......Page 371
CONCLUSION......Page 372
References......Page 373
INTRODUCTION......Page 378
The Retina......Page 379
Retinitis Pigmentosa......Page 381
Human Embryonic Stem Cell–Derived Retinal Pigment Epithelium......Page 382
Induced Pluripotent Stem Cell–Derived Retinal Pigment Epithelium......Page 383
Scaffolds for Retinal Pigment Epithelium Transplantation......Page 386
Photoreceptor Transplantation......Page 387
CELL-BASED NEUROPROTECTION......Page 388
DISEASE-IN-A-DISH MODELING FOR RETINAL DISORDERS......Page 389
Three-Dimensional Retinal Organoids......Page 390
CONCLUSION......Page 391
References......Page 392
Epidemiology......Page 396
Primary Versus Secondary Brain Injury......Page 397
Neuroinflammation......Page 398
Blood–Brain Barrier Permeability......Page 400
Cerebral Edema......Page 402
Mechanisms of Action......Page 403
Timing of Infusion......Page 404
Conventional Cell Delivery Routes, Continued......Page 405
Novel Cell Delivery Routes......Page 407
Reduction in Therapeutic Intensity: Pediatric Intensity Level of Therapy Scores......Page 408
Imaging Data......Page 409
Results......Page 410
Phase 1/2 Adipose-Derived Stem/Stromal Cells......Page 411
CONCLUSION......Page 412
Acknowledgments......Page 413
References......Page 414
INTRODUCTION......Page 418
Extracellular Matrix......Page 419
Ion Channels and Mechanoreceptors......Page 420
Cytoskeleton......Page 421
NUCLEUS AS THE CENTRAL ORGANELLE IN REGULATING MECHANOTRANSDUCTION......Page 423
CELLULAR MECHANOTRANSDUCTION MECHANISMS......Page 424
Mechanotransduction Through Cell–Cell Adhesions......Page 425
From Cells to Organs: How Mechanobiology Affects Tissue Development and Function......Page 426
CONCLUSIONS......Page 427
References......Page 428
INTRODUCTION......Page 432
BONE MORPHOGENETIC PROTEINS......Page 433
SCAFFOLDS OF EXTRACELLULAR MATRIX AND BIOMIMETIC BIOMATERIALS......Page 436
REGENERATIVE MEDICINE AND SURGERY OF ARTICULAR CARTILAGE......Page 439
REGENERATION OF ARTICULAR CARTILAGE SURFACE AND LUBRICATION......Page 440
References......Page 441
INTRODUCTION......Page 444
Strain......Page 445
Constitutive Relations......Page 448
Tissue Remodeling......Page 450
Mechanotransduction......Page 451
Mechanical Stimulation In Vivo......Page 454
Bone Bioreactors......Page 456
Blood Vessel Bioreactors......Page 458
CONCLUSIONS......Page 459
References......Page 460
CELL–EXTRACELLULAR MATRIX INTERACTIONS......Page 464
Effect of Physical Properties......Page 465
Stiffness and Compliance......Page 466
Surface Charge......Page 467
Methods of Altering Surface Chemistry......Page 468
Development of Bioactive Surfaces......Page 469
Cell Adhesion......Page 470
Cell Motility......Page 471
Cell Proliferation, Self-renewal, and Differentiation......Page 472
Fabrication Techniques......Page 473
Cellular Responses to Topographical Cues......Page 474
Electrically Conductive Substrate......Page 477
EFFECT OF DIMENSIONALITY......Page 478
Hydrogel Scaffolds......Page 479
Decellularized Tissue......Page 480
New Technology Development......Page 481
Cellular Responses to Three-Dimensional Substrates......Page 482
Cell Migration......Page 483
Effect of Externally Applied Mechanical Stimuli......Page 484
Mechanotransduction......Page 485
Cellular Responses in Modifying Extracellular Matrix......Page 486
CONCLUSION......Page 487
References......Page 488
Thermoresponsive Polymer for Biomedical Applications......Page 496
Controlled Grafting of Thermoresponsive Polymer on Culture Substrates......Page 497
Variety of Fabrication Techniques of Thermoresponsive Cell Culture Substrate......Page 498
Cornea Reconstruction......Page 499
Myocardium Regeneration......Page 500
Cell Sheet Layering Technique......Page 501
Vascularization in Cell Sheets for Large-scale Tissue Construction......Page 502
COMBINATION OF CELL SHEET ENGINEERING AND SCAFFOLD-BASED ENGINEERING......Page 504
Copatterning to Create a Cellular Microenvironment......Page 505
Intelligent Surfaces for Regulating Cell Orientation......Page 506
Skeletal Muscle Tissue Engineering......Page 507
CONCLUSIONS......Page 508
References......Page 509
INTRODUCTION......Page 512
Physical Properties......Page 513
Size......Page 514
Shape......Page 515
Surface Topography......Page 516
Optical Properties......Page 517
NANOBIOMATERIALS......Page 518
Bone Tissue......Page 520
Muscle Tissue......Page 521
Vascular Tissue......Page 522
Other Tissue......Page 523
Stem Cell Transfection......Page 524
Stem Cell Expansion......Page 525
References......Page 526
Mechanical Support......Page 532
Degradation Mechanisms......Page 534
Factors That Affect Degradation Rates......Page 535
Surface Modification for Degradation Control......Page 536
On-Demand Release......Page 537
Anisotropic and Gradient Scaffolds......Page 538
Surface Feature Manipulation......Page 539
SAFETY AND BIOCOMPATIBILITY REQUIREMENTS FOR BIOMATERIAL SCAFFOLDS......Page 540
Infection and Sterilization......Page 541
Hemocompatibility......Page 542
Foreign Body Response......Page 543
SUMMARY......Page 544
References......Page 545
WHY THE NEED FOR PRECISION CONTROL OF PROTEINS AT INTERFACES IN TISSUE ENGINEERING AND REGENERATIVE MEDICINE?......Page 550
SURFACE ANALYSIS AND ITS ROLE IN THE PRECISION DELIVERY OF BIOLOGICAL SIGNALS......Page 551
Sum Frequency Generation......Page 552
Quartz Crystal Microbalance With Dissipation Monitoring......Page 553
TECHNIQUES AND TECHNOLOGIES FOR PRECISION IMMOBILIZATION AT SURFACES......Page 554
Ionic Charge and Charge Control of Orientation......Page 555
Collagen to Control Protein Orientation......Page 556
Streptavidin for Biomolecular Orientation Control......Page 557
CONCLUSIONS......Page 558
References......Page 559
INTRODUCTION......Page 562
Processing Methods......Page 565
COLLAGEN......Page 567
Processing Methods......Page 568
Collagen in Bone Tissue Engineering Applications......Page 569
Processing Methods......Page 570
Gellan Gum in Bone Tissue Engineering Applications......Page 571
Processing Methods......Page 572
Polyhydroxyalkanoates in Bone Tissue Engineering Applications......Page 573
Silk Fibroin in Bone Tissue Engineering Applications......Page 574
STARCH......Page 575
Processing Methods......Page 576
NATURAL-BASED BIOCERAMICS......Page 577
CALCIUM PHOSPHATES......Page 578
Calcium Phosphate in Bone Tissue Engineering Applications......Page 579
Silicate in Bone Tissue Engineering Applications......Page 580
CONCLUSIONS......Page 581
References......Page 582
INTRODUCTION......Page 586
POLYMER SYNTHESIS......Page 587
Poly(ethylene), Poly(propylene), and Poly(styrene)......Page 588
Poly(meth)acrylates and Polyacrylamides......Page 589
Poly(N-isopropylacrylamide)......Page 590
Polyethers......Page 591
Polysiloxanes......Page 592
Hydrolytically Stable Polyurethanes......Page 593
Polyesters......Page 594
Polyesters of α-Hydroxy Acids......Page 595
Polyesters of Lactones......Page 597
Polyorthoesters......Page 598
Polyurethanes......Page 599
Amino Acid–Derived Polymers, Poly(amino Acids), and Peptides......Page 600
Polyanhydrides......Page 601
Biodegradable Cross-linked Polymer Networks......Page 602
Cross-linked Polyesters......Page 603
CONCLUSION/SUMMARY......Page 607
References......Page 608
Calcium Phosphate Bioceramics......Page 618
Basic Properties......Page 620
Apatite Cements......Page 621
Setting/Hardening Mechanism......Page 622
Hydrolysis Interaction......Page 623
Setting Times......Page 624
Strategies to Improve Setting Times......Page 625
Strategies to Improve Injectability......Page 626
Liquid-to-Powder Ratio......Page 627
STRATEGIES TO IMPROVE THE MECHANICAL PROPERTIES......Page 628
Porosity......Page 629
Dual Setting System......Page 631
Mechanics of Fiber-Reinforced Calcium Phosphate Cements......Page 632
Oral, Maxillofacial, and Craniofacial Applications......Page 633
CONCLUSION......Page 634
References......Page 635
EXTRACELLULAR MATRIX: FUNCTION AND COMPONENTS......Page 640
Collagen......Page 642
Fibronectin......Page 643
Glycosaminoglycans/Proteoglycans......Page 644
Matrix-Bound Nanovesicles......Page 645
Decellularization......Page 646
Hydrogels......Page 647
Whole-Organ Scaffolds......Page 648
List of Acronyms and Abbreviations......Page 649
References......Page 650
INTRODUCTION......Page 654
BIOMATERIALS TEMPLATES......Page 655
STRUCTURE–PROPERTY RELATIONSHIPS IN HYDROGELS......Page 658
Bioactive Forms of Poly(ethylene Glycol) as Exemplars of Increasing Sophistication......Page 659
Spatial Heterogeneity......Page 660
Matrix Mechanics......Page 662
Hydrogel Degradation......Page 663
Polymerization Mechanisms......Page 664
Injectable Systems......Page 665
Hyaluronic Acid......Page 666
Alginate......Page 667
Cellulose......Page 668
Collagen and Its Derivatives......Page 669
Elastin Derivatives......Page 670
Fibrin Derivatives......Page 671
Self-assembled Peptides......Page 672
SYNTHETIC HYDROGELS FOR TISSUE ENGINEERING TEMPLATES......Page 673
CONCLUSIONS......Page 675
References......Page 676
Overview of Surface Modification Strategies......Page 678
Topographical Modifications......Page 680
Noncovalent Coatings......Page 682
BIOLOGICAL MODIFICATION OF SURFACES......Page 683
References......Page 686
TISSUE COMPONENTS......Page 688
REGENERATION OF DISEASED TISSUES......Page 689
Cell Sources......Page 690
Porosity......Page 691
Degradation......Page 693
Importance of Microvasculature......Page 694
Hydrogels......Page 696
CONCLUSIONS......Page 697
References......Page 698
INTRODUCTION......Page 702
Blood–Material Interactions and Initiation of the Inflammatory Response......Page 703
Provisional Matrix Formation......Page 704
Temporal Sequence of Inflammation and Wound Healing......Page 705
Chronic Inflammation......Page 706
Macrophage Interactions......Page 707
Foreign Body Giant Cell Formation and Interactions......Page 709
FIBROSIS AND FIBROUS ENCAPSULATION......Page 710
IMMUNOTOXICITY (ACQUIRED IMMUNITY)......Page 711
References......Page 718
Further Reading......Page 721
INTRODUCTION......Page 722
FUNCTIONS OF SCAFFOLDING AND EXTRACELLULAR MATRIX......Page 723
SCAFFOLDING APPROACHES IN BONE TISSUE ENGINEERING......Page 724
Hydrogels......Page 725
Silk......Page 726
Collagen......Page 727
Hyaluronic Acid......Page 728
Alginate......Page 729
Peptide Hydrogels......Page 730
Copolymers......Page 731
Ceramic Scaffolds......Page 732
Bioglass......Page 733
Metallic Scaffolds......Page 734
Polymer–Ceramics Blends......Page 735
Metal–Ceramic Blends......Page 736
References......Page 737
INTRODUCTION AND OVERVIEW OF CANCER IMMUNOTHERAPY......Page 742
ADVANTAGES AND DISADVANTAGES OF CANCER IMMUNOTHERAPY......Page 744
Introduction of Nanomedicine in Cancer......Page 745
Effects of Nanoparticle Surface Functionalization......Page 747
Nanoparticle Targeting of the Tumor Microenvironment......Page 748
Nanoparticle Targeting of Antigen Presenting Cells......Page 749
Implantable Biomaterial Scaffolds as Cancer Vaccines......Page 754
Injectable Biomaterial Systems as Cancer Vaccines......Page 756
Implantable Biomaterial Scaffolds to Enhance Autologous T Cell Therapy......Page 758
CONCLUSION......Page 760
Glossary......Page 761
References......Page 763
Clustered Regularly Interspaced Short Palindromic Repeats......Page 768
Knockouts via Double-Strand Breaks......Page 769
Nickases......Page 770
Homology-Directed Repair......Page 771
SpCas 9 Variants and Orthologues......Page 772
Transcription Activator-like Effector Nucleases......Page 773
Recombinase......Page 774
Proteins......Page 775
DELIVERY METHODS......Page 776
Liver......Page 777
Muscle: Muscular Dystrophy......Page 778
Duchenne Muscular Dystrophy......Page 779
Blood......Page 780
Retina......Page 781
References......Page 782
BIOMINERALIZATION AND BONE REGENERATION......Page 788
Mesenchymal Stem Cells......Page 789
Biochemical Signaling: Growth Factors and Cell Signals......Page 790
In Vivo Preclinical Models......Page 791
Selection Considerations Based on Animal Species......Page 792
References......Page 793
INTRODUCTION......Page 796
ADVANCE OF IN VITRO ORGANOID DEVELOPMENT: PROGRESSION FROM TWO-DIMENSIONAL TO THREE-DIMENSIONAL MODELS......Page 797
Microengineering and Biofabrication......Page 799
Vessel-on-a-Chip......Page 800
Cancer-on-a-Chip......Page 801
BODY-ON-A-CHIP: MULTIORGAN SYSTEMS AND FUTURE APPLICATIONS......Page 802
Cancer......Page 803
Drug Testing and Toxicology......Page 804
Additional Disease Modeling......Page 805
The Ex Vivo Console of Human Organoids Platform......Page 806
Other Body-on-a-Chip Programs......Page 807
Organ-on-a-Chip Systems for Personalized Precision Medicine......Page 809
References......Page 810
DESIGN CONSIDERATIONS FOR CREATING BIOREACTORS......Page 814
Bioengineering Functional Lungs......Page 815
Bioreactors for Regeneration of Small Animal Lungs......Page 816
In Vivo Bioreactors for Lung Regeneration......Page 817
Bioreactors for Study of Lung Biology......Page 819
Evaluation of Bioengineered Lungs......Page 820
Perfusion Bioreactors for Bone Regeneration......Page 822
In Vivo Bone Bioreactors for Solving the Vascularization Problem......Page 824
Bioreactors for Studying Bone Development and Disease......Page 825
Monitoring the Environment and Tissue Development Within Bioreactors......Page 826
References......Page 828
FUNDAMENTALS OF THREE-DIMENSIONAL PRINTING......Page 832
Extrusion-Based Printing......Page 833
Inkjet Bioprinting......Page 834
BIOINKS......Page 835
Matrix or Matrix-Mimicking Bioinks......Page 837
Synthetic Materials......Page 838
Natural Materials......Page 840
Co-printing and Hybrid Bioinks......Page 843
Cell-Laden Bioinks......Page 846
Sacrificial Bioinks......Page 848
Supporting Bioinks and Supporting Baths......Page 850
In Vitro Applications......Page 852
CONCLUSION AND FUTURE DIRECTIONS......Page 853
References......Page 854
BIOPRINTING STRATEGY: FROM MEDICAL IMAGE TO PRINTED TISSUE......Page 858
Jetting-Based Printing......Page 859
Hybrid and Other Mechanisms......Page 861
Synthetic Hydrogels......Page 862
Naturally Derived Hydrogels......Page 863
Biodegradable Synthetic Polymers for Structural Integrity......Page 864
Three-Dimensional Bioprinted Vascular Structures......Page 865
Tumor Models......Page 866
Bone......Page 868
Cartilage......Page 869
Skeletal Muscle and Tendon......Page 872
Skin......Page 873
CONCLUSIONS AND FUTURE PERSPECTIVES......Page 874
Glossary......Page 875
References......Page 876
Fracture Healing......Page 880
Adipose-Derived Stem Cells......Page 881
Induced Pluripotent Stem Cells......Page 882
Porous and Highly Interconnected Scaffolds......Page 883
Nanofibrous Scaffolds for Bone Tissue Engineering......Page 884
Hydrogels......Page 885
Bone Morphogenetic Proteins......Page 886
Nucleotide Delivery and Gene Therapy......Page 887
IMMUNOMODULATION IN BONE REGENERATION......Page 888
T Cells......Page 889
References......Page 890
STRUCTURE OF THE INNER EAR......Page 894
HAIR CELL LOSS......Page 895
HISTORY OF HAIR CELL REGENERATION......Page 896
SPONTANEOUS HAIR CELL REGENERATION IN MAMMALIAN VESTIBULAR ORGANS......Page 897
INSIGHTS FROM DEVELOPMENTAL BIOLOGY......Page 898
INDUCTION OF HAIR CELL REGENERATION USING TRANSGENIC MICE......Page 902
STUDIES OF HAIR CELL REGENERATION USING THE LATERAL LINE......Page 903
FORMATION OF NEW NEUROMASTS FROM MULTIPOTENT PROGENITORS......Page 904
HAIR CELL REGENERATION IN THE LATERAL LINE......Page 905
PATHWAYS COORDINATING HAIR CELL REGENERATION IN THE LATERAL LINE......Page 906
OPEN QUESTIONS ABOUT LATERAL LINE REGENERATION......Page 907
References......Page 908
UNDERSTANDING THE CRANIOFACIAL REGENERATIVE ENVIRONMENT......Page 914
CURRENT METHODS OF MAXILLOFACIAL RECONSTRUCTION......Page 917
Ceramics......Page 918
Bioactive Molecules......Page 919
Platelet-Derived Growth Factor......Page 920
Bone Marrow Aspirate Concentrate Technique......Page 921
Bioreactors......Page 924
Antibiotics......Page 925
CONCLUSION......Page 926
List of Abbreviations......Page 928
References......Page 929
INTRODUCTION......Page 934
TOOTH DEVELOPMENT......Page 935
DENTAL STEM CELLS......Page 936
DENTAL TISSUE ENGINEERING......Page 937
Whole Tooth Engineering......Page 939
Dental Pulp and Dentin Regeneration......Page 940
Periodontal Regeneration......Page 941
Alveolar Bone Regeneration......Page 943
List of Abbreviations......Page 944
References......Page 945
INTRODUCTION......Page 950
Red Blood Cells Generated From Adult Stem Cells In Vitro......Page 951
Red Blood Cells Generated From Human Embryonic Stem Cells......Page 952
Red Blood Cells Generated From Human Induced Pluripotent Stem Cells......Page 954
Where Do We Go From Here?......Page 955
Generation of Megakaryocytes and Platelets From Adult Stem Cells and Somatic Cells......Page 956
Improving the Efficiency for In Vitro Platelet Production......Page 957
HEMATOPOIETIC STEM CELLS......Page 958
References......Page 960
Further Reading......Page 963
CARTILAGE AND CARTILAGE REPAIR......Page 964
Cartilage Surface Modification......Page 965
Bioscaffolds in Cartilage Repair......Page 966
Chitosan......Page 967
Synthetic Scaffolds......Page 968
Biological Factors......Page 969
Bioreactors......Page 970
Clinical Translation......Page 972
CURRENT AND FUTURE TRENDS IN CARTILAGE ENGINEERING......Page 973
References......Page 974
Regulatory and Financial Challenges to Stem Cell Therapies......Page 980
STEM CELL THERAPIES FOR MUSCULOSKELETAL DISEASES......Page 981
Bone......Page 982
Articular Cartilage......Page 983
Osteochondral Tissue......Page 985
Tendon and Ligament......Page 986
Tendon–Bone Interface: Enthesis......Page 987
Meniscus......Page 988
Intervertebral Disc......Page 989
Skeletal Muscle......Page 990
CHALLENGES AND PROSPECTS......Page 991
References......Page 993
SATELLITE CELL–DERIVED MYOBLASTS MEET THE PROPERTIES NEEDED FOR TRANSPLANTATION IN SKELETAL MUSCLES......Page 998
Gene Complementation......Page 999
Formation of New Myofibers......Page 1000
Formation of Graft-Derived Satellite Cells......Page 1002
Technical Approaches for Intramuscular Transplantation......Page 1003
Potential Risks of the Cell Injection Procedure......Page 1004
Improving the Efficiency of Cell Injections......Page 1006
Initial Survival......Page 1007
Long-term Survival......Page 1008
CONCLUSIONS......Page 1009
References......Page 1010
Background......Page 1014
History of Islet Transplantation......Page 1015
The Edmonton Protocol......Page 1016
Patient Assessment and Selection......Page 1017
Islet Transplantation Procedure......Page 1018
Immunosuppressive Therapy and Complications......Page 1019
Living Donor Islet Transplantation......Page 1020
Stem Cell Transplantation......Page 1022
Optimal Transplantation Site......Page 1024
Improving Engraftment Posttransplant......Page 1025
Improved Immunomodulation: Toward Donor-Specific Tolerance......Page 1026
References......Page 1028
FETAL DEVELOPMENT AND REGENERATIVE MEDICINE......Page 1036
PRECLINICAL ANIMAL STUDIES OF IN UTERO STEM CELL TRANSPLANTATION......Page 1038
Barriers to In Utero Stem Cell Transplantation Success......Page 1040
The Need for Better Hemophilia A Treatments......Page 1042
Feasibility and Justification for Treating Hemophilia A Before Birth......Page 1043
Genomic Integration-Associated Insertional Mutagenesis......Page 1045
Genome Editing......Page 1046
CLINICAL EXPERIENCE WITH IN UTERO STEM CELL TRANSPLANTATION......Page 1047
References......Page 1048
INTRODUCTION......Page 1056
Xenogenic Matrices......Page 1057
Tissue Engineering by Self-assembly......Page 1059
Extracellular Matrices Formed by Cell Culture and Synthetic Polymers......Page 1060
Nature-Derived Polymers and Synthetic Polymers......Page 1061
Synthetic Polymers With Seeded Cells......Page 1062
CONCLUSION......Page 1065
References......Page 1066
Young Populations......Page 1068
TISSUE ENGINEERED HEART VALVES......Page 1069
Considerations for Cell Source......Page 1070
Implant Function......Page 1071
Testing Tissue Engineered Heart Valve Function......Page 1072
Decellularized Bioscaffolds......Page 1073
Fibrin......Page 1074
Gelatin......Page 1075
Polyvinyl Alcohol......Page 1076
Hydrolytically Degradable Polymers......Page 1077
Tissue Engineered Heart Valve Fabrication Techniques......Page 1078
Future Direction in Tissue Engineered Heart Valves......Page 1079
CONCLUSIONS......Page 1080
References......Page 1081
LUNG DEVELOPMENT: A ROAD MAP TO REGENERATION......Page 1086
REPAIR AND REGENERATION IN THE NATIVE LUNG......Page 1087
NOVEL CELL POPULATIONS FOR LUNG REPAIR......Page 1088
BIOLOGICAL SCAFFOLDS TO SUPPORT REGENERATION......Page 1090
ADVANCES IN REBUILDING FUNCTIONAL LUNG TISSUE......Page 1092
List of Acronyms and Abbreviations......Page 1095
References......Page 1096
INTRODUCTION: FROM TISSUES TO ORGANS: KEY GOALS AND ISSUES......Page 1100
ENGINEERING OF CARDIAC PATCHES USING CELLS, SCAFFOLDS, AND BIOREACTORS......Page 1101
BIOPRINTING......Page 1105
CARDIAC ORGANOIDS AND ORGAN-ON-A-CHIP ENGINEERING......Page 1107
Oxygen Supply......Page 1110
Mechanical Stimulation......Page 1111
Electrical Stimulation......Page 1112
Tissue Architecture and Electrical Conduction......Page 1113
Vascularization......Page 1114
Host Response and Biocompatibility......Page 1115
In Situ Cardiac Tissue Engineering via Injection of Cells in Hydrogels......Page 1117
Implantation of Cardiac Patches......Page 1118
References......Page 1121
Further Reading......Page 1126
INTRODUCTION......Page 1128
Collagens......Page 1130
Decellularized Extracellular Matrix......Page 1131
Cancer Research......Page 1133
Bioartificial Liver and Transplantation Research......Page 1134
Limitations of Current In Vitro Liver Models to Test Drugs......Page 1135
Organoids in Drug Development......Page 1136
CONCLUSIONS AND FINAL PERSPECTIVES......Page 1137
References......Page 1138
Treatment Options, State of the Art, and Need for Corneal Regenerative Medicine......Page 1142
REGENERATIVE MEDICINE APPLIED TO KERATOPROSTHESIS DEVELOPMENT......Page 1143
Corneal Endothelium......Page 1145
FULLY CELL-BASED, SELF-ASSEMBLED CORNEAL CONSTRUCTS......Page 1146
Decellularized Extracellular Matrix as Implants......Page 1149
Peptide Analogs of Extracellular Matrix......Page 1150
CHALLENGES......Page 1152
References......Page 1153
ESOPHAGUS......Page 1158
STOMACH......Page 1161
SMALL INTESTINE......Page 1162
ANAL CANAL......Page 1168
IN VITRO MODELS......Page 1169
References......Page 1170
REQUIREMENTS OF A RENAL REPLACEMENT DEVICE......Page 1176
DEVICES USED IN CONVENTIONAL RENAL REPLACEMENT THERAPY......Page 1177
ADVANCEMENTS IN CONVENTIONAL RENAL REPLACEMENT THERAPY DEVICES......Page 1178
RENAL ASSIST DEVICE: A MORE COMPLETE RENAL REPLACEMENT THERAPY......Page 1179
RENAL ASSIST DEVICE THERAPY OF ACUTE KIDNEY INJURY CAUSED BY SEPSIS......Page 1180
IMMUNOMODULATORY EFFECT OF THE RENAL ASSIST DEVICE......Page 1181
SELECTIVE CYTOPHERETIC DEVICE......Page 1182
CHALLENGE: COST-EFFECTIVE STORAGE AND DISTRIBUTION FOR CELL DEVICES, BIOARTIFICIAL RENAL EPITHELIAL CELL SYSTEM DESIGN......Page 1183
BIOARTIFICIAL RENAL EPITHELIAL CELL SYSTEM AS AN EXTRACORPOREAL THERAPY TO TREAT ACUTE KIDNEY INJURY......Page 1184
WEARABLE BIOARTIFICIAL KIDNEY IN PRECLINICAL END-STAGE RENAL DISEASE MODEL......Page 1185
FUTURE ADVANCEMENTS FOR WEARABLE AND AMBULATORY RENAL REPLACEMENT THERAPIES......Page 1186
List of Acronyms and Abbreviations......Page 1187
References......Page 1188
INTRODUCTION......Page 1192
Cell Sources: Kidney Tissue–Derived Stem and Primary Cells......Page 1193
Other Cell Sources: Pluripotent, Fetal, or Adult Stem Cells......Page 1195
Engineering of Cell-Based Renal Constructs......Page 1196
CELL-FREE APPROACH: IN SITU RENAL REGENERATION......Page 1199
References......Page 1200
INTRODUCTION......Page 1206
Biomechanics......Page 1207
Uniaxial Tensile Testing......Page 1208
Contribution to Joint Function......Page 1209
HEALING OF LIGAMENTS AND TENDONS......Page 1210
Anterior Cruciate Ligament of the Knee......Page 1211
In Vitro Studies......Page 1212
Gene Therapy......Page 1213
Cell Therapy......Page 1214
Medial Collateral Ligament and Patellar Tendon Healing With Extracellular Matrix (Small Intestinal Submucosa)......Page 1215
Anterior Cruciate Ligament Healing......Page 1216
Mechanical Augmentation......Page 1217
Combined Biological and Mechanical Augmentation......Page 1218
SUMMARY AND FUTURE DIRECTIONS......Page 1219
References......Page 1220
INTRODUCTION......Page 1226
Endogenous Stem Cells......Page 1227
Biomolecule Delivery......Page 1228
Cell Therapy......Page 1229
Factors for Endogenous Stem Cell Stimulation......Page 1230
Cell Transplantation......Page 1232
Biomolecule Delivery......Page 1233
Guiding Axon Regrowth......Page 1234
Retinal Degeneration......Page 1235
Biomolecule Delivery......Page 1237
Cell Transplantation......Page 1238
List of Acronyms and Abbreviations......Page 1240
References......Page 1241
HISTORICAL BACKGROUND......Page 1250
Natural Scaffolds......Page 1251
Synthetic Scaffolds for Nerve Repair......Page 1252
Extracellular Matrix Molecules for Nerve Regeneration......Page 1253
Neurotrophic Factors and Cytokine Delivery for Nerve Regeneration......Page 1254
Seeding Neuronal Support Cells for Nerve Regeneration......Page 1255
ANISOTROPIC SCAFFOLDS FOR NERVE REGENERATION......Page 1256
Neurotrophic Factors......Page 1257
NATURAL NERVE GRAFTS......Page 1258
CONCLUSION......Page 1259
References......Page 1260
PRINCIPLES OF TISSUE ENGINEERING......Page 1264
THE VAGINA......Page 1265
Engineering of Functional Vaginal Tissue......Page 1266
Uterine Tissue Regeneration......Page 1267
THE OVARIES......Page 1269
Tissue Engineered Ovarian Follicles......Page 1270
Regenerating Ovarian Tissue From Stem Cells......Page 1271
Pelvic Organ Prolapse......Page 1272
References......Page 1273
TESTES......Page 1278
Spermatogonial Stem Cell Technology......Page 1279
Androgen Replacement Therapy......Page 1281
Spinal Ejaculation Generator......Page 1282
Penile Transplantation......Page 1284
References......Page 1285
INTRODUCTION......Page 1290
Stem Cell Sources......Page 1291
Multipotentiality......Page 1293
Paracrine Effects and Immunomodulatory Properties......Page 1294
Synthetic Scaffolds......Page 1295
Biodegradable Properties......Page 1296
Collagen......Page 1297
Fibrotic Bladder Model......Page 1298
Clinical Translation......Page 1300
Clinical Studies......Page 1301
References......Page 1302
INTRODUCTION......Page 1308
DEVELOPMENT, ANATOMY, AND FUNCTION OF SKIN......Page 1310
POTENTIAL PREREQUISITE REQUIREMENTS FOR TISSUE ENGINEERED SKIN SOLUTIONS......Page 1312
CURRENT TISSUE ENGINEERING SKIN TECHNOLOGIES......Page 1314
TISSUE ENGINEERING SKIN SOLUTIONS IN CLINICAL PRACTICE......Page 1316
THE FUTURE......Page 1317
References......Page 1319
INTRODUCTION......Page 1324
USE OF AUTOLOGOUS GROWTH FACTORS IN HAIR FOLLICLE REGENERATION......Page 1325
USE OF ADIPOSE-DERIVED STEM CELLS AND THEIR CONDITIONED MEDIUM FOR HAIR GROWTH......Page 1326
SIMULATING THE EMBRYONIC ENVIRONMENT......Page 1327
BIOENGINEERING A HUMAN HAIR FOLLICLE......Page 1331
References......Page 1333
Further Reading......Page 1335
Fetal Cells......Page 1336
Embryos......Page 1337
History of US Stem Cell Law and Policy......Page 1338
State Policy and Private Funding......Page 1340
The National Academies of Science......Page 1343
INTERNATIONAL COMPARISONS......Page 1345
SELECTED ETHICAL, LEGAL, SOCIAL, AND POLICY QUESTIONS OF STEM CELL RESEARCH......Page 1348
Compensating Egg Donors......Page 1350
Commercialization and Access to Treatments......Page 1351
Animal–Human Chimeras......Page 1352
List of Acronyms and Abbreviations......Page 1353
References......Page 1354
IS IT NECESSARY TO USE HUMAN EMBRYOS?......Page 1358
IS IT MORALLY PERMISSIBLE TO DESTROY A HUMAN EMBRYO?......Page 1359
MAY ONE BENEFIT FROM OTHERS' DESTRUCTION OF EMBRYOS?......Page 1360
MAY WE CLONE HUMAN EMBRYOS?......Page 1361
MAY WE USE HUMAN STEM CELLS TO CREATE CHIMERAS?......Page 1362
MAY WE GENETICALLY MODIFY HUMAN EMBRYOS?......Page 1363
Donor and Procurement Issues......Page 1364
Clinical Translation......Page 1365
References......Page 1367
BRIEF LEGISLATIVE HISTORY OF UNITED STATES FOOD AND DRUG ADMINISTRATION......Page 1372
LAWS, REGULATIONS, AND GUIDANCE......Page 1373
FOOD AND DRUG ADMINISTRATION ORGANIZATION AND JURISDICTIONAL ISSUES......Page 1374
APPROVAL MECHANISMS AND CLINICAL STUDIES......Page 1375
Regulation of Human Cells and Tissues Intended for Transplantation......Page 1377
Human Cellular Therapies......Page 1378
Xenotransplantation......Page 1380
Gene Therapy......Page 1381
Cell–Scaffold Combination Products......Page 1382
CLINICAL DEVELOPMENT PLAN......Page 1383
Food and Drug Administration's Standards Development Program......Page 1384
ADVISORY COMMITTEE MEETINGS......Page 1385
FOOD AND DRUG ADMINISTRATION RESEARCH AND CRITICAL PATH SCIENCE......Page 1386
OTHER COORDINATION EFFORTS......Page 1388
References......Page 1389
Primary Challenges for Widespread Adoption......Page 1394
Logistics......Page 1395
Scale-Up and Automation......Page 1396
ENVISIONED REGENERATIVE MEDICINE MANUFACTURING SYSTEMS OF THE FUTURE......Page 1397
Technical Societies......Page 1400
International Efforts......Page 1401
References......Page 1402
A......Page 1404
B......Page 1407
C......Page 1410
D......Page 1416
E......Page 1418
F......Page 1420
G......Page 1422
H......Page 1423
I......Page 1427
L......Page 1430
M......Page 1431
N......Page 1435
P......Page 1438
R......Page 1443
S......Page 1445
T......Page 1450
U......Page 1453
X......Page 1454
Z......Page 1455
Back Cover......Page 1456

Citation preview

PRINCIPLES OF REGENERATIVE MEDICINE

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PRINCIPLES OF REGENERATIVE MEDICINE THIRD EDITION Edited by

ANTHONY ATALA ROBERT LANZA ANTONIOS G. MIKOS ROBERT NEREM

Academic Press is an imprint of Elsevier 125 London Wall, London EC2Y 5AS, United Kingdom 525 B Street, Suite 1650, San Diego, CA 92101, United States 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom Copyright © 2019 Elsevier Inc. 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. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-12-809880-6 For information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: John Fedor Acquisition Editor: Mica Haley Editorial Project Manager: Timothy Bennett Production Project Manager: Punithavathy Govindaradjane Cover Designer: Miles Hitchen Typeset by TNQ Technologies

This book is dedicated to my family: Katherine, Christopher, and Zachary Anthony Atala To my family: Mary, Georgios, and Lydia Antonios G. Mikos To my wife, Marilyn, and to my four children, Nancy Black, Christy Maser, Steve Nerem, and Carol Wilcox Robert Nerem

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Contents Contributors Preface

4. The Molecular Circuitry Underlying Pluripotency in Embryonic and Induced Pluripotent Stem Cells

xix xxv

1. Molecular Organization of Cells

RACHEL H. KLEIN, PAUL S. KNOEPFLER

JON D. AHLSTROM

Introduction Ground State and Primed Embryonic Stem Cells Have Unique Signaling Networks Underlying Pluripotency Induced Pluripotent Stem Cells Leukemia Inhibitory Factor and Bone Morphogenic Protein Signaling Pathways Regulate Mouse Embryonic Stem Cell Self-Renewal Transforming Growth Factor b and Fibroblast Growth Factor Signaling Pathways Regulate Human Embryonic Stem Cell Self-Renewal Wnt Signaling Contributes to Maintenance of Pluripotency in Mouse Embryonic Stem Cells and to the Naive Human Embryonic Stem Cell State Three Transcription Factors, Octamer Binding Protein 4, SRY-Box 2, and Nanog, Form the Core Pluripotency Transcriptional Network MYC Links Cell Signaling to Pluripotency Gene Regulation A Specific Epigenetic Program Helps Maintain Pluripotency MicroRNAs Integrate With Cell Signaling and Transcription Factors to Regulate Stem Cell Proliferation and Differentiation Chromatin Structure Determines Regulatory Activity of Transcription Factor Binding to Pluripotency Genes Conclusions List of Acronyms and Abbreviations Acknowledgment References

Introduction Molecules That Organize Cells The EpithelialeMesenchymal Transition Transcriptional Program Molecular Control of the EpithelialeMesenchymal Transition Conclusion List of Acronyms and Abbreviations Glossary References

1 1 4 5 8 9 9 9

2. CelleExtracellular Matrix Interactions in Repair and Regeneration MELISSA PETREACA, MANUELA MARTINS-GREEN

Introduction Composition and Diversity of the Extracellular Matrix Receptors for Extracellular Matrix Molecules Signal Transduction Events During Celle Extracellular Matrix Interactions CelleExtracellular Matrix Interactions During Healing of Cutaneous Wounds CelleExtracellular Matrix Interactions During Regenerative Fetal Wound Healing Implications for Regenerative Medicine Acknowledgments References

15 15 16 18 25 28 30 31 31

3. Mechanisms of Blastema Formation and Growth in Regenerating Urodele Limbs

49 50

50

52

52

53 54 55

57

58 58 59 59 59

5. Scarless Wound Healing: From Experimental Target to Clinical Reality

DAVID L. STOCUM

Introduction Blastema Formation Blastema Growth List of Acronyms and Abbreviations Glossary Acknowledgments References

49

37 37 42 45 45 45 45

ALESSANDRA L. MOORE, CLEMENT D. MARSHALL, ALLISON NAUTA, HERMANN P. LORENZ, MICHAEL T. LONGAKER

Introduction Adult Skin Fetal Skin

vii

65 66 72

viii Regenerative Healing and Scar Reduction Theory Current Therapeutic Interventions Future Therapeutic Interventions Perspective List of Abbreviations References

CONTENTS

73 77 80 85 86 86

6. Progenitor and Stem Cell Heterogeneity: Using Big Data to Divide and Conquer MELANIE RODRIGUES, PAUL A. MITTERMILLER, JAGANNATH PADMANABHAN, GEOFFREY C. GURTNER

Introduction Single-Cell Isolation Acquiring Single-Cell Data Analyzing Single-Cell Data Determining Subpopulations Clinical Implications of Cellular Heterogeneity in Tissue Repair and Disease Conclusions References

93 95 96 100 103 105 108 109

7. Embryonic Stem Cells: Derivation, Properties, and Challenges IRINA KLIMANSKAYA

Introduction Derivation of Embryonic Stem Cells Sources of Human Embryonic Stem Cells Human Embryonic Stem Cell Maintenance Naive Embryonic Stem Cells Human Embryonic Stem Cell Differentiation and Manufacturing for Clinical Application Conclusions References Further Reading

113 113 114 116 119 119 120 120 123

8. Alternative Sources of Human Embryonic Stem Cells

10. Cord Blood Stem Cells KRISTIN M. PAGE, JESSICA M. SUN, JOANNE KURTZBERG

Introduction A Brief History Cord Blood Banking Public Versus Family (or Private) Banks Public Cord Blood Banking Procedures Clinical Uses of Umbilical Cord Blood Cord Blood Therapies for Inherited and Acquired Brain Diseases Investigations in the Treatment of Acquired Brain Injuries With Umbilical Cord Blood References

149 149 150 150 151 156 157 159 162

11. Induced Pluripotent Stem Cells ANDRES M. BRATT-LEAL, AI ZHANG, YANLING WANG, JEANNE F. LORING

Introduction Mechanisms of Reprogramming Epigenetic Remodeling Reprogramming Techniques Induced Transdifferentiation Genomic Stability Applications of Induced Pluripotent Stem Cells Disease Modeling Challenges and Future Possibilities in Disease Modeling Personalized Medicine Cell Therapy Conservation of Endangered Species Conclusions and Future Directions List of Acronyms and Abbreviations References

169 169 170 170 171 172 172 172 174 174 175 176 176 177 177

RANGARAJAN SAMBATHKUMAR, MANOJ KUMAR, CATHERINE M. VERFAILLIE

125 127 130 130

9. Stem Cells From the Amnion PAOLO DE COPPI, ANTHONY ATALA

Introduction Placenta: Function, Origin, and Composition Amniotic Fluid: Function, Origin, and Composition Amniotic Epithelial Cells Amniotic Mesenchymal Stem Cells Amniotic Fluid Stem Cells

144 144

12. Multipotent Adult Progenitor Cells

SVETLANA GAVRILOV, VIRGINIA E. PAPAIOANNOU, DONALD W. LANDRY

Introduction Organismically Dead Embryos Conclusion References

Conclusions References

133 133 134 134 135 138

Stem Cells Adult Stem Cells Isolation of Rodent Multipotent Adult Progenitor Cell Isolation of Human Multipotent Adult Progenitor Cells Differentiation Potential of Rodent and Human Multipotent Adult Progenitor Cells In Vitro Regenerative Capacities of Multipotent Adult Progenitor Cells Conclusion and Future Directions Conflict of Interest Statement References

181 181 182 183 183 183 187 187 187

CONTENTS

13. Hematopoietic Stem Cell Properties, Markers, and Therapeutics

17. Cardiac Stem Cells: Biology and Therapeutic Applications

JOHN D. JACKSON

KONSTANTINOS E. HATZISTERGOS, SARAH SELEM, WAYNE BALKAN, JOSHUA M. HARE

Introduction Hematopoietic Stem Cell Properties Hematopoietic Stem Cell Therapies Conclusion List of Acronyms and Abbreviations References

191 191 195 199 199 200

14. Mesenchymal Stem Cells ZULMA GAZIT, GADI PELLED, DMITRIY SHEYN, DORON C. YAKUBOVICH, DAN GAZIT

Introductory Overview Definition of Mesenchymal Stem Cells The Stem Cell Nature of Mesenchymal Stem Cells Which Tissues Contain Mesenchymal Stem Cells? Mesenchymal Stem Cell Isolation Techniques Mesenchymal Stem Cell Exosomes Immunomodulatory Effects of Mesenchymal Stem Cells Induced Pluripotent Stem CelleDerived Mesenchymal Stem Cells List of Acronyms and Abbreviations Acknowledgments References

205 206 206 207 208 208 211 212 213 213 214

15. Mesenchymal Stem Cells in Regenerative Medicine ARNOLD I. CAPLAN

Introduction and History New Insight All Mesenchymal Stem Cells Are Not Created Equal Clinical Trials The New Mesenchymal Stem Cells References

219 220 222 224 224 225

16. Cell Therapy of Liver Disease: From Hepatocytes to Stem Cells

247 248 252 254 255 265 265 267 267

18. Skeletal Muscle Stem Cells NORA YUCEL, HELEN M. BLAU

Introduction Satellite Cells Are Muscle Stem Cells The Molecular Characteristics of Muscle Stem Cells During Myogenesis in Regeneration Functional Characteristics of Muscle Stem Cells Isolation of Muscle Stem Cells Tracking Muscle Stem Cell Behavior Through Live Imaging (Bioluminescence Imaging and Intravital Imaging) Regulation of Muscle Stem Cells by Their Niche Satellite Cell Self-Renewal Mechanisms Muscle Stem CelleIntrinsic Defects in Aging and Disease Challenges in the Use of Satellite Cells in Regenerative Medicine Gene Editing Strategies Other Stem Cell Types Within Muscle Conclusions References

273 274 274 276 277

278 279 281 283 284 285 285 287 287

19. Stem Cells Derived From Fat JAMES C. BROWN, ADAM J. KATZ

STEPHEN C. STROM, CARL JORNS

Introduction Background Studies Clinical Hepatocyte Transplantation Hepatocyte Transplantation Novel Uses, Challenges, and Future Directions Summary References

Development of the Heart From Cardiac Stem/ Progenitor Cells Cardiac Stem/Progenitor Cells in the Adult Heart Cell-Based Therapeutics for Heart Disease Mechanisms of Action Clinical Trials Methods for Expansion of Adult Cardiac Stem Cells Combined Stem Cell Therapeutics Conclusions References

ix

229 231 232 237 241 242

Introduction Cellular Fractions Cellular Characterization Clinical Delivery of Adipose-Derived Cells Engineered Neo-Tissue Therapeutic Safety of Adipose-Derived Cells Conclusions References

295 295 296 297 300 301 302 302

x

CONTENTS

20. Peripheral Blood Stem Cells

List of Acronyms and Abbreviations Acknowledgments References

ABRITEE DAHL, GRAC ¸ A ALMEIDA-PORADA, CHRISTOPHER D. PORADA, SHAY SOKER

Introduction Types and Source of Stem Cells in the Peripheral Blood Endothelial Progenitor Cells Mesenchymal Stem/Marrow Stroma Cells Therapeutic Applications of Peripheral Blood Stem Cells Conclusions and Future Directions References

307 307 311 315 317 325 326

21. From Adult Pancreatic Islets to Stem Cells: Regenerative Strategies for the Treatment of Diabetes and Its Complications MARTA POKRYWCZYNSKA, GIACOMO LANZONI, CAMILLO RICORDI

Introduction From Adult Pancreatic Islets to Stem Cells Strategies for the Generation of b Cells for Replacement Therapy b Cells From Pluripotent Stem Cells (Embryonic Stem Cells and Induced Pluripotent Stem Cells) b Cells From Adult Stem/Progenitor Cells of the Biliary Tree and Pancreas Mesenchymal Stem Cells to Modulate Immunity and Promote Tissue Repair in Diabetes Conclusion References

335 335 336 336 341 344 345 346

22. Stem Cells for Diseases of the Retina AARON NAGIEL, STEVEN D. SCHWARTZ

Introduction Cell-Replacement Therapy Cell-Based Neuroprotection Disease-in-a-Dish Modeling for Retinal Disorders Conclusion References

351 355 361 362 364 365

23. Stem Cells for Traumatic Brain Injury CHRISTOPHER M. SCHNEIDER, MARGARET L. JACKSON, SUPINDER S. BEDI, CHARLES S. COX, JR.

Introduction Phases of Brain Injury Current Traumatic Brain Injury Management Strategies Preclinical Data Supporting Stem Cell Therapies for Traumatic Brain Injury Clinical Trials Conclusion

369 370 376 376 381 385

386 386 387

24. Mechanical Determinants of Tissue Development VOLHA LIAUDANSKAYA, DISHA SOOD, DAVID L. KAPLAN

Introduction Mechanotransduction Mechanisms and Major Effectors Nucleus as the Central Organelle in Regulating Mechanotransduction Cellular Mechanotransduction Mechanisms Conclusions Acknowledgments References

391 392 396 397 400 401 401

25. Morphogenesis, Bone Morphogenetic Proteins, and Regeneration of Bone and Articular Cartilage A.H. REDDI, KENJIRO IWASA

Introduction Bone Morphogenetic Proteins Stem Cells Scaffolds of Extracellular Matrix and Biomimetic Biomaterials Articular Cartilage Regeneration and Cartilage Morphogenetic Proteins Regenerative Medicine and Surgery of Articular Cartilage Regeneration of Articular Cartilage Surface and Lubrication List of Acronyms and Abbreviations Acknowledgments References

405 406 409 409 412 412 413 414 414 414

26. Physical Stress as a Factor in Tissue Growth and Remodeling JOEL D. BOERCKEL, CHRISTOPHER V. GEMMITI, DEVON E. MASON, YASH M. KOLAMBKAR, BLAISE D. PORTER, ROBERT E. GULDBERG

Introduction Describe the Mechanical Environment Understand the Role of Mechanical Stimuli Mechanical Regulation of Vascularized Tissue Regeneration Evaluate Functional Restoration Conclusions References

417 418 423 427 432 432 433

CONTENTS

27. CelleSubstrate Interactions

Manufacturability Summary References

MUHAMMAD RIZWAN, JOHN W. TSE, APARNA NORI, KAM W. LEONG, EVELYN K.F. YIM

Introduction CelleExtracellular Matrix Interactions CelleSubstrate Interactions Effect of Dimensionality Conclusion Acknowledgments References

437 437 438 451 460 461 461

28. Intelligent Surfaces for Cell Sheet Engineering HIRONOBU TAKAHASHI, TATSUYA SHIMIZU, TERUO OKANO

Introduction The Intelligence of Thermoresponsive Polymers for Cell Sheet Engineering Clinical Applications for Cell Sheet Engineering Cell Sheet Engineering Produces Scaffold-Free, Three-Dimensional Tissue Constructs Combination of Cell Sheet Engineering and Scaffold-Based Engineering Microfabricated Intelligent Surface for Engineering Complex Tissue Constructs Conclusions References

469 469 472 474 477 478 481 482

29. Applications of Nanotechnology for Regenerative Medicine; Healing Tissues at the Nanoscale YAFENG YANG, ADITYA CHAWLA, JIN ZHANG, ADAM ESA, HAE LIN JANG, ALI KHADEMHOSSEINI

Introduction Properties of Nanomaterials Nanobiomaterials Nanotechnology-Based Strategies in Regenerative Medicine Nanotechnology-Based Stem Cell Therapy Conclusion References

485 486 491 493 497 499 499

517 517 518

31. Proteins Controlled With Precision at Organic, Polymeric, and Biopolymer Interfaces for Tissue Engineering and Regenerative Medicine DAVID G. CASTNER, BUDDY D. RATNER

Why the Need for Precision Control of Proteins at Interfaces in Tissue Engineering and Regenerative Medicine? Surface Analysis and Its Role in the Precision Delivery of Biological Signals A Short Review of Key Surface Analysis Methods and Supporting Tools Techniques and Technologies for Precision Immobilization at Surfaces Conclusions References

523 524 525 527 531 532

32. Natural Origin Materials for Bone Tissue Engineering: Properties, Processing, and Performance F. RAQUEL MAIA, VITOR M. CORRELO, JOAQUIM M. OLIVEIRA, RUI L. REIS

Introduction Natural-Based Polymers Chitosan Collagen Gellan Gum Polyhydroxyalkanoates Silk Fibroin Starch Natural-Based Bioceramics Calcium Phosphates Silicate Ceramics Conclusions List of Acronyms and Abbreviations Acknowledgments References

535 538 538 540 543 545 547 548 550 551 553 554 555 555 555

33. Synthetic Polymers

30. Design Principles in Biomaterials and Scaffolds

MICHAEL C. HACKER, JAN KRIEGHOFF, ANTONIOS G. MIKOS

YANG ZHU, WILLIAM R. WAGNER

Function and Application-Oriented Design of Biomaterial Scaffolds Safety and Biocompatibility Requirements for Biomaterial Scaffolds

xi

505 513

Introduction Polymer Synthesis Nondegradable Synthetic Polymers Biodegradable Synthetic Polymers for Regenerative Medicine

559 560 561 567

xii

CONTENTS

Applications of Synthetic Polymers Conclusion/Summary References

580 580 581

34. Calcium Phosphate Bioceramics and Cements

591 594 595 601 606 607 608 608 608

35. Biologic Scaffolds Composed of Extracellular Matrix for Regenerative Medicine

613 613 619 622 622 622 623 623

DAVID F. WILLIAMS

627 628 631 632 639 646 648 649

37. Surface Modification of Biomaterials RACHIT AGARWAL, ANDRE´S J. GARCI´A

Introduction Physicochemical Surface Modifications Overcoating Technologies

The Need for Replacement Tissues Tissue Components Regeneration of Diseased Tissues Design Parameters for Histogenesis Synthetic Materials for Histogenesis of New Organs Future Directions in Three-Dimensional Scaffolds: Three-Dimensional Microfabrication Conclusions References

661 661 662 663 669 670 670 671

JAMES M. ANDERSON

36. Hydrogels in Regenerative Medicine Introduction Biomaterials Templates StructureeProperty Relationships in Hydrogels Increasing Sophistication of Synthetic Hydrogels for Tissue Engineering Natural Biopolymers as Extracellular MatrixeAnalog Hydrogels Synthetic Hydrogels for Tissue Engineering Templates Conclusions References

MELISSA K. MCHALE, NICOLE M. BERGMANN, JENNIFER L. WEST

39. Biocompatibility and Bioresponse to Biomaterials

MICHELLE SCARRITT, MARK MURDOCK, STEPHEN F. BADYLAK

Introduction Extracellular Matrix: Function and Components Intact and Solubilized Extracellular Matrix as a Scaffold Material Clinical and Commercial Applications Conclusions List of Acronyms and Abbreviations Glossary References

656 659 659 659

38. Histogenesis in Three-Dimensional Scaffolds

NATHAN W. KUCKO, RALF-PETER HERBER, SANDER C.G. LEEUWENBURGH, JOHN A. JANSEN

Introduction Classes of Calcium Phosphate Cements Physiochemical Properties Strategies to Improve the Mechanical Properties Clinical Applications Conclusion List of Acronyms and Abbreviations Glossary References

Biological Modification of Surfaces Surface Chemical Patterning Conclusion and Future Prospects References

651 653 655

Introduction Inflammation (Innate Immunity) and Wound Healing Fibrosis and Fibrous Encapsulation Immunotoxicity (Acquired Immunity) Conclusion References Further Reading

675 676 683 684 691 691 694

40. Hybrid Composite Biomaterials NIRMALYA TRIPATHY, ELUMALAI PERUMAL, RAFIQ AHMAD, JEONG EUN SONG, GILSON KHANG

Introduction Fundamentals of Bone Development and Defects Functions of Scaffolding and Extracellular Matrix Scaffolding Approaches in Bone Tissue Engineering Scaffolding Materials Conclusions and Future Prospects Acknowledgments References

695 696 696 697 699 710 710 710

41. Materials-Based Cancer Immunotherapies JARED M. NEWTON, ANDREW G. SIKORA, SIMON YOUNG

Introduction and Overview of Cancer Immunotherapy Advantages and Disadvantages of Cancer Immunotherapy Nanoparticle Biomaterials for Cancer Immunotherapy

715 717 718

CONTENTS

Macroscale Biomaterial Scaffolds for Cancer Immunotherapy Conclusion List of Acronyms and Abbreviations Glossary Acknowledgments References

727 733 734 734 736 736

YUNLAN FANG, XUGUANG CHEN, W.T. GODBEY

741 748 749 750 755 755

43. Preclinical Bone Repair Models in Regenerative Medicine 761 761 762 762 763 764 766 766

44. Body-on-a-Chip: Regenerative Medicine for Personalized Medicine ALEKSANDER SKARDAL, THOMAS SHUPE, ANTHONY ATALA

Introduction Advance of In Vitro Organoid Development: Progression From Two-Dimensional to Three-Dimensional Models Organ-on-a-Chip Technologies and Their Applications Body-on-a-Chip: Multiorgan Systems and Future Applications Conclusions and Perspectives References

769

Introduction Fundamentals of Three-Dimensional Printing Bioinks Conclusion and Future Directions List of Acronyms and Abbreviations References

805 805 808 826 827 827

47. Three-Dimensional Tissue and Organ Printing in Regenerative Medicine

Introduction Bioprinting Strategy: From Medical Image to Printed Tissue Bioprinting Mechanisms Biomaterials for Bioprinting: Bioinks Three-Dimensional Bioprinting in Regenerative Medicine Applications Conclusions and Future Perspectives List of Abbreviations Glossary References

831 831 832 835 838 847 848 848 849

48. Biomineralization and Bone Regeneration KUNAL J. RAMBHIA, PETER X. MA

770 772 775 783 783

45. Bioreactors in Regenerative Medicine

Development and Fracture of Bone Principles of Bone Tissue Engineering Stem Cells in Bone Tissue Engineering Scaffolds for Bone Tissue Engineering Growth and Differentiation Factors in Bone Tissue Engineering Immunomodulation in Bone Regeneration References

853 854 854 856 859 861 863

49. Hair Cell Regeneration in the Inner Ear and Lateral Line

JINHO KIM, KELSEY KENNEDY, GORDANA VUNJAK-NOVAKOVIC

Introduction Design Considerations for Creating Bioreactors Lung Bioreactors

46. Bioinks for Three-Dimensional Printing in Regenerative Medicine

GREGORY J. GILLISPIE, JIHOON PARK, JOSHUA S. COPUS, ANIL KUMAR PALLICKAVEEDU RAJAN ASARI, JAMES J. YOO, ANTHONY ATALA, SANG JIN LEE

ELVIS L. FRANCOIS, MICHAEL J. YASZEMSKI

Introduction Biomineralization and Bone Regeneration Cell Sources Embryonic Stem Cells Scaffolds Preclinical Models of Bone Tissue Regeneration Conclusions References

795 801 801 801

JAVIER NAVARRO, GISELE A. CALDERON, JORDAN S. MILLER, JOHN P. FISHER

42. Gene Editing in Regenerative Medicine Genome Editing Tools Delivery Cargo Delivery Methods Applications of Gene Editing in Regenerative Medicine Closing Remarks References

Bone Bioreactors Challenges and Future Directions Acknowledgments References

xiii

787 787 788

MATTHEW W. KELLEY, JASON R. MEYERS

Introduction Structure of the Inner Ear

867 867

xiv

CONTENTS

Hair Cell Loss History of Hair Cell Regeneration Spontaneous Hair Cell Regeneration in Mammalian Vestibular Organs Road Blocks to Regeneration Insights From Developmental Biology Induction of Hair Cell Regeneration Using Transgenic Mice Studies of Hair Cell Regeneration Using the Lateral Line Formation of New Neuromasts From Multipotent Progenitors Hair Cell Regeneration in the Lateral Line Pathways Coordinating Hair Cell Regeneration in the Lateral Line Open Questions About Lateral Line Regeneration Conclusions Clinical Trial References

868 869 870 871 871 875 876 877 878 879 880 881 881 881

50. Craniofacial Regenerative Medicine BRANDON T. SMITH, EMMA WATSON, ISSA A. HANNA, JAMES C. MELVILLE, ANTONIOS G. MIKOS, MARK E. WONG

Introduction Understanding the Craniofacial Regenerative Environment Current Methods of Maxillofacial Reconstruction Tissue Engineering Technologies Currently Used Conclusion List of Abbreviations Acknowledgments References

887

Cartilage and Cartilage Repair Tissue Engineering for Cartilage Repair Current and Future Trends in Cartilage Engineering References

946 947

BENJAMIN B. ROTHRAUFF, ALESSANDRO PIROSA, HANG LIN, JIHEE SOHN, MARK T. LANGHANS, ROCKY S. TUAN

Introduction Stem Cell Therapies for Musculoskeletal Diseases Challenges and Prospects List Abbreviations and Acronyms Acknowledgments References

953 954 964 966 966 966

55. Myoblast Transplantation in Skeletal Muscles DANIEL SKUK, JACQUES P. TREMBLAY

907 908 909 910 917 917 918 918

56. Islet Cell Transplantation

52. Cell Therapy for Blood Substitutes

937 938

54. Stem Cell Therapy for Musculoskeletal Diseases

Introduction Satellite CelleDerived Myoblasts Meet the Properties Needed for Transplantation in Skeletal Muscles Cell Administration Cell-Graft Survival Conclusions List of Abbreviations Glossary References

NELSON MONTEIRO, PAMELA C. YELICK

971

971 976 980 982 983 983 983

JULIET A. EMAMAULLEE, ANDREW PEPPER, A.M. JAMES SHAPIRO

Introduction Clinical Islet Transplantation Future Challenges Summary and Conclusions References

987 990 993 1001 1001

57. Prenatal Cell- and Gene-Based Therapies for Regenerative Medicine

SHI-JIANG LU, ROBERT LANZA

Introduction Red Blood Cells Megakaryocytes and Platelets Hematopoietic Stem Cells Perspectives Financial and Competing Interest Disclosure References Further Reading

HEATHER J. FAUST, QIONGYU GUO, JENNIFER H. ELISSEEFF

887 890 891 899 901 902 902

51. Dental Tissue Engineering Introduction Tooth Development Dental Stem Cells Dental Tissue Engineering Conclusions List of Abbreviations Acknowledgments References

53. Cartilage Tissue Engineering

923 924 929 931 933 933 933 936

GRAC ¸ A ALMEIDA-PORADA, CHRISTOPHER D. PORADA

Introduction Fetal Development and Regenerative Medicine Preclinical Animal Studies of In Utero Stem Cell Transplantation

1009 1009 1011

CONTENTS

Clinical Experience With In Utero Stem Cell Transplantation Conclusions and Future Directions References

1020 1021 1021

58. Engineering of Large Diameter Vessels HIDEKI MIYACHI, TOSHIHIRO SHOJI, SHINKA MIYAMOTO, TOSHIHARU SHINOKA

Introduction Materials for and Approaches to Fabricating Tissue Engineered Vascular Grafts Conclusion List of Acronyms and Abbreviations References

1029 1030 1038 1039 1039

59. Regenerative Medicine Approaches for Tissue Engineered Heart Valves 1041 1042 1042 1046 1053 1054

60. Regenerative Medicine of the Respiratory Tract SARAH E. GILPIN, PHILIPP T. MOSER, HARALD C. OTT

Lung Development: A Road Map to Regeneration Repair and Regeneration in the Native Lung Novel Cell Populations for Lung Repair Biological Scaffolds to Support Regeneration Advances in Rebuilding Functional Lung Tissue Clinical Translation and Future Considerations List of Acronyms and Abbreviations References

1059 1060 1061 1063 1065 1068 1068 1069

61. Cardiac Tissue SERENA MANDLA, MILICA RADISIC

Introduction: From Tissues to Organs: Key Goals and Issues Engineering of Cardiac Patches Using Cells, Scaffolds, and Bioreactors Bioprinting Cardiac Organoids and Organ-on-a-Chip Engineering Engineering the Entire Ventricle Bioreactors and Conditioning Tissue and Organ Function In Vivo Studies Summary References Further Reading

62. Bioengineering of Liver Tissue

PILAR SAINZ-ARNAL, IRIS PLA-PALACI´N, ´ NCHEZ-ROMERO, MANUEL ALMEIDA, NATALIA SA SARA MORINI, ESTELA SOLANAS, ALBERTO LUE, ´ , PEDRO M. BAPTISTA TRINIDAD SERRANO-AULLO

Introduction Hepatic Tissue Engineering Liver Spheroids, Organoids, and Aggregates: Cancer, Bioartificial Liver, Transplantation Research, and Toxicology and Drug Development Conclusions and Final Perspectives Acknowledgments References

1101 1103

1106 1110 1111 1111

63. Regenerative Medicine in the Cornea

¨ NS HILBORN, FIONA SIMPSON, EMILIO I. ALARCON, JO ISABELLE BRUNETTE, MAY GRIFFITH

JAMES K. WILLIAMS, JAMES J. YOO, ANTHONY ATALA

Introduction Clinical Options Tissue Engineered Heart Valves Biomaterials for Tissue Engineered Heart Valves Conclusions References

xv

1073 1074 1078 1080 1083 1083 1086 1090 1094 1094 1099

Introduction Regenerative Medicine Applied to Keratoprosthesis Development Regeneration of Corneal Layers Fully Cell-Based, Self-Assembled Corneal Constructs Cell-Free Biomaterials CelleBiomaterial Composites Composite Implants Incorporating Specific Bioactive Functions Challenges Conclusions and Future Perspective References

1115 1116 1118 1119 1122 1125 1125 1125 1126 1126

64. Alimentary Tract RICHARD M. DAY

Introduction Esophagus Stomach Small Intestine Colon Anal Canal In Vitro Models Conclusion References

1131 1131 1134 1135 1141 1141 1142 1143 1143

65. Extracorporeal Renal Replacement CHRISTOPHER J. PINO, H. DAVID HUMES

Introduction Requirements of a Renal Replacement Device Devices Used in Conventional Renal Replacement Therapy Advancements in Conventional Renal Replacement Therapy Devices Renal Assist Device: A More Complete Renal Replacement Therapy

1149 1149 1150 1151 1152

xvi Renal Assist Device Therapy of Acute Kidney Injury Caused by Sepsis Clinical Experience With a Renal Assist Device to Treat Acute Kidney Injury Immunomodulatory Effect of the Renal Assist Device Selective Cytopheretic Device Challenge of Cell-Based Device: Robust Cell Source Challenge: Cost-Effective Storage and Distribution for Cell Devices, Bioartificial Renal Epithelial Cell System Design Bioartificial Renal Epithelial Cell System as an Extracorporeal Therapy to Treat Acute Kidney Injury Wearable Bioartificial Kidney in Preclinical End-Stage Renal Disease Model Complete Bioartificial Kidney System for Use in End-Stage Renal Disease Future Advancements for Wearable and Ambulatory Renal Replacement Therapies Conclusion List of Acronyms and Abbreviations Glossary References

CONTENTS

1153 1154 1154 1155 1156

1156

1157 1158 1159 1159 1160 1160 1161 1161

66. Regenerative Medicine Approaches for the Kidney 1165 1166 1172 1173 1173 1173

67. Functional Tissue Engineering of Ligament and Tendon Injuries

MAHESH C. DODLA, MELISSA ALVARADO-VELEZ, VIVEK J. MUKHATYAR, RAVI V. BELLAMKONDA

Problems and Challenges With Peripheral Nerve Injuries Historical Background Current Strategies for Peripheral Nerve Regeneration Isotropic Scaffolds for Nerve Regeneration Anisotropic Scaffolds for Nerve Regeneration Natural Nerve Grafts Animal Models Conclusion References

1223 1223 1224 1224 1229 1231 1232 1232 1233

70. Regenerative Medicine for the Female Reproductive System Introduction Principles of Tissue Engineering The Vagina The Uterus The Ovaries Other Tissue Engineering Applications in the Female Reproductive System Conclusions and Future Trends References

1237 1237 1238 1240 1242 1245 1246 1246

HOOMAN SADRI-ARDEKANI, ANTHONY ATALA

1179 1180 1183 1185 1188 1192 1193

Introduction Testes Ejaculatory System Penis Conclusion Acknowledgment References

1251 1251 1255 1257 1258 1258 1258

72. Regenerative Medicine of the Bladder

68. Central Nervous System

YUANYUAN ZHANG, ANTHONY ATALA

SAMANTHA L. PAYNE, BRIAN G. BALLIOS, M. DOUGLAS BAUMANN, MICHAEL J. COOKE, MOLLY S. SHOICHET

Introduction Wound Response and Barriers to Regeneration Therapeutic Strategies in the Central Nervous System Case Studies in Tissue Therapy in the Central Nervous System

69. Peripheral Nerve Regeneration

71. Regenerative Medicine for the Male Reproductive System

SAVIO L-Y. WOO, JONQUIL R. MAU, HUIJUN KANG, RUI LIANG, ALEJANDRO J. ALMARZA, MATTHEW B. FISHER

Introduction Normal Ligaments and Tendons Healing of Ligaments and Tendons Application of Functional Tissue Engineering Healing of Ligaments and Tendons Summary and Future Directions References

1213 1213 1214 1214

RENATA S. MAGALHAES, ANTHONY ATALA

IN KAP KO, JAMES J. YOO, ANTHONY ATALA

Introduction Cell-Based Therapy Cell-Free Approach: In Situ Renal Regeneration Conclusions and Future Perspectives Acknowledgments References

Conclusions and Outlook List of Acronyms and Abbreviations Acknowledgments References

1199 1200 1201 1203

Introduction Cell Sources Biodegradable Biomaterials Preclinical Models Clinical Trials Conclusion References

1263 1264 1268 1271 1273 1275 1275

CONTENTS

73. Therapeutic Applications: Tissue Engineering of Skin FIONA M. WOOD

Introduction Development, Anatomy, and Function of Skin Potential Prerequisite Requirements for Tissue Engineered Skin Solutions Current Tissue Engineering Skin Technologies Tissue Engineering Skin Solutions in Clinical Practice The Future Conclusion References

1281 1283 1285 1287 1289 1290 1292 1292

74. Regenerative Medicine Approaches for Engineering a Human Hair Follicle GAIL K. NAUGHTON

Introduction Use of Autologous Growth Factors in Hair Follicle Regeneration Use of Adipose-Derived Stem Cells and Their Conditioned Medium for Hair Growth Tissue-Derived Materials for Hair Regeneration Additional Studies on Secreted Growth Factors and Hair Growth Simulating the Embryonic Environment Bioengineering a Human Hair Follicle Summary References Further Reading

1297 1298 1299 1300 1300 1300 1304 1306 1306 1308

75. US Stem Cell Research Policy JOSEPHINE JOHNSTON, RACHEL L. ZACHARIAS

Introduction Sources of Stem Cells United States Federal and State Stem Cell Policy Stem Cell Research Guidelines International Comparisons Selected Ethical, Legal, Social, and Policy Questions of Stem Cell Research Conclusion List of Acronyms and Abbreviations Acknowledgments References

Introduction and Chapter Overview Brief Legislative History of United States Food and Drug Administration Laws, Regulations, and Guidance Food and Drug Administration Organization and Jurisdictional Issues Approval Mechanisms and Clinical Studies Meetings With Industry, Professional Groups, and Sponsors Regulations and Guidance of Special Interest for Regenerative Medicine Preclinical Development Plan Clinical Development Plan Advisory Committee Meetings Food and Drug Administration Research and Critical Path Science Other Coordination Efforts Conclusion References

1321 1326 1326 1327 1327

Why Regenerative Medicine Manufacturing? Current Challenges and Opportunities in Regenerative Medicine Manufacturing Envisioned Regenerative Medicine Manufacturing Systems of the Future Global Landscape for Regenerative Medicine Manufacturing Summary and Conclusions References

1331 1331 1332 1333

1337 1340 1340

CAROLYN YONG, DAVID S. KAPLAN, ANDREA GRAY, LAURA RICLES, ANNA KWILAS, SCOTT BRUBAKER, JUDITH ARCIDIACONO, LEI XU, CYNTHIA CHANG, REBECCA ROBINSON, RICHARD MCFARLAND

78. Regenerative Medicine ManufacturingdChallenges and Opportunities

RONALD M. GREEN

1334 1334 1335 1336

77. Overview of the US Food and Drug Administration Regulatory Process

1309 1309 1311 1316 1318

76. Ethical Considerations Introduction Is It Necessary to Use Human Embryos? Is It Morally Permissible to Destroy a Human Embryo? May One Benefit From Others’ Destruction of Embryos?

May We Create an Embryo to Destroy It? May We Clone Human Embryos? May We Use Human Stem Cells to Create Chimeras? May We Genetically Modify Human Embryos? Are There Special Considerations Governing the Use of Stem Cells in Clinical Research and Clinical Applications? Conclusion References

xvii

1345 1345 1346 1347 1348 1350 1350 1356 1356 1358 1359 1361 1362 1362

PAUL COHEN, JOSHUA G. HUNSBERGER, ANTHONY ATALA

Index

1367 1367 1370 1373 1375 1375

1377

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Contributors Rachit Agarwal Centre for BioSystems Science and Engineering, Indian Institute of Science, Bangalore, India

Joel D. Boerckel University of Pennsylvania, Philadelphia, PA, United States

Jon D. Ahlstrom States

Andres M. Bratt-Leal The Scripps Research Institute, San Diego, CA, United States; Summit for Stem Cell Foundation, San Diego, CA, United States

PolarityTE, Salt Lake City, UT, United

Rafiq Ahmad Chonbuk National University, Jeonju-si, Republic of Korea

James C. Brown United States

Emilio I. Alarcon University of Ottawa Heart Institute, Ottawa, ON, Canada

Scott Brubaker Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States

Alejandro J. Almarza University of Pittsburgh, Pittsburgh, PA, United States

Isabelle Brunette Maisonneuve-Rosemont Hospital Research Centre, Montreal, QC, Canada; University of Montreal, Montreal, QC, Canada

Grac¸a Almeida-Porada Wake Forest Institute for Regenerative Medicine, Wake Forest University, WinstonSalem, NC, United States

Gisele A. Calderon States

Manuel Almeida Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain Melissa Alvarado-Velez United States

Duke University, Durham, NC,

David G. Castner United States

Aditya Chawla Harvard Medical School, Brigham and Women’s Hospital, Boston, MA, United States; Massachusetts Institute of Technology, Cambridge, MA, United States; Harvard University, Boston, MA, United States

Anthony Atala Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Stephen F. Badylak University of Pittsburgh, Pittsburgh, PA, United States

Xuguang Chen Salubris Biotherapeutics, Inc., Gaithersburg, MD, United States

University of Miami, Miami, FL, United University of Toronto, Toronto, ON, Canada

Pedro M. Baptista Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain; Center for Biomedical Research Network Liver and Digestive Diseases (CIBERehd), Zaragoza, Spain; Instituto de Investigacio´n Sanitaria de la Fundacio´n Jime´nez Dı´az, Madrid, Spain; Universidad Carlos III de Madrid, Madrid, Spain M. Douglas Baumann Canada

Paul Cohen North Carolina State University, Raleigh, NC, United States Michael J. Cooke Canada

University of Toronto, Toronto, ON,

Joshua S. Copus Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

Syngenta Canada Inc., Guelph, ON,

Supinder S. Bedi McGovern Medical School at the University of Texas Health Science Center at Houston, Houston, TX, United States Ravi V. Bellamkonda States

Duke University, Durham, NC, United

Nicole M. Bergmann States

Rice University, Houston, TX, United

Helen M. Blau Stanford University School of Medicine, Stanford, CA, United States

University of Washington, Seattle, WA,

Cynthia Chang Center for Devices and Radiological Health, FDA, Silver Spring, MD, United States

Judith Arcidiacono Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States

Brian G. Ballios

Rice University, Houston, TX, United

Arnold I. Caplan Case Western Reserve University, Cleveland, OH, United States

James M. Anderson Case Western Reserve University, Cleveland, OH, United States

Wayne Balkan States

University of Florida, Gainesville, FL,

Vitor M. Correlo 3B’s Research Group, I3Bs e Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Guimara˜es, Portugal; ICVS/3B’sePT Government Associate Laboratory, Braga/Guimara˜es, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimara˜es, Portugal Charles S. Cox, Jr. McGovern Medical School at the University of Texas Health Science Center at Houston, Houston, TX, United States Abritee Dahl Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

xix

xx

CONTRIBUTORS

Richard M. Day University College London, London, United Kingdom

Robert E. Guldberg Georgia Institute of Technology, Atlanta, GA, United States

Paolo De Coppi UCL Institute of Child Health and Great Ormond Street Hospital, London, United Kingdom; Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

Qiongyu Guo States

Mahesh C. Dodla Duke University, Durham, NC, United States Jennifer H. Elisseeff Johns Hopkins University, Baltimore, MD, United States Juliet A. Emamaullee Department of Surgery, University of Alberta, Edmonton, AB, Canada Adam Esa Cardiff University, Cardiff Wales, United Kingdom Yunlan Fang XenoBiotic Laboratories, Inc., Plainsboro Township, NJ, United States Heather J. Faust United States

Johns Hopkins University, Baltimore, MD,

John P. Fisher University of Maryland, College Park, MD, United States; Center for Engineering Complex Tissues, College Park, MD, United States Matthew B. Fisher North Carolina State University, Raleigh, NC, United States Elvis L. Francois Department of Orthopedic Surgery, Mayo Clinic, Rochester, MN, United States Andre´s J. Garcı´a Woodruff School of Mechanical Engineering and Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States Svetlana Gavrilov College of Physicians and Surgeons of Columbia University, New York, NY, United States Dan Gazit Cedars-Sinai Medical Center, Los Angeles, CA, United States; Hebrew University of Jerusalem, Jerusalem, Israel Zulma Gazit Cedars-Sinai Medical Center, Los Angeles, CA, United States; Hebrew University of Jerusalem, Jerusalem, Israel Christopher V. Gemmiti Georgia Institute of Technology, Atlanta, GA, United States Gregory J. Gillispie Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Sarah E. Gilpin Massachusetts General Hospital, Boston, MA, United States; Harvard Medical School, Boston, MA, United States W.T. Godbey States

Tulane University, New Orleans, LA, United

Andrea Gray Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States Ronald M. Green Dartmouth College, Hanover, NH, United States May Griffith Maisonneuve-Rosemont Hospital Research Centre, Montreal, QC, Canada; University of Montreal, Montreal, QC, Canada

Lehigh University, Bethlehem, PA, United

Geoffrey C. Gurtner United States Michael C. Hacker

Stanford University, Stanford, CA, University of Leipzig, Leipzig, Germany

Issa A. Hanna University of Texas Health Science Center at Houston, Houston, TX, United States Joshua M. Hare States

University of Miami, Miami, FL, United

Konstantinos E. Hatzistergos FL, United States Ralf-Peter Herber Netherlands Jo¨ns Hilborn

University of Miami, Miami,

Cam Bioceramics BV, Leiden, The

Uppsala University, Uppsala, Sweden

H. David Humes Innovative Biotherapies, Ann Arbor, MI, United States; University of Michigan, Ann Arbor, MI, United States Joshua G. Hunsberger Wake Forest Institute for Regenerative Medicine, Wake Forest University, WinstonSalem, NC, United States Kenjiro Iwasa University of California, Davis, Sacramento, CA, United States John D. Jackson Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Margaret L. Jackson McGovern Medical School at the University of Texas Health Science Center at Houston, Houston, TX, United States Hae Lin Jang Harvard Medical School, Brigham and Women’s Hospital, Boston, MA, United States; Massachusetts Institute of Technology, Cambridge, MA, United States; Harvard University, Boston, MA, United States John A. Jansen Radboudumc, Nijmegen, The Netherlands Josephine Johnston United States

The Hastings Center, Garrison, NY,

Carl Jorns Department of Transplantation Surgery, Karolinska University Hospital, Karolinska Institute, Stockholm, Sweden Huijun Kang University of Pittsburgh, Pittsburgh, PA, United States David L. Kaplan States

Tufts University, Medford, MA, United

David S. Kaplan Center for Devices and Radiological Health, FDA, Silver Spring, MD, United States Adam J. Katz States

University of Florida, Gainesville, FL, United

Matthew W. Kelley National Institutes of Health, Bethesda, MD, United States Kelsey Kennedy United States

Columbia University, New York, NY,

xxi

CONTRIBUTORS

Ali Khademhosseini Harvard Medical School, Brigham and Women’s Hospital, Boston, MA, United States; Massachusetts Institute of Technology, Cambridge, MA, United States; Harvard University, Boston, MA, United States; Konkuk University, Seoul, Republic of Korea; King Abdulaziz University, Jeddah, Saudi Arabia Gilson Khang Chonbuk National University, Jeonju-si, Republic of Korea Jinho Kim Columbia University, New York, NY, United States Rachel H. Klein United States

University of California Davis, Davis, CA,

Irina Klimanskaya Astellas Institute for Regenerative Medicine, Marlboro, MA, United States Paul S. Knoepfler United States

University of California Davis, Davis, CA,

Jeanne F. Loring The Scripps Research Institute, San Diego, CA, United States Shi-Jiang Lu Vcanbio Center for Translational Biotechnology, Natick, MA, United States Alberto Lue Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain; Lozano Blesa University Hospital, Zaragoza, Spain Peter X. Ma The University of Michigan, Ann Arbor, MI, United States Renata S. Magalhaes Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Serena Mandla Institute of Biomaterials and Biomedical Engineering, Department of Chemical Engineering and Applied Chemistry, University of Toronto, Toronto, ON, Canada

In Kap Ko Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

Clement D. Marshall Stanford University School of Medicine, Palo Alto, CA, United States

Yash M. Kolambkar Georgia Institute of Technology, Atlanta, GA, United States

Manuela Martins-Green CA, United States

Jan Krieghoff University of Leipzig, Leipzig, Germany

Devon E. Mason University of Pennsylvania, Philadelphia, PA, United States

Nathan W. Kucko Radboudumc, Nijmegen, The Netherlands; Cam Bioceramics BV, Leiden, The Netherlands Manoj Kumar

University of KU Leuven, Leuven, Belgium

Joanne Kurtzberg States

Duke University, Durham, NC, United

Anna Kwilas Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States Donald W. Landry College of Physicians and Surgeons of Columbia University, New York, NY, United States Mark T. Langhans University of Pittsburgh School of Medicine, Pittsburgh, PA, United States Robert Lanza Astellas Institute of Regenerative Medicine, Marlborough, MA, United States Giacomo Lanzoni States

University of Miami, Miami, FL, United

Sang Jin Lee Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Sander C.G. Leeuwenburgh Netherlands

Jonquil R. Mau United States

University of California, Riverside,

University of Pittsburgh, Pittsburgh, PA,

Richard McFarland Advanced Regenerative Manufacturing Institute, Manchester, NH, United States Melissa K. McHale States

Rice University, Houston, TX, United

James C. Melville University of Texas Health Science Center at Houston, Houston, TX, United States Jason R. Meyers States

Colgate University, Hamilton, NY, United

Antonios G. Mikos Rice University, Houston, TX, United States Jordan S. Miller Rice University, Houston, TX, United States Paul A. Mittermiller United States

Stanford University, Stanford, CA,

Hideki Miyachi Nationwide Children’s Hospital, Columbus, OH, United States

Radboudumc, Nijmegen, The

Shinka Miyamoto Nationwide Children’s Hospital, Columbus, OH, United States

Kam W. Leong Duke University, Durham, NC, United States; Columbia University, New York, NY, United States

Nelson Monteiro University of Connecticut Health, Farmington, CT, United States

Rui Liang University of Pittsburgh, Pittsburgh, PA, United States

Alessandra L. Moore Stanford University School of Medicine, Palo Alto, CA, United States; Brigham and Women’s Hospital, Boston, MA, United States

Volha Liaudanskaya States

Tufts University, Medford, MA, United

Hang Lin University of Pittsburgh School of Medicine, Pittsburgh, PA, United States Michael T. Longaker Stanford University School of Medicine, Palo Alto, CA, United States Hermann P. Lorenz Stanford University School of Medicine, Palo Alto, CA, United States

Sara Morini Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain; Universidade de Lisboa, Lisbon, Portugal Philipp T. Moser Massachusetts General Hospital, Boston, MA, United States Vivek J. Mukhatyar States

Duke University, Durham, NC, United

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CONTRIBUTORS

Mark Murdock University of Pittsburgh, Pittsburgh, PA, United States

Alessandro Pirosa University of Pittsburgh School of Medicine, Pittsburgh, PA, United States

Aaron Nagiel University of California Los Angeles Geffen School of Medicine, Los Angeles, CA, United States

Iris Pla-Palacı´n Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain

Gail K. Naughton States

Marta Pokrywczynska Nicolaus Copernicus University in Torun, Ludwik Rydygier Medical College in Bydgoszcz, Bydgoszcz, Poland

Histogen, Inc., San Diego, CA, United

Allison Nauta Stanford University School of Medicine, Palo Alto, CA, United States; Oregon Health and Sciences University, Portland, OR, United States Javier Navarro University of Maryland, College Park, MD, United States; Center for Engineering Complex Tissues, College Park, MD, United States Jared M. Newton United States Aparna Nori

Baylor College of Medicine, Houston, TX,

Duke University, Durham, NC, United States

Teruo Okano Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, Tokyo, Japan Joaquim M. Oliveira 3B’s Research Group, I3Bs e Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Guimara˜es, Portugal; ICVS/3B’sePT Government Associate Laboratory, Braga/Guimara˜es, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimara˜es, Portugal Harald C. Ott Massachusetts General Hospital, Boston, MA, United States; Harvard Medical School, Boston, MA, United States Jagannath Padmanabhan Stanford University, Stanford, CA, United States Kristin M. Page States

Duke University, Durham, NC, United

Anil Kumar Pallickaveedu Rajan Asari Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Virginia E. Papaioannou College of Physicians and Surgeons of Columbia University, New York, NY, United States Jihoon Park Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Samantha L. Payne Canada

University of Toronto, Toronto, ON,

Gadi Pelled Cedars-Sinai Medical Center, Los Angeles, CA, United States; Hebrew University of Jerusalem, Jerusalem, Israel Andrew Pepper Department of Surgery, University of Alberta, Edmonton, AB, Canada Elumalai Perumal The Catholic University of Korea, Seochogu, Republic of Korea Melissa Petreaca United States

DePauw University, Greencastle, IN,

Christopher J. Pino United States

Innovative Biotherapies, Ann Arbor, MI,

Christopher D. Porada Wake Forest Institute for Regenerative Medicine, Wake Forest University, WinstonSalem, NC, United States Blaise D. Porter Georgia Institute of Technology, Atlanta, GA, United States Milica Radisic Institute of Biomaterials and Biomedical Engineering, Department of Chemical Engineering and Applied Chemistry, University of Toronto, Toronto, ON, Canada; Toronto General Research Institute, Toronto, ON, Canada Kunal J. Rambhia The University of Michigan, Ann Arbor, MI, United States F. Raquel Maia 3B’s Research Group, I3Bs e Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Guimara˜es, Portugal; ICVS/3B’sePT Government Associate Laboratory, Braga/Guimara˜es, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimara˜es, Portugal Buddy D. Ratner University of Washington, Seattle, WA, United States A.H. Reddi University of California, Davis, Sacramento, CA, United States Rui L. Reis 3B’s Research Group, I3Bs e Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Guimara˜es, Portugal; ICVS/3B’sePT Government Associate Laboratory, Braga/Guimara˜es, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimara˜es, Portugal Laura Ricles Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States Camillo Ricordi University of Miami, Miami, FL, United States Muhammad Rizwan University of Waterloo, Waterloo, ON, Canada Rebecca Robinson Advanced Regenerative Manufacturing Institute, Manchester, NH, United States Melanie Rodrigues United States

Stanford University, Stanford, CA,

Benjamin B. Rothrauff University of Pittsburgh School of Medicine, Pittsburgh, PA, United States Hooman Sadri-Ardekani Wake Forest Institute for Regenerative Medicine, Wake Forest University, WinstonSalem, NC, United States Pilar Sainz-Arnal Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain; Instituto Aragone´s de Ciencias de la Salud (IACS), Zaragoza, Spain

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CONTRIBUTORS

Rangarajan Sambathkumar Leuven, Belgium

University of KU Leuven,

Estela Solanas Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain

Natalia Sa´nchez-Romero Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain

Jeong Eun Song Chonbuk National University, Jeonju-si, Republic of Korea

Michelle Scarritt United States

Disha Sood

University of Pittsburgh, Pittsburgh, PA,

Christopher M. Schneider McGovern Medical School at the University of Texas Health Science Center at Houston, Houston, TX, United States Steven D. Schwartz University of California Los Angeles Geffen School of Medicine, Los Angeles, CA, United States Sarah Selem University of Miami, Miami, FL, United States Trinidad Serrano-Aullo´ Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain; Lozano Blesa University Hospital, Zaragoza, Spain A.M. James Shapiro Department of Surgery, University of Alberta, Edmonton, AB, Canada Dmitriy Sheyn Cedars-Sinai Medical Center, Los Angeles, CA, United States Tatsuya Shimizu Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, Tokyo, Japan Toshiharu Shinoka Nationwide Children’s Hospital, Columbus, OH, United States; Ohio State University, Columbus, OH, United States Molly S. Shoichet Canada

University of Toronto, Toronto, ON,

Toshihiro Shoji Nationwide Children’s Hospital, Columbus, OH, United States Thomas Shupe Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Andrew G. Sikora Baylor College of Medicine, Houston, TX, United States Fiona Simpson Maisonneuve-Rosemont Hospital Research Centre, Montreal, QC, Canada Aleksander Skardal Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States; Virginia Tech-Wake Forest School of Biomedical Engineering and Sciences, Wake Forest University, Winston-Salem, NC, United States; Comprehensive Cancer Center at Wake Forest Baptist Medical, Winston-Salem, NC, United States; Department of Cancer Biology, Wake Forest University, Winston-Salem, NC, United States Daniel Skuk Axe Neurosciences, Research Center of the CHU de QuebeceCHUL, Quebec, QC, Canada Brandon T. Smith Rice University, Houston, TX, United States Jihee Sohn University of Pittsburgh School of Medicine, Pittsburgh, PA, United States Shay Soker Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

Tufts University, Medford, MA, United States

David L. Stocum Indiana University-Purdue University, Indianapolis, IN, United States Stephen C. Strom Department of Laboratory Medicine, Karolinska Institute and Division of Pathology, Karolinska University Hospital, Stockholm, Sweden Jessica M. Sun

Duke University, Durham, NC, United States

Hironobu Takahashi Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, Tokyo, Japan Jacques P. Tremblay Axe Neurosciences, Research Center of the CHU de QuebeceCHUL, Quebec, QC, Canada Nirmalya Tripathy Chonbuk National University, Jeonju-si, Republic of Korea John W. Tse

University of Waterloo, Waterloo, ON, Canada

Rocky S. Tuan University of Pittsburgh School of Medicine, Pittsburgh, PA, United States Catherine M. Verfaillie Belgium

University of KU Leuven, Leuven,

Gordana Vunjak-Novakovic York, NY, United States

Columbia University, New

William R. Wagner University of Pittsburgh, Pittsburgh, PA, United States Yanling Wang The Scripps Research Institute, San Diego, CA, United States; Summit for Stem Cell Foundation, San Diego, CA, United States Emma Watson

Rice University, Houston, TX, United States

Jennifer L. West Duke University, Durham, NC, United States David F. Williams Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States James K. Williams Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Mark E. Wong University of Texas Health Science Center at Houston, Houston, TX, United States Savio L-Y. Woo University of Pittsburgh, Pittsburgh, PA, United States Fiona M. Wood Australia

University of Western Australia, Perth,

Lei Xu Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States Doron C. Yakubovich Jerusalem, Israel

Hebrew University of Jerusalem,

Yafeng Yang Harvard Medical School, Brigham and Women’s Hospital, Boston, MA, United States; Massachusetts Institute of Technology, Cambridge, MA, United States

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CONTRIBUTORS

Michael J. Yaszemski Departments of Orthopedic Surgery and Biomedical Engineering, Mayo Clinic, Rochester, MN, United States Pamela C. Yelick Tufts University School of Dental Medicine, Boston, MA, United States Evelyn K.F. Yim University of Waterloo, Waterloo, ON, Canada; National University of Singapore, Singapore, Singapore Carolyn Yong Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States James J. Yoo Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Simon Young The University of Texas Health Science Center at Houston, School of Dentistry, Houston, TX, United States

Nora Yucel Stanford University School of Medicine, Stanford, CA, United States Rachel L. Zacharias United States

The Hastings Center, Garrison, NY,

Yuanyuan Zhang Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States Ai Zhang The Scripps Research Institute, San Diego, CA, United States Jin Zhang Harvard Medical School, Brigham and Women’s Hospital, Boston, MA, United States; Massachusetts Institute of Technology, Cambridge, MA, United States Yang Zhu States

University of Pittsburgh, Pittsburgh, PA, United

Preface

The textbook Principles of Regenerative Medicine was created as a primary resource for scientists, clinicians, teachers, and students from academia, industry, and government, as well as the public at large. The initial edition was the first comprehensive body of work dedicated entirely to the field and quickly became the most relevant textbook in the arena of regenerative medicine. The first and second editions of the textbook have had broad appeal and are currently ranked as the most distributed textbooks in the field. I am honored to have had the opportunity to continue to edit the textbook, now in its third edition, with our coeditors, Robert Lanza, Tony Mikos, and Robert Nerem. We welcome Tony Mikos as a new editor and would like to thank Jamie Thomson for his contributions to the first two editions. The contributions of the editors cannot be overestimated. We are indebted to their vision and the strong foundation they have created, upon which the current text was built. The specialty of regenerative medicine continues to grow and change rapidly. There have been major areas of advances in just the past few years. The field encompasses multiple areas of scientific inquiry, each of which is complex, but together they are a powerful combination of technologies, such as stem cells, gene editing, nuclear transfer, proteomics, pharmacology, nanotechnology, tissue engineering, three-dimensional printing, and biomanufacturing. We are now in an era of translation of bench site discoveries to clinical therapies. We hope that this book will enlighten all of these areas and supply guidance where it is most needed. The textbook was organized in a manner that builds upon the basic science of the field and goes forward into clinical applications and possible clinical utility. The textbook is organized into seven major areas, starting with chapters that encompass some of the fundamentals of the field. The biologic and molecular bases of regenerative medicine are covered with the molecular, mechanistic, and phenotypic aspects of cells and the extracellular matrix. The second section encompasses cells and tissue development, dealing with the various types of cells and determinants of tissue formation. The third section explores the mechanical, physical, and morphogenetic aspects of regenerative medicine, including nanotechnology applications. The fourth section is dedicated to the area of biomaterials for regenerative medicine. The fifth section covers enabling technologies for regenerative medicine, including infection control, gene editing, preclinical models, bioreactors, bioprinters, and body-on-a-chip applications. The sixth section discusses the topics of therapeutic applications and deals mostly with specific tissue and organ types. The last section of the book is dedicated to the regulatory, manufacturing, policy, and ethical aspects of the field. This area is becoming increasingly more important because the nexus between science, safety, and ethics is constantly changing. The third edition of the Principles of Regenerative Medicine builds on the knowledge base of the prior editions; most relevant chapters have been updated and other relevant topics are added with the inclusion of new chapters. Once again, the leading scientists and physicians in the field have contributed their expertise to this comprehensive collection. We would like to thank all of the authors for their contributions; the publisher, Elsevier, and its staff, including Timothy Bennett, who worked diligently to complete the book; and our commissioning editors, Micah Lewis and Elizabeth Brown, who nurtured this project from its concept to completion. Anthony Atala (on behalf of the Editors)

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C H A P T E R

1 Molecular Organization of Cells Jon D. Ahlstrom PolarityTE, Salt Lake City, UT, United States

INTRODUCTION Multicellular tissues exist in one of two types of cellular arrangements, epithelial or mesenchymal. Epithelial cells adhere tightly to each other at their lateral surfaces and to an organized extracellular matrix (ECM) at their basal domain, thereby producing a sheet of cells resting on a basal lamina with an apical surface. Mesenchymal cells, in contrast, are individual cells with a bipolar morphology that are held together as a tissue within a threedimensional (3D) ECM (Fig. 1.1). The conversion of epithelial cells into mesenchymal cells, an “epitheliale mesenchymal transition” (EMT), is central to many aspects of embryonic morphogenesis and adult tissue repair, as well as a number of disease states [1e3]. The reverse process whereby mesenchymal cells coalesce into an epithelium is a “mesenchymaleepithelial transition” (MET). Understanding the molecules that regulate this transition between epithelial and mesenchymal states offers important insights into how cells and tissues are organized. The early embryo is structured as one or more epithelia. An EMT allows the rearrangements of cells to create additional morphological features. Well-studied examples of EMTs during embryonic development include gastrulation in Drosophila [3], the emigration of primary mesenchyme cells (PMCs) in sea urchin embryos [4], and gastrulation in amniotes (reptiles, birds, and mammals) at the primitive streak [2]. EMTs also occur later in vertebrate development, such as during the emigration of neural crest cells from the neural tube [5], the formation of the sclerotome from epithelial somites, and palate fusion [2]. The reverse process of MET is likewise crucial to development; examples include the condensation of mesenchymal cells to form the notochord and somites [6], kidney tubule formation from nephrogenic mesenchyme [7], and the creation of heart valves from cardiac mesenchyme [8]. In the adult organism, EMTs and METs occur during wound healing and tissue remodeling [9]. The conversion of neoplastic epithelial cells into invasive cancer cells has long been considered an EMT process [1,10]. However, there are also examples of tumor cells that have functional cellecell adhesion junctions yet are still migratory and invasive as a group [11]. This “collective migration” also occurs during development [11]. Hence, there is debate whether an EMT model accurately describes all epithelial metastatic cancers. Similarly, the fibrosis of cardiac, kidney, lens, and liver epithelial tissue has also long been categorized as an EMT event [6,12]. However, research in the kidney in vivo shows that the myofibroblasts induced after kidney injury are derived from mesenchymal pericytes rather than the proximal epithelial cells [13]. Therefore, defining the origin of the cells that contribute to fibrotic tissue scarring (epithelial or otherwise) will require further investigation. The focus of this chapter is on the molecules that regulate the organization of cells into epithelium or mesenchyme. We will first discuss the cellular changes that occur during an EMT, including changes in cellecell and celle ECM adhesions, changes in cell polarity, and the stimulation of invasive cell motility. Then we will consider the molecules and mechanisms that control the EMT or MET, including the structural molecules, transcription factors, and signaling pathways that regulate EMTs.

MOLECULES THAT ORGANIZE CELLS The conversion of an epithelial sheet into individual migratory cells and back again requires the coordinated changes of many distinct families of molecules. Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00001-1

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Copyright © 2019 Elsevier Inc. All rights reserved.

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1. MOLECULAR ORGANIZATION OF CELLS

FIGURE 1.1 Epithelial versus Mesenchymal. Epithelial cells adhere tightly together by tight junctions and adherens junctions localized near the apical surface. Epithelial cells also have a basal surface that rests on a basal lamina. In contrast, mesenchymal cells do not have well-defined cellecell adhesion complexes; they have front-end/back-end polarity instead of apicobasal polarity, and mesenchymal cells are characterized by their ability to invade the basal lamina.

Changes in CelleCell Adhesion Epithelial cells are held together by specialized cellecell junctions, including adherens junctions, desmosomes, and tight junctions [14]. These junctions are localized in the lateral domain near the apical surface and establish the apical polarity of the epithelium. For an epithelial sheet to produce individual mesenchymal cells, cellecell adhesions must be disrupted. The principal transmembrane proteins that mediate cellecell adhesions are members of the cadherin superfamily [15]. E-cadherin and N-cadherin are classical cadherins that interact homotypically through their extracellular immunoglobulin G domains with like-cadherins on adjacent cells. Cadherins are important mediators of cellecell adhesion. For example, misexpression of E-cadherin is sufficient to promote cellecell adhesion and assembly of adherens junctions in fibroblasts [16]. In epithelial cancers (carcinomas), E-cadherin acts as a tumor suppressor [10]. In a mouse model for b-cell pancreatic cancer, the loss of E-cadherin is the ratelimiting step for transformed epithelial cells to become invasive [17]. Although the loss of cadherin-mediated celle cell adhesion is necessary for an EMT, the loss of cadherins is not always sufficient to generate a complete EMT in vivo. For example, the neural tube epithelium in mice expresses N-cadherin, but in the N-cadherin knockout mouse an EMT is not induced in the neural tube [18]. Hence, cadherins are essential for maintaining epithelial integrity, and the loss of cellecell adhesion caused by the reduction of cadherin function is an important step for an EMT. One characteristic of an EMT is “cadherin switching.” Often, epithelia that express E-cadherin will downregulate E-cadherin expression at the time of the EMT and express different cadherins such as N-cadherin [19]. Cadherin switching may promote motility. For instance, in mammary epithelial cell lines, the misexpression of N-cadherin is sufficient for increased cell motility. However, blocking N-cadherin expression does not result in motility even though the adherens junctions are reduced. Hence, cadherin switching may be necessary for cell motility, but cadherin switching alone is not sufficient to bring about a complete EMT [20]. There are several ways in which cadherin expression and function are regulated. Transcription factors that are central to most EMTs, such as Snail-1, Snail-2, Zinc finger E-box-binding (Zeb)1, Zeb2, Twist, and E2A, all bind to E-boxes on the E-cadherin promoter and repress the transcription of E-cadherin [21]. Posttranscriptionally, the Ecadherin protein is ubiquitinated by the E3-ligase, Hakai, which targets E-cadherin to the proteasome [22]. E-cadherin turnover at the membrane is regulated by either caveolae-dependent endocytosis or clathrindependent endocytosis [23], and p120-catenin prevents endocytosis of E-cadherin at the membrane [24]. E-cadherin function can also be disrupted by matrix metalloproteases, which degrade the extracellular domain of E-cadherin [25]. Some or all of these mechanisms may occur during an EMT to disrupt cellecell adhesion. Cellecell adhesion is maintained principally by cadherins, and changes in cadherin expression are typical of an EMT.

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Changes in CelleExtracellular Matrix Adhesion Altering the way in which a cell interacts with the ECM is also important in EMTs. For example, at the time that sea urchin PMCs ingress, the cells have increased adhesiveness for ECM [4]. CelleECM adhesion is mediated principally by integrins. Integrins are transmembrane proteins composed of two noncovalently linked subunits, a and b, that bind to ECM components such as fibronectin, laminin, and collagen. The cytoplasmic domain of integrins links to the cytoskeleton and interacts with signaling molecules. Changes in integrin function are required for many EMTs, including neural crest emigration [26], mouse primitive streak formation [2], and cancer metastasis [27]. However, the misexpression of integrin subunits is not sufficient to bring about a full EMT in vitro [28] or in vivo [29]. The presence and function of integrins are modulated in several ways. For example, the promoter of the integrin b6 gene is activated by the transcription factor Ets-1 during colon carcinoma metastasis [30]. Most integrins can also cycle between “on” (high-affinity) or “off” (low-affinity) states. This “inside-out” regulation of integrin adhesion occurs at the integrin cytoplasmic tail [31]. In addition to integrin activation, the “clustering” of integrins on the cell surface affects the overall strength of integrineECM interactions. The increased adhesiveness of integrins caused by clustering, known as avidity, can be activated by chemokines and depends on RhoA and phosphatidylinositol 30 kinase (PI3K) activity [31]. Changes in ECM adhesion are required for an EMT. CelleECM adhesions are maintained by integrins, which have varying degrees of adhesiveness depending on the presence, activity, or avidity of the integrin subunits.

Changes in Cell Polarity and Stimulation of Cell Motility Cellular polarity is defined by the distinct arrangement of cytoskeletal elements and organelles in epithelial versus mesenchymal cells. Epithelial polarity is characterized by cellecell junctions found near the apicolateral domain (nonadhesive surface), and a basal lamina opposite of the apical surface (adhesive surface). Mesenchymal cells, in contrast, do not have apicobasal polarity, but rather front-end/back-end polarity, with actin-rich lamellipodia and Golgi localized at the leading edge [2]. Molecules that establish cell polarity include Cdc42, PAK1, PI3K, PTEN, Rac, Rho, and the PAR proteins [32,33]. Changes in cell polarity help to promote an EMT. In mammary epithelial cells, the activated transforming growth factor-b (TGF-b) receptor II causes Par6 to activate the E3 ubiquitin ligase Smurf1, which then targets RhoA to the proteasome. The loss of RhoA activity results in the loss of cellecell adhesion and epithelial cell polarity [34]. For mesenchymal cells to leave the epithelium, they must become motile. Many of the same polarity (Crumbs, PAR, and Scribble complexes), structural (actin and microtubules), and regulatory molecules (Cdc42, Rac1, and RhoA) that govern epithelial polarity are also central to cell motility [35]. Cell motility mechanisms also vary depending on whether the environment is two-dimensional or 3D [36]. Many mesenchymal cells express the intermediate filament vimentin, which may be responsible for several aspects of the EMT phenotype [37]. In short, a wide variety of structural, polarity, and regulatory molecules must be reassigned as cells transition between epithelial polarity and mesenchymal migration.

Invasion of the Basal Lamina In most EMTs, the emerging mesenchymal cells must penetrate a basal lamina, which consists of ECM components such as collagen type IV, fibronectin, and laminin. The basal lamina functions to stabilize the epithelium and is a barrier to migratory cells [38]. One mechanism that mesenchymal cells use to breach the basal lamina is to produce enzymes that degrade it. Plasminogen activator is one protease associated with a number of EMTs, including neural crest emigration [38] and the formation of cardiac cushion cells during heart morphogenesis [39]. The type II serine protease TMPRSS4 also promotes an EMT and metastasis when overexpressed in vitro and in vivo [40]. Matrix-metalloprotease (MMPs) are also important for many EMTs. When MMP-2 activity is blocked in the neural crest EMT, neural crest emigration is inhibited, but not neural crest motility [41]. In mouse mammary cells, MMP-3 overexpression is sufficient to induce an EMT in vitro and in vivo [42]. Misexpressing MMP-3 in cultured cells induces an alternatively spliced form of Rac1 (Rac1b), which then causes an increase in reactive oxygen species (ROS) intracellularly, and Snail-1 expression. Either Rac1b activity or ROS is necessary and sufficient for an MMP-3einduced EMT [43]. Hence, a number of extracellular proteases are important to bring about an EMT.

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Although epithelial cells undergoing an EMT eventually lose cellecell adhesion and apicobasal polarity and gain invasive motility, the EMT program is not necessarily ordered or linear. For example, in a study in which neural crest cells were labeled with cell adhesion or polarity markers and individual live cells were observed to undergo the EMT in slice culture, neural crest cells changed epithelial polarity either before or after the complete loss of cellecell adhesion, or lost cellecell adhesion either before or after cell migration commenced [44]. Therefore, whereas an EMT consists of several distinct phases, these steps may occur in different orders or combinations, some of which (e.g., the complete loss of cell-cell adhesion) may not always be necessary. Changes in a wide range of molecules are needed for an EMT because epithelial cells lose cellecell adhesion, change cellular polarity, and gain invasive cell motility.

THE EPITHELIALeMESENCHYMAL TRANSITION TRANSCRIPTIONAL PROGRAM At the foundation of every EMT or MET program are the transcription factors that regulate the gene expression required for these cellular transitions. Whereas many of the transcription factors that regulate EMTs have been identified, the complex regulatory networks are still incomplete. Here we review the transcription factors that are known to promote the various phases of an EMT. Then we examine how these EMT transcription factors themselves are regulated at the promoter and posttranscriptional levels.

Transcription Factors That Regulate EpithelialeMesenchymal Transition The Snail family of zinc-finger transcription factors, including Snail-1 and Snail-2 (formerly Snail and Slug), are direct regulators of cellecell adhesion and motility during EMTs [21,45]. The knockout of Snail-1 in mice is lethal early in gestation, and the presumptive primitive streak cells that normally undergo an EMT retain apicobasal polarity and adherens junctions, and express E-cadherin messenger RNA [46]. Snail-1 misexpression is sufficient for breast cancer recurrence in a mouse model in vivo, and high levels of Snail-1 predict the relapse of human breast cancer [47]. Snail-2 is necessary for the chicken primitive streak and neural crest EMTs [48]. One way in which Snail-1 or Snail-2 causes decreases in cellecell adhesion is by repressing the E-cadherin promoter [21]. This repression requires the mSin3A co-repressor complex, histone deacetylases, and components of the Polycomb 2 complex [49]. Snail-1 is also a transcriptional repressor of the tight junction genes Claudin and Occludin [21] and the polarity gene Crumbs3 [50]. The misexpression of Snail-1 and Snail-2 further leads to the transcription of proteins important for cell motility, such as fibronectin, vimentin [51], and RhoB [52]. Moreover, Snail-1 promotes invasion across the basal lamina. In MadineDarby Canine Kidney (MDCK) cells, the misexpression of Snail-1 represses laminin (basement membrane) production [53] and indirectly upregulates mmp-9 transcription [54]. Snail and Twist also make cancer cells more resistant to senescence, chemotherapy, and apoptosis, and endow cancer cells with “stem cell” properties [6]. Hence, Snail-1 or Snail-2 is necessary and sufficient for bringing about many of the steps of an EMT, including loss of cellecell adhesion, changes in cell polarity, gain of cell motility, invasion of the basal lamina, and increased proliferation and survival. Other zinc finger transcription factors important for EMTs are Zeb homeobox 1 (Zeb1; also known as dEF1), and Zeb2 (also known as Smad-interacting protein-1; Sip1). Both Zeb1 and Zeb2 bind to the E-cadherin promoter and repress transcription [21]. Zeb1 can also bind to and repress the transcription of the polarity proteins Crumbs3, Pals1-associated tight junction proteins, and Lethal giant larvae 2 [55]. Zeb2 is structurally similar to Zeb1, and Zeb2 overexpression is sufficient to downregulate E-cadherin, dissociate adherens junctions, and increase motility in MDCK cells [56]. The lymphoid enhancer-binding factor/T-cell factor (LEF/TCF) transcription factors also have an important role in EMTs. For instance, the misexpression of Lef-1 in cultured colon cancer cells reversibly causes the loss of cellecell adhesion [57]. LEF/TCF transcription factors directly activate genes that regulate cell motility, such as the L1 adhesion molecule [58], and the fibronectin gene [59]. LEF/TCF transcription factors also upregulate genes required for basal lamina invasion, including mmp-3 and mmp-7 [60]. Other transcription factors that have a role in promoting EMTs are the class I basic helix-loop-helix factors E2-2A and E2-2B [61], the forkhead box transcription factor FOXC2 [62], the homeobox protein Goosecoid [63], and the homeoprotein Six1 [64,65]. Transcription factors that regulate an EMT often do so by directly repressing cell adhesion and epithelial polarity molecules and by upregulating genes required for cell motility and basal lamina invasion.

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Regulation at the Promoter Level Given the importance of the Snail, Zeb, and LEF/TCF transcription factors in orchestrating the various phases of an EMT, it is essential to understand the upstream events that regulate these EMT-promoting transcription factors. The activation of Snail-1 transcription in Drosophila requires the transcription factors Dorsal (nuclear factor kB [NF-kB]) and Twist [21]. The human Snail-1 promoter also has functional NF-kB sites [66], and blocking NF-kB reduces Snail-1 transcription [67]. In addition, a region of the Snail-1 promoter is responsive to integrin-linked kinase (ILK) [21], and ILK can activate Snail-1 expression via poly-adenosine phosphate-ribose polymerase [68]. In mouse mammary epithelial cells, high-mobility group protein A2 and Smads activate Snail-1 expression and subsequently Snail-2, Twist, and Id2 transcription [69]. For Snail-2 expression, myocardin-related transcription factors (MRTFs) interact with Smads to induce Snail-2 [70] and MRTFs may have a role in metastasis [71] and fibrosis [72]. There are also several Snail-1 transcriptional repressors. In breast cancer cell lines, metastasis-associated protein 3 binds directly to and represses the transcription of Snail-1 in combination with the Mi-2/nucleosome remodeling deacetylase complex [73], as does lysine-specific demethylase [74]. The Ajuba LIM proteins (Ajuba, LIMD1, and WTIP) are additional transcriptional corepressors of the Snail family [75]. The transcription of LEF/TCF genes such as Lef-1 are activated by Smads [76]. The misexpression of Snail-1 results in the transcription of dEF-1 and Lef-1 through a yet unknown mechanism [21].

Posttranscriptional Regulation of EpithelialeMesenchymal Transition Transcription Factors The activity of EMT transcription factors is also regulated at the protein level, including translational control, protein stability (targeting to the proteasome), and nuclear localization. Noncoding RNAs are emerging as important regulator EMTs. In a breast cancer model, Myc activates the expression of microRNA-9 (miR-9), and miR-9 directly binds to and represses the E-cadherin promoter [77]. Members of the miR-200 family repress the translation of Zeb1, and the expression of these miR-200 family members is repressed by Snail-1. In addition, Zeb2 transcription can be activated by naturally occurring RNA antisense transcripts [78]. It is not yet known whether there are noncoding RNAs that regulate Snail family members. However, the Y-box binding protein-1 is important for the selective activation of Snail-1 translation [79]. Protein stability is another layer of EMT control. Snail-1 is phosphorylated by glycogen synthase kinase 3b (GSK3b) and targeted for destruction [80]. Therefore, the inhibition of GSK-3b activity by Wnt signaling may have multiple roles in an EMT, leading to the stabilization of both b-catenin and Snail-1. Some proteins that prevent GSK 3be mediated phosphorylation (and thus promote Snail-1 activation) are lysyl oxidaseelike proteins (LOXL)2, LOXL3 [81], and ILK [82]. A Snail 1especific phosphatase (Snail-1 activator) is C-terminal domain phosphatase [83]. Snail-2 is targeted for degradation by the direct action of p53 and the ubiquitin ligase Mdm2 [84]. In addition to protein translation and stability, the function of Snail-1 depends on nuclear localization mediated by Snail-1’s nuclear localization sequence. The phosphorylation of human Snail-1 by p21-activated kinase 1 promotes the nuclear localization of Snail-1 (and therefore Snail-1 activation) in breast cancer cells [85]. In zebrafish, LIV-1 promotes the translocation of Snail-1 into the nucleus [86]. Snail-1 also contains a nuclear export sequence (NES) that depends on the calreticulin nuclear export pathway [87]. This NES sequence is activated by the phosphorylation of the same lysine residues targeted by GSK-3b, which suggests a mechanism whereby phosphorylation of Snail-1 by GSK-3b results in the export of Snail-1 from the nucleus and subsequent degradation. LEF/TCF activity is also regulated by other proteins. b-Catenin is required as a cofactor for LEF/TCF-mediated activation of transcription, and Lef-1 can also associate with cofactor Smads to activate the transcription of additional EMT genes [88]. In colon cancer cells, thymosin b4 stabilizes ILK activity [89]. EMT transcription factors such as Snail-1, Zeb1, and Lef-1 are regulated by a variety of mechanisms at both the transcriptional and posttranscriptional levels by noncoding RNA translation control, protein degradation, nuclear localization, and cofactors such as b-catenin.

MOLECULAR CONTROL OF THE EPITHELIALeMESENCHYMAL TRANSITION The initiation of an EMT or MET is a tightly regulated event during development and tissue repair because deregulation of cellular organization is disastrous to the organism. A variety of external and internal signaling mechanisms coordinate the complex events of the EMT, and these same signaling pathways are often disrupted or reactivated during disease. EMTs or METs can be induced by either diffusible signaling molecules or ECM

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components. We next discuss the role of signaling molecules and ECM in triggering an EMT, and then present a summary model for EMT induction.

Ligand-Receptor Signaling During development, five main ligand-receptor signaling pathways are employed: TGF-b, Wnt, receptor tyrosine kinase (RTK), Notch, and Hedgehog. These pathways, among others, have a role in triggering EMTs. Although the activation of a single signaling pathway can be sufficient for an EMT, in most cases an EMT or MET is initiated by multiple signaling pathways acting in concert. Growth Factor-b Pathway The TGF-b superfamily includes the TGF-b, activin, and bone morphogenetic protein (BMP) families. These ligands operate through receptor serine/threonine kinases to activate a variety of signaling molecules including Smads, mitogen-activated protein kinase (MAPK), PI3K, and ILK. Most EMTs studied to date are induced in part, or solely, by TGF-b superfamily members [90]. During embryonic heart development, TGF-b2 and TGF-b3 have sequential and necessary roles in activating the endocardium to invade the cardiac jelly and from the endocardial cushions [91]. In the avian neural crest, BMP4 induces Snail-2 expression [92]. In the EMT that transforms epithelial tissue into metastatic cancer cells, TGF-b acts as a tumor suppressor during early stages of tumor development, but as a tumor/EMT inducer at later stages [90,93]. TGF-b signaling may combine with other signaling pathways to induce an EMT. For example, in cultured breast cancer cells, activated Ras and TGF-b induce an irreversible EMT [94], and in pig thyroid epithelial cells, TGF-b and epidermal growth factor (EGF) synergistically stimulate the EMT [95]. One outcome of TGF-b signaling is to change epithelial cell polarity immediately. In a TGF beinduced EMT of mammary epithelial cells, TGF-bRII directly phosphorylates the polarity protein, Par6, leading to the dissolution of tight junctions [34]. TGF-b signaling also regulates gene expression through the phosphorylation and activation of Smads. Smads are important cofactors in the stimulation of an EMT. For example, Smad3 is necessary for a TGF beinduced EMT in lens and kidney tissue in vivo [96]. The Smad3/4 also complex with Snail-1 and co-repress the promoters of cellecell adhesion molecules [97]. Furthermore, TGF-bRI directly binds to and activates PI3K [98], which in turn activates ILK and downstream pathways. ILK is emerging as an important positive regulator of EMTs [99]. ILK interacts directly with growth factor receptors (TGF-b, Wnt, or RTK), integrins, the actin skeleton, PI3K, and focal adhesion complexes. ILK directly phosphorylates Akt and GSK-3b, and results in the subsequent activation of transcription factors such as AP-1, NF-kB, and Lef-1. Overexpression of ILK in cultured cells causes the suppression of GSK-3b activity [82], translocation of b-catenin to the nucleus, activation of Lef-1/b-catenin transcription factors, and the downregulation of E-cadherin [100]. Inhibition of ILK in cultured colon cancer cells leads to the stabilization of GSK-3b activity, decreased nuclear b-catenin localization, the suppression of Lef-1 and Snail-1 transcription, and reduced invasive behavior of colon cancer cells [101]. ILK activity also results in Lef 1emediated transcriptional upregulation of MMPs [60]. Hence, ILK (inducible by TGF-b signaling) is capable of orchestrating most of the major events in an EMT, including the loss of cellecell adhesion and invasion across the basal lamina. Wnt Pathway Many EMTs or METs are also regulated by Wnt signaling. Wnts signal through seven-pass transmembrane proteins of the Frizzled family, which activates G-proteins and PI3K, inhibits GSK-3b, and promotes nuclear b-catenin signaling. For example, during zebrafish gastrulation, Wnt11 activates the GTPase Rab5c, which results in the endocytosis of E-cadherin [102]. Wnt6 signaling is sufficient for increased transcription of Snail-2 in the avian neural crest [103]. Snail-1 expression increases Wnt signaling [104], which suggests a positive feedback loop. One of the downstream signaling molecules activated by Wnt signaling is b-catenin. b-Catenin is a structural component of adherens junctions. Nuclear b-catenin is also a limiting factor for the activation of LEF/TCF transcription factors. b-Catenin is pivotal for regulating most EMTs. Interfering with nuclear b-catenin signaling blocks the ingression of sea urchin PMCs [105], and in b-catenin mouse knockouts, the primitive streak EMT does not occur and no mesoderm is formed [106]. b-Catenin is also necessary for the EMT that occurs during cardiac cushion development [107]. In breast cancer, b-catenin expression is highly correlated with metastasis and poor survival [108], and blocking b-catenin function in tumor cells inhibits invasion in vitro [109]. It is unclear whether b-catenin overexpression alone is sufficient for all EMTs. If b-catenin is misexpressed in cultured cells, it causes apoptosis [110]. However,

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the misexpression of a stabilized from of b-catenin in mouse epithelial cells in vivo results in metastatic skin tumors [111]. Signaling by Receptor Tyrosine Kinase Ligands The RTK family of receptors and the growth factors that activate them also regulate EMTs or METs. Ligand binding promotes RTK dimerization and activation of the intracellular kinase domains by autophosphorylation of tyrosine residues. These phosphotyrosines act as docking sites for intracellular signaling molecules, which can activate signaling cascades such as Ras/MAPK, PI3K/Akt, JAK/STAT, or ILK. We will cite a few examples of RTK signaling in EMTs and METs. Hepatocyte growth factor (HGF; also known as scatter factor) acts through the RTK c-met. HGF is important for the MET in the developing kidney [112]. HGF signaling is required for the EMT that produces myoblasts (limb muscle precursors) from somite tissue in the mouse [10]. In epithelial cells, HGF causes an EMT through MAPK and early growth response factor-1 signaling [113]. Fibroblast growth factor (FGF) signaling regulates mouse primitive streak formation [114]. FGF signaling also stimulates cell motility and activates MMPs [115,116]. EGF promotes the endocytosis of E-cadherin [117]. EGF can also increase Snail-1 activity via the inactivation of GSK3-b [118], and EGF promotes increased Twist expression through a JAK/STAT3 pathway [119]. Insulin growth factor (IGF) signaling induces an EMT in breast cancer cell lines by activating Akt2 and suppressing Akt1 [120]. In prostate cancer cells, IGF-1 promotes Zeb-1 expression [121]. In fibroblast cells, constitutively activated IGF-IR increases NF-kB activity and Snail-1 levels [122]. In several cultured epithelial cell lines, IGFR1 is associated with the complex of E-cadherin and b-catenin, and the ligand IGF-II causes the redistribution of b-catenin from the membrane to the nucleus, activation of the transcription factor TCF-3, and a subsequent EMT [123]. Another RTK known for its role in EMTs is the ErbB2/HER-2/Neu receptor, whose ligand is heregulin/neuregulin. Overexpression of HER-2 occurs in 25% of human breast cancers, and the misexpression of HER-2 in mouse mammary tissue in vivo is sufficient to cause metastatic breast cancer [124]. Herceptin (antibody against the HER-2 receptor) treatment is effective in reducing the recurrence of HER 2epositive metastatic breast cancers. HER-2 signaling activates Snail-1 expression in breast cancer through an unknown mechanism [47]. The RTK Axl is also required for breast cancer carcinoma invasiveness [125]. Vascular endothelial growth factor (VEGF) signaling promotes Snail-1 activity by suppressing GSK3-b [126] and results in increased levels of Snail-1, Snail-2, and Twist [127]. Snail-1 can also activate the expression of VEGF [128]. RTK signaling is important for many EMTs. Notch Pathway The Notch signaling family also regulates EMTs. When the Notch receptor is activated by its ligand Delta, an intracellular portion of the Notch receptor ligand is cleaved and transported to the nucleus, where it regulates target genes. Notch1 is required for cardiac endothelial cells to undergo an EMT to make cardiac cushions, and the role of Notch may be to make cells competent to respond to TGF-b2 [129]. In the avian neural crest EMT, Notch signaling is required for the induction and/or maintenance of BMP4 expression [130]. Similarly, Notch signaling is required for the TGF beinduced EMT of epithelial cell lines [131], and Notch promotes Snail-2 expression in cardiac cushion cells [132] and cultured cells [133]. Hedgehog Pathway The hedgehog pathway is also involved in EMTs. Metastatic prostate cancer cells express high levels of hedgehog and Snail-1. If prostate cancer cell lines are treated with the hedgehog-pathway inhibitor, cyclopamine, levels of Snail-1 are decreased. If the hedgehog-activated transcription factor, Gli, is misexpressed, Snail-1 expression increases [134].

Additional Signaling Pathways Other signaling pathways that activate EMTs include inflammatory signaling molecules, lipid hormones, ROS species, and hypoxia. Interleukin-6 (inflammatory and immune response) can promote Snail-1 expression in breast cancer cells [135], and Snail-1, in turn, can activate interleukin-6 expression [136], providing a link between inflammation and EMTs [137]. The lipid hormone prostaglandin E2 (PGE2) induces Zeb1 and Snail activity in lung cancer cells [138], and Snail-1 can also induce PGE2 expression [139]. ROS species can also activate EMTs by PKC and

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FIGURE 1.2 Induction of an epithelialemesenchymal transition (EMT). This figure summarizes some of the important molecular pathways that bring about an EMT. Many of the signaling pathways converge on the activation of Snail-1 and nuclear b-catenin signaling to change gene expression, which results in the loss of epithelial cell polarity, the loss of cellecell adhesion, and increased invasive cell motility. BMP, bone morphogenetic protein; CalR, calreticulin; GSK-3b, glycogen synthase kinase 3b; Igl2, immunoglobulin 2; ILK, integrin-linked kinase; LOX, lysyl oxidase; miR-200, microRNA-200; mmp, matrix-metalloprotease; NF-kB, nuclear factor kB; RhoA, Ras homolog gene family, member A; RTK, receptor tyrosine kinase; TGF-b, transforming growth factor-b; Zeb1, zinc finger E-box-binding homeobox 1.

MAPK signaling [140]. Hypoxia is important for initiating EMTs during development [141] and disease [137], often through hypoxia-inducible factor-1, which directly activates Twist expression [142]. Hypoxia also activates lysyl oxidases, which stabilize Snail-1 expression [143] by inhibiting GSK-3b activity [144]. In addition to diffusible signaling molecules, ECM molecules regulate EMTs or METs. This was first dramatically demonstrated when lens or thyroid epithelium was embedded in collagen gels, and then promptly underwent an EMT [2]. Integrin signaling appears to be important in this process [145] and involves ILK-mediated activation of NF-kB, Snail-1, and Lef-1 [146]. Other ECM components that regulate EMTs include hyaluronan [147], the g-2 chain of laminin 5 [148], periostin [149], and podoplanin [150,151]. A variety of diffusible signals and ECM components can stimulate EMTs or METs.

A Model for EpithelialeMesenchymal Transition Induction Many experimental studies on EMT mechanisms are piecework, and although great progress has been made in discovering EMT pathways, the entire signaling network is still incomplete. Fig. 1.2 shows many of the various signaling mechanisms, although in actuality only a few of the inductive pathways will be used for individual EMTs. From experimental evidence, it appears that many EMT signaling pathways converge on ILK, the inhibition of GSK-3b, and stimulation of nuclear b-catenin signaling to activate Snail and LEF/TCF transcription factors. Snail, Zeb, and LEF/TCF transcription factors then act on a variety of targets to suppress cellecell adhesion, induce changes in cell polarity, stimulate cell motility, and promote invasion of the basal lamina.

CONCLUSION Since the term “EMT” was coined [10], important insights have been made in this rapidly expanding field of research. EMT and MET events occur during development, tissue repair, and disease, and many molecules that regulate the various EMTs or METs have been characterized, thanks in large part to the advent of cell culture models. However, the EMT regulatory network as a whole is still incomplete. An improved understanding of EMT and MET pathways will lead to more effective strategies for tissue engineering and novel therapeutic targets for the treatment of disease.

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List of Acronyms and Abbreviations BMP Bone morphogenetic protein ECM Extracellular matrix EGF Epidermal growth factor EMT Epithelialemesenchymal transition FGF Fibroblast growth factor GSK-3b Glycogen synthase kinase 3b HGF Hepatocyte growth factor IGF Insulin growth factor ILK Integrin-linked kinase LEF/TCF Lymphoid enhancer-binding factor/T-cell factor LOXL proteins Lysyl oxidaseelike proteins MDCK cells MadineDarby Canine Kidney cells MET Mesenchymal-epithelial transition MMPs Matrix-metalloproteases MRTFs Myocardin-related transcription factors NES Nuclear export sequence PGE2 Prostaglandin E2 PI3K Phosphatidylinositol 3 kinase PMC Primary mesenchyme cells ROS Reactive oxygen species RTK Receptor tyrosine kinase TGF-b Transforming growth factor-b VEGF Vascular endothelial growth factor Zeb Zinc finger E-box-binding

Glossary Apical Surface of the epithelial layer where adherens junctions and tight junctions are located. This is opposite the basal surface. Basal Surface of the epithelial layer where the basal lamina is found. This is opposite the apical surface. Basal Lamina Consists of extracellular matrix components such as collagen type IV, fibronectin, and laminin. The basal lamina functions to stabilize the epithelium and is a barrier against migratory cells. Epithelial Epithelial cells adhere tightly to each other at their lateral surfaces and to an organized extracellular matrix at their basal domain, thereby producing a sheet of cells resting on a basal lamina with an apical surface. EpithelialeMesenchymal Transition The conversion of epithelial cells into mesenchymal cells. Mesenchymal Mesenchymal cells are individual cells with a bipolar morphology that are held together as a tissue within a three-dimensional extracellular matrix. MesenchymaleEpithelial Transition The conversion of mesenchymal cells into epithelial cells.

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[67] Strippoli R, Benedicto I, Perez Lozano ML, Cerezo A, Lopez-Cabrera M, del Pozo MA. Epithelial-to-mesenchymal transition of peritoneal mesothelial cells is regulated by an ERK/NF-kB/Snail1 pathway. Dis Model Mech 2008;1(4e5):264e74. [68] Lee JM, Dedhar S, Kalluri R, Thompson EW. The epithelial-mesenchymal transition: new insights in signaling, development, and disease. J Cell Biol 2006;172(7):973e81. [69] Thuault S, Tan EJ, Peinado H, Cano A, Heldin C-H, Moustakas A. HMGA2 and Smads co-regulate SNAIL1 expression during Induction of epithelial-to mesenchymal transition. J Biol Chem 2008;283(48):33437e46. [70] Morita T, Mayanagi T, Sobue K. Dual roles of myocardin-related transcription factors in epithelial mesenchymal transition via slug induction and actin remodeling. J Cell Biol 2007;179(5):1027e42. [71] Medjkane S, Perez-Sanchez C, Gaggioli C, Sahai E, Treisman R. Myocardin-related transcription factors and SRF are required for cytoskeletal dynamics and experimental metastasis. Nat Cell Biol 2009;11(3):257e68. Nature Publishing Group. [72] Fan L, Sebe A, Peterfi Z, Masszi A, Thirone ACP, Rotstein OD, et al. Cell contact-dependent regulation of epithelial-myofibroblast transition via the Rho-Rho kinase-phospho-myosin pathway. Mol Biol Cell 2007;18(3). E06e07e0602. [73] Fujita N, Jaye DL, Kajita M, Geigerman C, Moreno CS, Wade PA. MTA3, a Mi-2/NuRD complex subunit, regulates an invasive growth pathway in breast cancer. Cell 2003;113(2):207e19. [74] Wang Y, Zhang H, Chen Y, Sun Y, Yang F, Yu W, et al. LSD1 is a subunit of the NuRD complex and targets the metastasis programs in breast cancer. Cell 2009;138(4):660e72. [75] Langer EM, Feng Y, Zhaoyuan H, Rauscher III FJ, Kroll KL, Longmore GD. Ajuba LIM proteins are Snail/Slug corepressors required for neural crest development in Xenopus. Dev Cell 2008;14(3):424e36. [76] Nawshad A, Hay ED. 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Dual regulation of Snail by GSK-3ß-mediated phosphorylation in control of epithelialmesenchymal transition. Nat Cell Biol 2004;6(10):931e40. [81] Peinado H, Olmeda D, Cano A. Snail, Zeb and bHLH factors in tumour progression: an alliance against the epithelial phenotype? Nat Rev Cancer 2007;7(6):415e28. [82] Delcommenne M, Tan C, Gray V, Rue L, Woodgett J, Dedhar S. Phosphoinositide-3-OH kinase-dependent regulation of glycogen synthase kinase 3 and protein kinase B/AKT by the integrin-linked kinase. Proc Natl Acad Sci USA 1998;95(19):11211e6. [83] Wu Y, Evers BM, Zhou BP. Small C-terminal domain phosphatase enhances snail activity through dephosphorylation. J Biol Chem January 2, 2009;284(1):640e8. [84] Wang S-P, Wang W-L, Chang Y-L, Wu C-T, Chao Y-C, Kao S-H, et al. p53 controls cancer cell invasion by inducing the MDM2-mediated degradation of Slug. Nat Cell Biol 2009;11(6):694e704. Nature Publishing Group. [85] Yang Z, Rayala S, Nguyen D, Vadlamudi RK, Chen S, Kumar R. 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Snail promotes Wnt target gene expression and interacts with beta-catenin. Oncogene 2008; 27(37):5075e80. [105] Logan CY, Miller JR, Ferkowicz MJ, McClay DR. Nuclear beta-catenin is required to specify vegetal cell fates in the sea urchin embryo. Development 1999;126(2):345e57. [106] Huelsken J, Vogel R, Brinkmann V, Erdmann B, Birchmeier C, Birchmeier W. Requirement for beta-catenin in anterior-posterior axis formation in mice. J Cell Biol 2000;148(3):567e78. [107] Liebner S, Cattelino A, Gallini R, Rudini N, Iurlaro M, Piccolo S, et al. ß-Catenin is required for endothelial-mesenchymal transformation during heart cushion development in the mouse. J Cell Biol 2004;166(3):359e67. [108] Cowin P, Rowlands TM, Hatsell SJ. Cadherins and catenins in breast cancer. Curr Opin Cell Biol 2005;17(5):499e508. [109] Wong AST, Gumbiner BM. Adhesion-independent mechanism for suppression of tumor cell invasion by E-cadherin. J Cell Biol 2003;161(6): 1191e203. [110] Kim K, Pang KM, Evans M, Hay ED. Overexpression of ß-Catenin induces apoptosis independent of its transactivation function with LEF-1 or the involvement of major G1 cell cycle regulators. Mol Biol Cell 2000;11(10):3509e23. [111] Gat U, DasGupta R, Degenstein L, Fuchs E. De novo hair follicle morphogenesis and hair tumors in mice expressing a truncated ß-Catenin in skin. Cell 1998;95(5):605e14. [112] Woolf AS, Kolatsi-Joannou M, Hardman P, Andermarcher E, Moorby C, Fine LG, et al. Roles of hepatocyte growth factor/scatter factor and the met receptor in the early development of the metanephros. J Cell Biol 1995;128(1e2):171e84. [113] Grotegut S, von Schweinitz D, Christofori G, Lehembre F. Hepatocyte growth factor induces cell scattering through MAPK/Egr-1-mediated upregulation of Snail. EMBO J 2006;25(15):3534e45. [114] Ciruna B, Rossant J. FGF signaling regulates mesoderm cell fate specification and morphogenetic movement at the primitive streak. Dev Cell 2001;1(1):37e49. [115] Suyama K, Shapiro I, Guttman M, Hazan RB. A signaling pathway leading to metastasis is controlled by N-cadherin and the FGF receptor. Cancer Cell 2002;2(4):301e14. [116] Billottet C, Tuefferd M, Gentien D, Rapinat A, Thiery J-P, Broe¨t P, et al. Modulation of several waves of gene expression during FGF-1 induced epithelial-mesenchymal transition of carcinoma cells. J Cell Biochem 2008;104(3):826e39. [117] Lu Z, Ghosh S, Wang Z, Hunter T. Downregulation of caveolin-1 function by EGF leads to the loss of E-cadherin, increased transcriptional activity of ß-catenin, and enhanced tumor cell invasion. Cancer Cell 2003;4(6):499e515. [118] Lee M-Y, Chou C-Y, Tang M-J, Shen M-R. Epithelial-mesenchymal transition in cervical cancer: correlation with tumor progression, epidermal growth factor receptor overexpression, and Snail up-regulation. Clin Cancer Res 2008;14(15):4743e50.

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C H A P T E R

2 CelleExtracellular Matrix Interactions in Repair and Regeneration Melissa Petreaca1, Manuela Martins-Green2 1

DePauw University, Greencastle, IN, United States; 2University of California, Riverside, CA, United States

INTRODUCTION For many years, the extracellular matrix (ECM) was thought to serve only as a structural support for tissues. Several studies conducted in the latter half of the 20th century dispelled this notion, providing evidence that matrix molecules promote the conversion of myoblasts to myotubes and facilitate the morphogenesis of multiple glands and organs [1]. Data from these and other studies collectively implicated the ECM in embryonic inductions and suggested the presence of matrix binding sites on the surface of cells responding to specific matrix molecules. These ideas opened an entire field of inquiry investigating the detailed mechanisms through which ECM molecules influence cell behavior. Bissell et al. proposed the model of “dynamic reciprocity,” in which ECM molecules transmit signals across the cell membrane via cell surface receptors, stimulating signaling pathways that change expression of specific genes whose protein products then affect or alter the ECM [1]. It has become clear that this concept is correct, because celleECM interactions activate intracellular signaling, modulate cytokine and growth factor activities, and regulate cell adhesion, migration, growth, differentiation, and programmed cell death. Much of our current understanding of the molecular basis of celleECM interactions in these events comes from in vitro cell or organ cultures and in vivo experiments involving changes in the composition or function of matrix molecules. Here, we will first briefly discuss the composition and diversity of some of the better-known ECM molecules and their receptors, and then discuss selected examples that illustrate the dynamics of celleECM interactions during regenerative (scarless) and nonregenerative (scar-forming) wound healing, as well as the potential mechanisms through which matrix-induced signaling affects wound repair. Finally, we will discuss the implications of celleECM interactions in regenerative medicine.

COMPOSITION AND DIVERSITY OF THE EXTRACELLULAR MATRIX The ECM is a molecular complex with many components, including, but not limited to, collagens, hyaluronan, proteoglycans, glycosaminoglycans (GAGs), and elastins. These molecules interact with each other and with growth factors, cytokines, and matrix-degrading enzymes and their inhibitors. The distribution and organization of matrix molecules are not static, but rather vary from tissue to tissue and, during development and tissue repair after injury, from stage to stage, conferring distinct properties and functions on the tissue in question [1,2]. For example, mesenchymal cells are immersed in an interstitial matrix that confers specific biomechanical and functional properties to connective tissue, whereas epithelial and endothelial cells contact the specialized matrix of the basement membrane via only their basal surfaces, conferring mechanical strength and specific physiological properties to the epithelia [3]. The temporal and spatial presence and distribution of specific matrix molecules are critical for the function of both mature tissues and new tissues generated by development or repair, as shown by the impact of mutations in matrix molecules and their associated receptors on these processes [3]. This diversity of composition, organization, and distribution of ECM among different tissues and the same tissue under different conditions results from changes Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00002-3

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Copyright © 2019 Elsevier Inc. All rights reserved.

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2. CELLeEXTRACELLULAR MATRIX INTERACTIONS IN REPAIR AND REGENERATION

in gene expression, splicing, and posttranslational modifications of matrix molecules. For example, alternative splicing or proteolytic cleavage may change the binding potential of matrix proteins to their receptors or to other matrix molecules, whereas altered patterns of glycosylation affect cell adhesion and migration [4e6]. Changes in the structure, organization, and components of the ECM also can affect the distribution and function of growth factors and cytokines that interact with the ECM, influencing downstream signaling in multiple ways. Through its binding to growth factors and cytokines, matrix molecules can protect these molecules from degradation, regulate their local concentrations, facilitate their formation of stable gradients within a tissue, and/or present them more effectively to their receptors, all of which increase downstream signaling and thereby alter signalinginduced cell behaviors [4,7e10]. One example of this matrix-facilitated growth factor signaling involves the binding of vascular endothelial growth factor (VEGF) to fibronectin or heparin sulfate proteoglycans, which increases endothelial proliferation and migration [9]. However, matrix molecules can also bind and sequester growth factors, thereby preventing ligandereceptor interactions and downstream signaling. Binding of heparan sulfate proteoglycans (HSPGs) to heparin-binding endothelial growth factor (EGF)-like factor growth factor (HB-EGF) prevents receptor binding and downstream signaling until proteolytic release of the ligand from the HSPGs [11]. Further complicating the impact of matrix molecules on growth factor signaling, some matrix molecules and matricryptins, proteolytic fragments of matrix molecules, can bind directly to growth factor receptors and either activate or inhibit downstream signaling [4]. For example, specific domains of matrix molecules, including the epidermal growth factor (EGF)-like repeats of laminin-332 and tenascin C, bind and activate the epidermal growth factor receptor (EGFR) [12]. These EGF-like repeats may function as matricryptins after their proteolytic release from their molecules of origin, allowing EGF-like matrix-derived peptides to function as soluble ligands that activate the EGFR, or the EGF-like repeats may function as EGFR ligands in the context of the intact matrix molecule [8,12]. If these EGFlike repeats can induce signaling as part of intact matrix molecules, these repeats could function as persistent, stable inducers of EGFR signaling that could regulate cell function over long periods [8]. In contrast, some matrix molecules and many matricryptins inhibit growth factor receptors, suppressing or antagonizing their downstream effects on cell survival, proliferation, and migration. For example, the proteoglycan decorin binds multiple growth factor receptors and interferes with downstream signaling, whereas fragments of collagen XVIII (endostatin), collagen IV (tumstatin), and the proteoglycan perlecan (endorepellin) bind and inhibit or downregulate VEGF receptors, decreasing endothelial cell survival and migration [10,12].

RECEPTORS FOR EXTRACELLULAR MATRIX MOLECULES To establish the direct effects of ECM molecules on cell behavior, it was important to identify transmembrane receptors for the specific sequences present on these molecules. Early investigations of salivary gland morphogenesis showed that intracellular microfilaments contracted near the sites of glycosaminoglycan deposition, which suggested that the matrix could regulate microfilament function through the binding of matrix molecules with cell surface receptors [12a]. Later experiments demonstrated that ECM molecules contain specific amino acid motifs enabling their direct binding to cell surface receptors, the best-characterized of which is the tripeptide RGD, first found in fibronectin but later identified in many other matrix molecules [8]. The RGD motifs found in multiple matrix molecules serve as ligands for a subset of receptor proteins called integrins. Integrins, a family of heterodimeric transmembrane proteins composed of a and b subunits, were the first ECM receptors to be identified [13]. The 18 a and 8 b integrin subunits interact with each other in multiple combinations to generate a diverse family of matrix-binding receptors (Fig. 2.1). Some integrins have restricted ligand specificity; others bind multiple epitopes located on the same or different ECM molecules (Fig. 2.1) [13]. In contrast to the extracellular ligand-binding domains of the integrin heterodimers, the intracellular domains of these receptors are relatively small. Despite the relatively short length of their cytoplasmic domains, integrins interact with an array of intracellular signaling proteins that facilitate integrin-associated, matrix-induced signal transduction [14]. The simultaneous binding of some integrins, including integrins avb3 and a5b1, to both a matrix ligand and a growth factor receptor attached to its growth factor ligand facilitates growth factor receptor signaling, expanding the role of integrinematrix binding in the regulation of cell behavior [15]. Although not as extensively studied as the integrins, several proteoglycans, including members of the syndecan family, CD44, and RHAMM (receptor for hyaluronate-mediated motility) also function as ECM receptors [6]. The extracellular domains of syndecans interact with multiple ligands, including growth factors and matrix molecules

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RECEPTORS FOR EXTRACELLULAR MATRIX MOLECULES

β6

β8 FN

VN LAP-β1

FG

β5

FN

VN

Least Selective OSP

αv

FG

FN

TSP

VN

CO

β3

vWF

FG

FN VN

TSP

αllb

vWF

CO FN FG VN

α9

α1

vWF

P OS

LN

TN

α2

CO

α8 FN

CO

β1 LN

CO

LN

FN

FN

LN

α7

αIEL

α3

LN

FN E

C

1

A

D

H

AM

VC

FN

β4

LN

α6

β7

α4 α5

FN -1

AM

IC

Very Selective

αM

αH

αD

C3bi

FG

FX

-3

ICAM -2

-1

ICAM

αL

ICAM

β2 FN

C3bi

αX

RGD mediated binding

FIGURE 2.1 Representative members of the integrin family of extracellular matrix (ECM) receptors and their respective ligands. These heterodimeric receptors are composed of one a and one b subunit and are capable of binding a variety of ligands, including immunoglobulin superfamily cell adhesion molecules, complement factors, and clotting factors, in addition to ECM molecules. Cellecell adhesion is largely mediated through integrin heterodimers containing the b2 subunits, whereas cellematrix adhesion is mediated primarily via integrin heterodimers containing the b1 and b3 subunits. In general, the b1 integrins interact with ligands found in the connective tissue matrix, including laminin, fibronectin, and collagen, whereas the b3 integrins interact with vascular ligands, including thrombospondin, vitronectin, fibrinogen, and von Willebrand factor. CO, collagens; C3bi, complement component; FG, fibrinogen; FN, fibronectin; FX, Factor X; ICAM-1, intercellular adhesion molecule-1; ICAM-2, intercellular adhesion molecule-2; ICAM-3, intercellular adhesion molecule-3; LN, laminin; OSP, osteopontin; TN, tenascin; TSP, thrombospondin; VCAM-1, vascular cell adhesion molecule-1; VN, vitronectin; vWF, von Willebrand factor.

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such as fibronectin and multiple collagens, via chondroitin- and heparan-sulfate GAGs, whose composition varies based on the specific syndecan family member and the tissue in which it is expressed [16]. In contrast to the variable extracellular domain, the syndecan transmembrane and intracellular domains are small and relatively conserved, interacting with the actin cytoskeleton and several associated signaling molecules, including kinases in the Src and PKC families as well as Tiam1, a guanine nucleotide exchange factor for and activator of the Rho GTPases [17]. Although these intracellular signaling molecules are likely responsible for syndecan-mediated cytoskeletal reorganization, syndecan binding to nonmatrix molecules, including growth factors, growth factor receptors, and integrins, makes the identification of matrix-specific, syndecan-mediated signal transduction challenging [16]. Like syndecans, the CD44 receptor carries chondroitin sulfate and heparan sulfate chains on its extracellular domain and undergoes tissue-specific splicing and glycosylation to yield multiple isoforms [6]. Although hyaluronan is its primary ligand, CD44 interacts with other matrix molecules, including fibronectin, laminin, collagen IV, and collagen XIV. Furthermore, the ability of the heparan- and chondroitin-sulfate GAGs on CD44 to bind growth factors, combined with the interactions of CD44 with growth factor receptors, such as EGFR and transforming growth factor-bR (TGFbR), suggests a role for CD44 in modulating growth factor signaling [6,18]. In contrast to the transmembrane CD44 and syndecan proteoglycans, RHAMM, another proteoglycan able to both bind matrix molecules and induce signaling, is associated with the cell membrane through a glycosylphosphatidylinositol (GPI) linkage and not a transmembrane domain. As such, RHAMM located on the cell surface likely activates intracellular signaling through indirect mechanisms, via interactions with transmembrane receptors such as CD44 or growth factor receptors [19]. Interestingly, a variant of RHAMM lacking the GPI anchor resides in the cytoplasm and/or nucleus, where it affects intracellular signaling and cytoskeletal organization through its binding to intracellular signaling molecules and the cytoskeleton, expanding the role of RHAMM as a regulator of cell signaling [19]. Cell surface receptors other than integrins or proteoglycans, including the elastin receptor complex (ERC), CD36, annexin II, Toll-like receptors, and discoidin domain receptors (DDRs) can also serve as receptors for ECM molecules. The ERC is a complex of proteins, including elastin-binding protein (EBP), a splice variant of b-galactosidase, as well as neuraminidase 1 and cathepsin A, that serves as a receptor for elastin, laminin, fibrillin, and peptides derived from these ECM molecules. Signaling activated by this receptor is necessary for elastin deposition and participates in signaling induced by elastin and laminin during mechanotransduction [20]. Another nonintegrin, nonproteoglycan receptor, CD36, better known for its function as a scavenger receptor for long-chain fatty acids and oxidized low-density lipoprotein, binds thrombospondin, collagen I, and collagen IV [21]. CD36-thrombospondin binding activates a variety of signal transduction molecules, ultimately leading to inhibition of angiogenesis via increased endothelial cell apoptosis [22]. Cell surface annexin II, yet another matrix receptor, interacts with alternative splice variants of tenascin-C and facilitates proliferative and migratory effects of these tenascin C splice variants [23]. Although not a matrix receptor per se, Toll-like receptors can function as receptors for fragments of multiple matrix molecules, including fibronectin and lowemolecular weight fragments of hyaluronan, which suggests that such fragments may function as danger signals that induce inflammation in response to matrix degradation [18,24]. Receptor tyrosine kinases can also serve as receptors for matrix molecules, as described earlier for EGFR, which binds matrix molecules through their EGF-like domains, as well as the collagen-binding discoidin domain receptors DDR1 and DDR2. Whereas the DDR proteins, like all receptor tyrosine kinases, promote receptor phosphorylation in response to ligand binding, DDRs form constitutive dimers required for ligand binding, unlike most of this receptor family, which dimerize after ligand binding [25]. Also, in contrast to other receptor tyrosine kinase activation, DDR activation by their collagen ligands induces prolonged rather than transient signaling events [26]. The combination of long-lived collagen ligands and the relatively long-term signaling induced by these ligands through the DDRs could provide sustained regulation of cell survival, proliferation, and/or migration in cells expressing these receptors. In the next section, we will focus on ways in which ECM molecules and their various receptors affect signal transduction, the multiple cell behaviors regulated by this signaling, and the impact of this signaling on wound healing under both nonregenerative and regenerative conditions.

SIGNAL TRANSDUCTION EVENTS DURING CELLeEXTRACELLULAR MATRIX INTERACTIONS As a result of interactions between ECM molecules and their receptors, as described previously, signals can be transmitted directly or indirectly to signaling molecules within the cell, activating a cascade of events and the coordinated expression of a variety of genes involved in cell adhesion, migration, proliferation, differentiation, and

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Recruitment of adaptor proteins Activation of Signal Transduction Cascades

Changes in Cell Adhesion, Migration, Proliferation, Apoptosis

α/β integrins

Growth factor receptor

Reparative and Regenerative Tissue Responses Non-integrin ECM receptor

Collagen

Fibronectin

Laminin

FIGURE 2.2 Schematic diagram of celleextracellular matrix (ECM) interactions present during healing and regenerative responses. Such

interactions between the ECM receptors and their respective ligands initiate signal transduction cascades culminating in a variety of cellular events important in repair and regeneration, including changes in cellular adhesion and migration and altered rates of proliferation and apoptosis. The presence and/or extent of such changes may influence the balance of repair and regenerative responses to favor one outcome over another; thus, interventions that alter ECM signaling events may shift this balance to favor tissue regeneration and decrease scarring.

programmed cell death (Fig. 2.2). Some of the increasing evidence that celleECM interactions through multiple types of matrix receptors activate a variety of signaling pathways affecting these specific cell outcomes is described subsequently.

Adhesion and Migration When considering the importance of cellematrix interactions in adhesion and migration, it is important to recognize that some receptors for ECM molecules, such as the integrins, can participate in both traditional “outside-in” signaling, leading to the activation of intracellular signaling, and “inside-out” signaling, in which intracellular signaling activates the receptor by increasing its affinity for an ECM molecule. This is further complicated by the fact that receptor activation and ligand binding can also initiate outside-in signaling. In the following section, the signaling events refer to outside-in signaling unless otherwise specified. The outside-in signaling activated by integrin binding to its ligand requires the indirect interaction of the integrin cytoplasmic domain with the cytoskeleton, through a variable group of proteins that collectively comprise the integrin adhesion complex [14]. Some of the proteins of the integrin adhesion complex facilitate integrin-mediated mechanotransduction, whereas other proteins connect the integrins to other types of downstream signal transduction pathways [14]. Integrin-mediated ECM signaling uses these associated proteins to induce changes in cell shape and lead to proliferation, migration, and/or differentiation [27]. Some of the best-known components of the integrin adhesion complex include the signaling adaptor proteins focal adhesion kinase (FAK) and paxillin, the mechanotransducing proteins vinculin and talin, the actin regulatory proteins zyxin and VASP, and the actin-binding protein a-actinin [14]. When the integrin heterodimer interacts with its matrix ligand, FAK, which is associated with the

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membrane-proximal portion of the integrin cytoplasmic domain, becomes autophosphorylated on tyrosine 397 (Y387) [28]. The Y397 autophosphorylation may be associated with mechanosensation, as shown by the positive correlation between Y397 phosphorylation and adhesion to fibronectin of increasing stiffness [29]. However, the role of mechanosensation in FAK Y397 phosphorylation may depend on both the matrix molecule and the integrin in question, because soluble collagen but not soluble fibronectin can induce FAK autophosphorylation in suspended cells [29]. Upon integrinematrix binding and FAK activation, the phosphorylated Y397 serves as a binding site for the SH2 domain of the nonreceptor tyrosine kinase c-Src, which then enhances FAK activity by phosphorylating additional tyrosine residues [28,30]. Src-mediated FAK phosphorylation on tyrosine 925 generates a binding site for the Grb2/Sos complex, with subsequent activation of Ras and the mitogen-activated protein kinase (MAPK) cascade, which may be involved in adhesion/deadhesion and migration [28,31]. In addition to activation of the Ras/MAPK pathway, the FAK/Src complex phosphorylates and regulates the activity of other components of the integrin adhesion complexes, including paxillin, p130Cas, vinculin, and tensin [32]. All of these proteins are associated with cell migration, as shown by the migration defects observed in cells either lacking these molecules altogether and/or expressing mutants affecting their activity or localization [33e36]. For example, paxillin-deficient fibroblasts exhibit both reduced phosphorylation of signaling molecules downstream of integrin ligation and decreased cell motility [33]. The phosphorylation of paxillin by the FAK/Src complex promotes paxillin binding to the SH2 domain of the Crk/DOCK180/ELMO complex. DOCK180, a guanine nucleotide exchange factor (GEF) for Rac1, then activates Rac1 and promotes cell migration [37]. Similarly, p130Cas phosphorylation by the FAK/Src complex promotes Rac1 activation and migration through Crk/DOCK180/ELMO-mediated Rac activation [28]. Beyond their activation of Rac, phosphorylated integrin adhesion complex components regulate other members of the Rho GTPase family. For example, phospho-paxillin also appears to activate p190RhoGAP, leading to localized inhibition of RhoA activity [37]. Additional regulators of RhoA and Rac activity in migrating cells include members of the tensin family, which, like paxillin and p130Cas, are phosphorylated by FAK/Src complexes. In contrast to the RhoGAP-activating function of phosphorylated paxillin, however, tensin1 increases RhoA activity by inhibiting the RhoGAP activity of DLC1 [34]. The conventional understanding of Rac and RhoA dynamics during cell migration suggests that the combination of enhanced Rac1 activity and decreased RhoA activity at the leading edge of migrating cells decreases cell adhesion and promotes protrusion formation, thereby facilitating cell migration. However, Fo¨rster resonance energy transferebased sensors demonstrate both RhoA and Rac activity at the leading edge of migrating cells, suggesting a more complex relationship between these GTPases and membrane protrusion than previously thought and demonstrating the need for more research in this area [38]. Although FAK and c-Src are best known for their roles in outside-in signaling, as described earlier, these kinases are also involved in inside-out signaling, which promotes integrineligand binding. FAK promotes integrin activation, cell adhesion to fibronectin, and strengthening of focal adhesions [39]. These effects appear to require Src binding and/or activity, because a Y397 F mutation that prevents FAK autophosphorylation and Src binding at this site also prevents FAK-mediated adhesion [39]. FAK-induced integrin binding to ECM molecules can then initiate outside-in signaling, leading to more FAK activation, FAK-Src interaction, and downstream signaling that promotes deadhesion and migration. This suggests a cycle of FAK and Src activity, in which they initially promote deadhesion and migration, followed by the formation of new adhesions at the leading edge. In support of this FAK/Src cycle of activity, active Src moves from the focal adhesions to the membrane ruffles at the leading edge during cell migration [39a]. The activation of integrins downstream of FAK/Src signaling may mediated by talins and kindlins, two families of proteins that bind directly to the cytoplasmic domains of b integrins. FAK is necessary for talin recruitment to integrins at newly forming adhesion sites, and cells either lacking specific talin family members or expressing mutants that abolish talineintegrin interactions prevent integrin activation induced by various stimuli, leading to defects in integrin-mediated adhesion [40]. Deficiencies in kindlin expression are associated with similar defects in integrin activation and cell adhesion, which suggests the possibility that talins and kindlins work together, potentially downstream of FAK, to activate integrins and promote cell adhesion [41]. However, kindlin signaling is not limited to inside-out signaling; it also participates in outside-in signaling after integrin engagement with matrix ligands. The interaction of kindlin-2 with Src is not required for inside-out signaling and cell adhesion to fibronectin but is necessary for paxillin phosphorylation downstream of integrin ligation and for platelet-derived growth factor (PDGF)-induced mesangial cell proliferation and migration [42]. These findings suggest that kindlin-2 may have multiple roles in integrin-mediated signaling, both promoting integrin-mediated adhesion and facilitating integrin-induced signaling after matrix adhesion, underscoring the interplay between inside-out and outside-in integrin signaling. A separate form of integrin activation occurs downstream of growth factor receptors. VEGF binding to the VEGFR2 in endothelial cells activates c-Src, which directly phosphorylates b3 integrin, increasing endothelial cell

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adhesion to the avb3 ligand vitronectin [43]. Interestingly, Src is also necessary for integrin avb3 interaction with VEGFR2, which is also required for maximal VEGFR2 phosphorylation and VEGF-induced endothelial cell migration in vitro and angiogenesis in vivo [43]. Furthermore, VEGF-induced endothelial cell migration requires FAK activity [44]. These data suggest that the interactions between VEGFR2 and integrin avb3 couple integrin activation with VEGFR2 signaling, promoting FAK activation and signaling downstream of FAK (described earlier), thereby linking cellematrix adhesion with promigratory signaling. Beyond integrin signaling in cell adhesion and migration, nonintegrin ECM receptors, including proteoglycan receptors such as syndecans, CD44, and RHAMM, as well as the elastin-laminin receptor, the EGFR, and DDR1, also participate in cell adhesion and migration through a variety of mechanisms. For example, syndecans facilitate migration through both activation of signaling through the cytoplasmic domains of the syndecans themselves and syndecan-mediated activation of growth factor receptors and integrins [17]. The PDZ-binding motif, located on syndecan intracellular domains, binds several signaling proteins with PDZ domains, including the Rac1 GEF Tiam1, which suggests a mechanism by which syndecan binding to matrix ligands directly regulates cell migration [17]. Syndecans also collaborate with integrins and growth factor receptors in the induction of cell adhesion and migration. In the case of integrins, syndecan-4 regulates the formation and generation of specific integrin adhesion complexes by promoting the internalization and degradation of some integrin heterodimers while increasing surface levels of different heterodimers [16,45]. These alterations in integrin composition induced by syndecan-4 may affect both cell adhesion and integrin-induced signaling important in cell migration. In support of this idea, syndecan cooperation with integrin heterodimers mediates cell adhesion to vitronectin and laminin and induces cell migration, and syndecan-4 promotes Rac1 activation and localization to the leading edge of migrating cells, facilitating directional cell migration in response to fibronectin [46]. Syndecans also participate in growth factor receptor signaling through a complex web of interactions in which the syndecans bind growth factors and prevent receptor binding until proteolytic release (e.g., HB-EGF) or bind both growth factor and its receptor to promote receptor activation (e.g., fibroblast growth factor [FGF] and VEGF) [16]. These growth factoreproteoglycan interactions may provide a stable growth factor gradient that facilitates directional cell migration [8,9]. Two non-syndecan proteoglycan receptors, CD44 and RHAMM, both bind hyaluronan and collaborate to promote hyaluronan-induced signaling. Hyaluronan is a matrix molecule that is generally produced in a highe molecular weight form that can cross-link its cell surface receptors, promoting cell adhesion and inhibiting cell migration [19]. However, loweremolecular weight hyaluronan fragments found in injured and/or inflamed tissues promote cell migration associated with inflammation and angiogenesis [47]. Some of these differences in function between highe and lowemolecular weight hyaluronan likely result from differences in receptor selection, because lowemolecular weight forms are more likely to bind Toll-like receptors, which can then promote inflammation through nuclear factor kBeinduced expression of cytokines and chemokines [47]. However, the angiogenesis induced by hyaluronan fragments requires CD44, which suggests that hyaluronan fragments of different sizes may induce different conformational changes in CD44 or selectively bind different CD44 splice or glycosylation variants, in either case inducing different downstream signaling pathways that either promote angiostatic or angiogenic outcomes [47,48]. Some of these CD44 variants can interact with growth factors and/or their receptors, modulating downstream signaling. For example, the splice variant CD44v6 binds both to VEGF isoforms and to VEGFR2, and CD44v6 inhibition decreases VEGF-induced activation of VEGFR2 and downstream phosphorylation of ERK in endothelial cells [49]. Furthermore, CD44v6 inhibition significantly decreases VEGF-induced endothelial cell migration in vitro and angiogenesis in vivo, which underscores the importance of this CD44 variant in promigratory VEGFR2 signaling [49]. Fragments of matrix components, likely generated under proteolytic conditions such as those occurring during wound healing and inflammation, bind multiple matrix receptors in addition to the receptors described earlier, including the EGFR and the elastin receptor, thereby promoting cell adhesion and migration in various cell types. The EGF-like repeats in tenascin-C, laminin 332, thrombospondin, and secreted protein acidic and rich in cysteine protein (SPARC) can decrease cellematrix adhesions and promote cell migration [12]. Although receptor binding and signaling induced by these matrix molecules are complicated by their ability to bind and activate integrins, there is compelling evidence that the EGF-like repeats of tenascin C and laminin 332 promote changes in cell adhesion and motility in an EGFR-dependent manner [12]. Another matrix fragment-binding receptor, the ERC, binds fragments of elastin, laminin, and fibrillin. Although ERC binding to ligands initially promotes cell adhesion by promoting elastin deposition into the ECM, interaction of the EBP component of the ERC with elastin-derived peptides can promote migration of multiple cell types, including monocytes, keratinocytes, fibroblasts, smooth muscle cells, and endothelial cells [50]. The proliferative and migratory effects of elastin-derived peptides may result from the activation of multiple MAPK cascades downstream of the ERC [51].

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In contrast to the ERC and EGFR, which can bind matrix fragments, DDR1 and DDR2 bind intact collagen molecules rather than fragments [25]. Signaling induced by collagen through the DDRs seem to have cell typeespecific effects, promoting migration in some cell types while repressing it in others [25]. In the case of DDR2, collagen binding promotes the movement of fibroblasts through matrix, likely through induction of matrix metalloproteinase (MMP)-2 expression and activation, suggests a role in tissue invasion [26]. Because collagen induces relatively long-term DDR2 activation, DDR2-induced cell behaviors, such as matrix invasion, could be maintained over substantial periods [26].

Proliferation and Survival ECM interaction with its receptors can promote cell proliferation and survival, as is demonstrated clearly by the anchorage dependence of cell growth. Even in the presence of growth factors, cells will not enter the S phase of the cell cycle without matrix adhesion [52]. In addition, cell detachment from the matrix or expression of matrix receptors in the absence of their intact ligands often promotes apoptosis, a process known as anoikis [15]. Thus, adhesion of cells to ECM molecules has an important role in regulating cell survival and proliferation. Because much of our understanding of matrix-induced proliferation and survival is related to matrix interaction with integrins, we will begin with integrin-associated proliferation and survival. Multiple studies in which integrins are either inhibited or deficient demonstrate that integrin signaling is critical for cell proliferation [52]. For example, the fibroblasts of mice lacking the a1b1 integrin, a primary collagen receptor, have reduced proliferation even though they exhibit normal adhesion [53]. The loss of integrin-mediated adhesion inhibits cell survival by inducing the movement of the proapoptotic protein Bax from the cytoplasm to the mitochondria, promoting apoptosis [54]. Integrin-mediated, adhesion-induced cell survival requires FAK activity, as shown by the Bax translocation and apoptosis of cells expressing dominant negative FAK and the survival of detached cancer cells overexpressing FAK [28,54]. Integrinematrix binding activates the FAKeSrc complex, which interacts with and activates PI3K, leading to the downstream activation of Akt [52]. Akt then alters the ratios of proapoptotic and antiapoptotic Bcl-2 family members, increasing the levels of the antiapoptotic Bcl-xl and Mcl-1 and decreasing the levels of proapoptotic Bax and Bak, thereby promoting cell survival [55]. Whereas the prosurvival signals downstream of Fak in epithelial cells require Src activity, as described previously, prosurvival signaling in fibroblasts instead involves p130Cas activation, which suggests that the mechanisms involved in FAK-induced survival are cell type specific [56]. Although maintenance of cell survival by integrin-mediated matrix adhesion is necessary for cell division, it is not sufficient. Many signaling pathways induced by integrin-ECM binding, notably the activation of MAPK pathways, promote cell division rather than simply survival [52]. As mentioned in an earlier section, FAK-induced p130Cas activation is upstream of Rac1, and once Rac1 is activated by this pathway, it can promote cell proliferation downstream of multiple effectors [28]. For example, Rac1-induced JNK activity promotes cell division through c-Jun-induced cyclin expression, whereas Rac1-mediated activation of Pak1 induces proliferation through its phosphorylation of Raf and downstream activation of extracellular signal-regulated kinase (ERK) [52]. Integrin ligation also activates ERK1/2 through FAK-mediated recruitment of Shc, an adaptor protein that binds Grb2/Sos and induces the Ras/ERK cascade [57]. Furthermore, cell-ECM binding and Rac activation promote the degradation of cyclin-dependent kinase (CDK) inhibitors, which would otherwise block cell proliferation [52]. Matrix molecules can also induce cell proliferation through cooperation with growth factor receptor signaling [8]. Such cooperative effects may occur in a direct manner, because some matrix molecules bind and activate growth factor receptors, leading to cell proliferation, whereas other matrix molecules bind growth factors and regulate their interaction with their receptors [23]. The matrix molecules that bind growth factor receptors directly can promote receptor activation and downstream signaling. For example, laminin and tenascin-C can bind and activate EGFR and associated downstream signaling through their EGF-like domains [12], which suggests the possibility that they or their proteolytic fragments may promote cell division downstream of EGFR. In contrast, decorin binding to EGFR and VEGFR2 inhibits proliferation induced by VEGF and EGF, potentially through internalization or degradation of the receptors, which has been shown for EGFR [12,58]. Although there is some evidence, described earlier, that matrixegrowth factor receptor engagement can promote cell proliferation, substantially more evidence demonstrates the importance of growth factor binding to matrix molecules on growth factor receptor activation and induction of cell proliferation. Multiple ECM molecules are able to bind to either growth factors or their receptors to regulate their activity [8]. For example, VEGF binding to vitronectin promotes the formation of a VEGFR2/integrin avb3 complex that greatly increases VEGF-induced VEGFR2

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phosphorylation and endothelial cell proliferation, and the binding of VEGF to fibronectin has a similar effect on VEGFR2/integrin a5b1 complex formation, VEGFR2 activation, and cell proliferation [43,59]. Several growth factors, including FGFs, VEGFs, PDGFs, and TGFb, can interact with HSPGs, which can either sequester these factors within the matrix, such that they are released upon matrix degradation, or can present them to their receptors [16]. For example, FGF binding to heparan sulfate moieties on syndecans facilitates FGF binding to and activation of FGFR, increasing the duration of signaling downstream of FGFR and promoting cell proliferation [16,45]. Similarly, VEGF binding to HSPGs increases binding to VEGFR2 and promotes cell proliferation [45]. In addition to matrix binding to growth factor receptors and growth factor binding to matrix, matrixegrowth factor cooperation can occur through direct interactions between integrins and growth factors or growth factor receptors. VEGF binds directly to integrin a9b1, which is necessary for VEGF-induced angiogenesis in vivo [59]. Other growth factors are also able to bind integrins directly; for example, insulin-like growth factor (IGF)-1 and FGF-2 both interact with integrin avb3, which is necessary for the cell proliferation induced by both factors [57]. Growth factor receptor binding to integrins may also enhance growth factoreinduced signaling. Both PDGFRb and VEGFR2 physically interact with integrin subunits, and concomitant integrin-mediated cell adhesion further increases both receptor activation and mitogenicity [52,59]. In the case of VEGFR2, integrin binding participates in FAK-induced Ras/ERK activation in endothelial cells, providing a likely mechanism for the role of integrin binding in proliferation [60]. Nonintegrin ECM receptors, including the hyaluronan receptor CD44, the ERC, and the collagen-binding DDRs, have also been implicated in cell proliferation and survival. Whereas highemolecular weight hyaluronan inhibits cell proliferation, the loweremolecular weight forms present in sites of injury and inflammation induce proliferation of multiple cell types, including fibroblasts and smooth muscle cells [19,47]. Hyaluronan-induced proliferation in fibroblasts appears to be mediated, at least in part, by CD44 and its downstream activation of ERK and Akt [61]. Similarly, binding of hyaluronan fragments to CD44 promotes smooth muscle cell proliferation via downstream activation of the RaseERK pathway [47]. The binding of elastin-derived peptides to the ERC promotes smooth muscle cell proliferation by activating multiple signaling pathways that culminate in activation of the MAPK cascade and upregulation of multiple cyclins and CDKs [62]. The ERC may also inhibit proliferation induced by growth factor receptors, because the neuraminidase subunit of the ERC desialylates cell surface growth factor receptors, preventing growth factorereceptor binding and downstream signaling, complicating the role that the ERC has in regulating proliferation [63]. Finally, both DDR2, a collagen receptor, and annexin II, a tenascin C receptor, can promote division in some cell types [23,26]. To date, we have few clues regarding signaling activated in matrix-induced proliferation via nonintegrin receptors, which suggests the need for more research in this area.

Differentiation Interaction of cells with ECM molecules, hormones, and growth factors is required to activate genes needed for the differentiation of multiple cell types, including keratinocytes, endothelial cells, and fibroblasts. Differentiation of keratinocytes is carefully regulated by multiple cellematrix and cellecell interactions. Keratinocytes in contact with the basement membrane proliferate and do not express proteins associated with terminal differentiation, whereas those cells that are not attached to the basement membrane in suprabasal layers begin to express these markers of terminal differentiation. This suggests that keratinocyte differentiation is repressed by adhesion to the basement membrane. In support of this idea, integrin activation in cultured cells inhibits differentiation, whereas impaired adhesion promotes differentiation [64]. In contrast, keratinocytes of mice deficient in various integrins can still differentiate, which suggests that there may be some redundancy in the system [65]. There may be additional cues provided by matrix molecules present within the epidermis itself that regulate differentiation. For example, hyaluronan is present within the epidermis, where it facilitates keratinocyte terminal differentiation in a CD44dependent manner [66]. These results are contradicted by separate studies suggesting that epidermal hyaluronan prevents premature keratinocyte differentiation [18]. One potential explanation for this discrepancy is that enzymatic digestion of hyaluronan may have generated hyaluronan fragments with binding and/or signaling differences from intact hyaluronan. Regardless, it appears that hyaluronan and CD44 are involved in the keratinocyte differentiation process. Other differentiated phenotypes also require integrin-mediated signaling events. TGFb-mediated differentiation of fibroblasts into myofibroblasts, contractile cells that produce ECM, requires cell integrin-mediated adhesion to extra domain A (EDA) of fibronectin [27]. Integrins avb5 and avb6 and integrin heterodimers containing the b1 subunit are associated with TGFb activation and are thus important in myofibroblast differentiation

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[27,65]. Because integrin a5b1 interacts with the EDA of fibronectin and integrin b1-deficient mice exhibit defective myofibroblast differentiation, the connection between integrin a5b1 and the fibronectin EDA may be particularly important in myofibroblast differentiation [27]. Pathways downstream of integrinefibronectin interaction, including FAK and integrin-linked kinase activation, are necessary for TGFb-induced myofibroblast differentiation, providing some potential mechanisms for its observed dependence on EDA-containing fibronectin [27]. Myofibroblast differentiation is also associated with matrix stiffness; increasing matrix stiffness is associated with increasing differentiation, which suggests a role for integrin-mediated mechanotransduction, likely involving FAK activation, in this differentiation process [2,5]. However, the roles of hyaluronan and CD44 activation in myofibroblast differentiation remain unclear owing to conflicting evidence in which both decreased hyaluronan production and hyaluronan-induced CD44 activation increase myofibroblast differentiation [18]. This discrepancy may result from differences in the ability of CD44 to engage in cross-talk with the TGFb receptor under different circumstances [18]. Differentiation of endothelial cells is more difficult to identify than is differentiation of keratinocytes or myofibroblasts, both of which have specific protein markers of differentiation. In contrast, endothelial cell differentiation occurs when endothelial cells form mature vascular structures, which is challenging to mimic in vitro. Early experiments involving cultured endothelial cells demonstrated growth of endothelial cells on an ECM comparable to the basement membrane, Matrigel, induced the formation of capillary-like tubes whose formation could serve as an in vitro proxy for endothelial cell differentiation [67]. These results suggested that the adhesion of endothelial cells with some component or components of Matrigel could promote their differentiation. The main matrix components of both Matrigel and basement membranes in blood vessels in vivo are laminins, heterotrimeric proteins composed of a, b, and g subunits [68]. In Matrigel, the primary laminin is laminin-111 (a1b1g1), which, in the absence of other matrix components, can promote endothelial tube formation in vitro [69]. The basement membranes of intact vessels in vivo contain laminin-411 (a4b1g1) primarily, rather than laminin-111, and laminin-411 participates in endothelial tube formation in vitro and angiogenesis in vivo; the latter is shown by the formation of abnormal vessels in laminin a4- and g1-deficient mice [68]. Peptides derived from laminins a1 and g1 promote angiogenesis through integrins avb3 and a5b1, which suggests the potential role of laminin degradation in new vessel formation [68]. After the formation of these nascent vessels, they must be stabilized and matured by the deposition of new basement membrane and the recruitment and differentiation of pericytes, smooth muscle cells that increase endothelial barrier function and participate in depositing a new basement membrane [9,70]. Fibronectin, laminin-411, collagen IV, and HSPGs are all required for the formation of the new basement membrane, whereas integrin-mediated interactions between the developing matrix and pericytes is necessary for continued deposition of basement membrane matrix molecules [68,71]. As such, matrixeintegrin interactions are critical for the deposition of the basement membrane and recruitment of pericytes, which, in turn, are necessary for new vessel maturation.

Apoptosis Many cell types undergo apoptosis, or programmed cell death, through well-known signal transduction pathways involving the activation of proteases from the caspase family. Cell-matrix signaling tends to promote cell adhesion and survival, repressing apoptosis. However, even when cells remain attached to some ECM molecules, the inability of some integrins to bind their matrix ligands induces a specific form of apoptosis called integrinmediated death through integrin-mediated recruitment and activation of caspase 8 [72]. In addition to promoting apoptotic signaling, the activation of caspase 3 downstream of caspase 8 may prevent prosurvival integrin signaling from separate matrix-bound integrins by cleaving paxillin and kindlin-3, blocking prosurvival outside-in signaling downstream of these proteins [37,41]. Integrin ligation by soluble, rather than intact, ligands also promotes apoptosis through the recruitment and activation of caspase 8 by the clustered integrins [73]. Such soluble ligands may be created by matrix degradation during tissue remodeling. A fragment of collagen XVIII, endostatin, binds to a5b1 integrin and decreases the expression of Bcl-xl and Bcl-2, prosurvival Bcl family members, promoting endothelial cell apoptosis [10]. Similarly, a fragment of collagen IV, tumstatin, induces apoptosis in endothelial cells via integrin avb3 [74]. Fragments of matrix molecules can also induce apoptosis through nonintegrin receptors. Lowemolecular weight fragments of hyaluronan can induce apoptosis in some cells through CD44, likely by interfering with prosurvival signaling downstream of CD44 [19]. Similarly, elastin-derived fragments promote fibroblast and lymphocyte apoptosis through the ERC [22,51]. Taken together, these pieces of evidence suggest that disrupting normal matrixe receptor adhesions can promote apoptosis in various cell types.

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CELLeEXTRACELLULAR MATRIX INTERACTIONS DURING HEALING OF CUTANEOUS WOUNDS Interactions of cells with ECM molecules have a crucial role during wound healing and regeneration. Continuous cross-talk between cells and the surrounding matrix environment contributes to the processes of clot formation, inflammation, granulation tissue development, and remodeling; during regeneration, matrix interactions are important in restoring damaged tissue. Many lines of experimental evidence demonstrate that the basic cellular mechanisms resulting in wound healing involve cell adhesionedeadhesion, migration, proliferation, differentiation, and apoptosis (Fig. 2.2).

Adhesion and Migration Shortly after tissue damage and during the early stages of wound healing, multiple factors and blood cells enter into the wound area, activating the coagulation cascade when coagulation factors from the blood encounter tissue factor expressed on endothelial cells, tissue factor expressed by nonvascular cells exposed by injury, or collagen, also exposed by injury [75]. This cascade ultimately results in the activation of thrombin, an enzyme that cleaves fibrinogen to generate fibrin, which polymerizes to form a fibrin clot. Injury to the endothelium simultaneously promotes the adhesion of platelets to subendothelial von Willebrand Factor and ECM components, causing platelet aggregation, activation, and adhesion to fibrin, trapping them within the fibrin clot [75]. Activated platelets release additional coagulation factors that promote and stabilize the fibrin clot, which serves as a vascular plug. In addition, to fibrin and platelets, the fibrin clot contains matrix components including plasma fibronectin and a variety of chemokines, cytokines, and growth factors released by activated platelets [7,76]. As such, in addition to its hemostatic function, the fibrin clot facilitates wound healing by serving as a provisional matrix for cell migration and a reservoir of cytokines, thrombin, and growth factors that collectively promote the later phases of inflammation and granulation tissue formation [77]. During the clotting process, platelets and activated mast cells degranulate, releasing vasodilating and chemotactic factors that chemoattract inflammatory leukocytes to the wound site, initiating the inflammatory response. Leukocyte extravasation from blood vessels requires several adhesion and signaling events, including the binding of leukocyte integrins with endothelial intercellular cell adhesion molecules and vascular cell adhesion molecules rather than matrix molecules, and the binding of leukocyte chemokine receptors with chemokines associated with endothelial cell surface HSPGs [27,78]. Some of the first matrix molecules that the leukocytes encounter during emigration from the bloodstream are located within the endothelial basement membrane and the provisional matrix. Neutrophils adhere to and migrate within the basement membrane as they move through the blood vessel, where they encounter laminins-411 and -511, fibronectin, and vitronectin, and subsequently interact with fibrin and fibronectin in the provisional matrix [76,79]. Neutrophils also secrete laminin-411, which likely participates in their extravasation through integrin aMb2, because inhibition of this integrin blocks leukocyte extravasation [79]. After leukocyte extravasation from the vasculature, the leukocytes are directed to the site of injury by the chemokines that form relatively stable gradients through interactions with endothelial cell surface proteins and ECM molecules, thereby promoting directional cell migration through the provisional matrix [6]. The fibrinefibronectin provisional matrix serves as substrate for the migration of leukocytes and later keratinocytes during the early stages of healing when inflammation and reepithelialization occur. Leukocyte interactions with ECM molecules via integrin receptors affect many of the functions of these cells, in particular those associated with cell adhesion, migration, the production of inflammatory mediators, and antimicrobial functions. As mentioned, neutrophils interact with fibronectin in the basement membrane and the provisional matrix, and several types of inflammatory cells interact with fibrinogen, the primary component of the provisional matrix, through integrins aMb2 and integrin aXb2 [77,80]. Neutrophil binding to thrombospondin 4, another ligand of integrin aMb2, also induces expression of the chemokine CXCL8, whereas monocyte interactions with fibrin in the provisional matrix induces expression of multiple proinflammatory cytokines and chemokines, including TNFa, interleukin (IL)-6, IL-1b, and several CC chemokines [80,81]. Intact ECM molecules then facilitate additional leukocyte chemotaxis into the inflamed area by binding these chemokines, thus creating a stable chemotactic gradient to promote specific directional migration [7,82]. Although intact ECM molecules regulate inflammation in multiple ways, matrix fragments also participate in this process. Inflamed tissues contain many proteases that can cleave matrix molecules, generating fragments with

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altered matrix binding, downstream signaling, and effects on inflammatory processes. For example, platelet-derived hyaluronidase cleaves hyaluronan, generating lowemolecular weight hyaluronan fragments; neutrophil elastase cleaves elastin, fibronectin, and collagen XVII; and urinary plasminogen activatore and tissue plasminogen activatoreinduced plasmin activation promote cleavage of collagen XIX, laminin-332, SPARC, and syndecan-4 [4,6,47]. Several of the fragments generated in this manner then promote inflammation in a positive feedback loop. For example, lowemolecular weight hyaluronan fragments induce cytokine and chemokine production through CD44, and collagen XVII and elastin fragments generated by neutrophil elastase promote neutrophil chemotaxis [4,19]. Ultimately, the matrix molecules encountered by leukocytes substantially influence the course of inflammation. The inflammatory phase of wound healing is followed by a proliferative phase. Shortly after wounding, activated platelets secrete a variety of growth factors, including EGF and TGFb, that stimulate the keratinocytes at the wound edge to proliferate and migrate to cover the wounded area, a process known as reepithelialization [83]. FGFs and EGFs are produced and/or released from sequestering matrix molecules at later times by neutrophils, macrophages, endothelial cells, and fibroblasts, and may maintain the proliferative and promigratory signals needed for reepithelialization. During reepithelialization, the keratinocytes migrate beneath the provisional ECM, composed primarily of fibrin and fibronectin [27,65]. The lack of keratinocyte migration on top of the fibrin-based clot may result from their lack of integrin avb3 expression as well as the ability of fibrin to prevent keratinocyte adhesion to other provisional matrix components, including fibronectin [77]. These cells instead migrate on the nascent, provisional basement membrane composed largely of including laminin-332, fibronectin, and tenascin-C, through the a2b1, a3b1, a5b1, a6b1, a9b1, a2b4, and av integrins expressed by these cells [27]. That these matrixereceptor interactions are critical for reepithelialization is clear from studies investigating keratinocyte migration and reepithelialization in cultured keratinocytes migrating on specific matrices, in mice lacking these molecules, and in human patients with matrix mutations. Interestingly, keratinocytes at the migration front produce laminin-332, facilitating the migration of keratinocytes behind the migration front on this matrix component [27]. Because the keratinocytes migrate between the fibrin-based clot and the underlying tissue, their migration is also associated with the activity of multiple proteases, including MMPs and plasmin [27,84]. These enzymes may facilitate keratinocyte migration by promoting their deadhesion from matrix molecules that would otherwise promote adhesion over migration and/or through releasing matrix-bound growth factors, or EGF-like domains from matrix molecules themselves, that then induce promigratory signaling [65,84]. At the same time as reepithelialization, granulation tissue, a provisional connective tissue containing nascent blood vessels and multiple types of ECM molecules, including tenascin-C, cellular fibronectin, SPARC, and various collagens, begins to form [7,26,27]. Granulation tissue serves as substrate for the migration of the keratinocytes (see earlier discussion), the endothelial cells that form the vasculature of the wound bed, fibroblasts, myofibroblasts, and leukocytes that are chemoattracted to the wound site by chemokines secreted by multiple cells within the wound [26]. Chemokine-mediated chemoattraction of cells involved in granulated tissue formation, in conjunction with the interaction of these cells with ECM via cell surface receptors, participates in functions during the formation of the granulation tissue. For example, the neutrophil chemoattractant CXCL8, produced by multiple cell types in the wound environment, induces fibroblast recruitment to the granulation tissue and their deposition of tenascinC, fibronectin, and collagen I [85]. Interactions of both intact matrix molecules and their degradation products with endothelial cells facilitate their migration through the nascent connective tissue to generate new blood vessels during the formation of granulation tissue [27]. As described in more detail previously, fibronectin and vitronectin may promote endothelial cell migration and angiogenesis by enhancing VEGF-induced promigratory signaling through the formation of integrine VEGFR2eVEGFematrix molecule complexes [59]. Proteoglycans may promote endothelial cell migration and angiogenesis through their association with angiogenic factors, including VEGF, CXCL8, and FGF-2, generating a stable gradient to promote directional migration and vessel formation [9,22,85]. During angiogenesis, endothelial cells release and activate matrix-degrading enzymes, including MMPs and cathepsins, which can facilitate the migration and invasion of endothelial cells into the surrounding tissue and generate bioactive matrix fragments that provide additional angiogenic stimuli [4]. Complicating the role of matrix fragments in angiogenesis is that some matrix fragments promote angiogenesis whereas others inhibit angiogenesis by inhibiting endothelial cell migration and/or inducing their apoptosis [4]. This raises the possibility that after angiogenic matrix molecules and fragments induce angiogenesis, these molecules could be cleaved further to generate antiangiogenic fragments that then promote vessel stabilization, basement membrane deposition, and maturation [86]. Therefore, the way matrix molecules are locally cleaved and/or factors are locally released could have important consequences for the formation of the granulation tissue.

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Proliferation During wound reepithelialization, keratinocytes trailing behind those at the front edge of migration replicate to provide a source of cells that cover the wound. Basement membrane-type ECM still present on the basal surface of these keratinocytes may be important in maintaining this less migratory, proliferative state. In a dermal wound model, basement membrane matrices are able to sustain the proliferation of keratinocytes for several days [87]. One component of the basement membrane involved in this proliferation is likely laminin. For example, laminin511 and -521 can promote keratinocyte proliferation in vitro, and keratinocyte proliferation requires the laminin receptor integrin a6b4 [52,68]. Integrin a9b1, which binds several matrix molecules, is critical for keratinocyte proliferation during wound healing and thus in reepithelialization [65]. In contrast, specific cellematrix interactions may prevent excessive proliferation. For example, the keratinocytes of mice deficient in either fibrinogen or emilin-1, a matrix ligand for integrins a4b1 and a9b1, hyperproliferate during reepithelialization, which suggests the importance of these molecules in limiting proliferation during normal reepithelialization [27,52]. As described earlier, granulation tissue forms at the same time as reepithelialization. In the granulation tissue, several types of cells proliferate, including fibroblasts and endothelial cells of the microvasculature. The ECM molecules present in the granulation tissue, in conjunction with growth factors released by the platelets and secreted by the cells present in this tissue, provide signals to promote cell proliferation [8]. ECM molecules themselves, including fibronectin and specific fragments of fibronectin, tenascin-C, laminins, collagen VI, SPARC, and hyaluronan, can stimulate fibroblast and endothelial cell proliferation [2]. For example, fibroblast proliferation requires collagen-induced activation of DDR2, a tyrosine kinase receptor, as shown by the reduced proliferation of DDR2 / in vitro and decreased numbers in wound granulation tissue in vivo, although the decreased numbers in vivo could result from a combination of decreased proliferation and migration [26]. ECM molecules may cooperate with growth factors to promote fibroblast and endothelial cell proliferation. In fibroblasts, proliferation induced by TGFb1 requires fibronectin [88]. Angiogenesis requires both endothelial cell migration (described previously) and proliferation, and angiogenic factors such as VEGFs and FGFs associate with ECM molecules, increasing their ability to activate their receptors and thereby stimulate the proliferation of endothelial cells, which then migrate to form the new microvessels [9,59,85]. Some antiangiogenic molecules, including thrombospondin and endostatin, may inhibit angiogenesis by competing with these growth factors for either growth factor receptor binding or matrix binding [9,10]. In contrast, ECM molecules and/or peptides derived from their proteolysis can have inhibitory effects on cell proliferation. Intact decorin and SPARC, as well as peptides derived from decorin, SPARC, collagens XVIII and XV (endostatin), collagen IV (tumstatin), and tenascin-C have antiangiogenic effects owing to their inhibition of endothelial cell proliferation [4,22].

Differentiation As healing progresses, the healing wound shifts from granulation tissue formation to matrix remodeling, gradually removing the provisional matrix molecules of the granulation tissue and replacing them with a more mature connective tissue rich in collagen I [2]. This process is associated with the differentiation of some fibroblasts into myofibroblasts, acquiring the morphological and biochemical characteristics of smooth muscle cells by expressing a-smooth muscle actin [84]. This differentiation process requires specific growth factors and/or chemokines, including TGFb and CXCL8, as well as fibroblast interactions with multiple types of matrix molecules [84,85]. TGFb1-induced myofibroblast differentiation requires adhesion to the EDA-containing splice variant of fibronectin, likely mediated by integrin a4b7 [5,89]. However, the combined presence of EDA-containing fibronectin and TGFb1 is not sufficient to induce myofibroblast differentiation, which also requires fibroblast adhesion to stiff collagen matrices [2]. After myofibroblast differentiation, these cells secrete copious amounts of matrix molecules, particularly multiple isoforms of collagens, release enzymes that cross-link and thereby stiffen collagen fibrils further, and contract to promote wound closure [26,84]. These myofibroblast activities could cause more fibroblasts to differentiate into myofibroblasts in a positive feedback loop that, if unchecked, could promote excessive fibrosis and abnormal scarring that interferes with normal tissue function [84]. Decorin, a small proteoglycan present in normal wound healing, decreases TGFb1 and collagen production, providing a possible mechanism that could limit myofibroblast differentiation and function to prevent excessive scarring [84]. Differentiation of keratinocytes, endothelial cells, and pericytes is also regulated by cellematrix interactions. In keratinocytes, some matrixeintegrin interactions seem to inhibit terminal differentiation and promote proliferation of basement membraneeassociated cells, whereas hyaluronaneCD44 binding may promote terminal differentiation [64,66]. The differentiation of endothelial cells in mature blood vessels requires physical interactions with basement

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membrane components such as laminin-411, as well as interactions with pericytes [68]. Pericyte differentiation, in turn, requires integrin b1edependent cellematrix interactions [90].

Apoptosis In healing wounds, many cells that are needed for one specific phase of the healing process undergo apoptosis after completing their respective functions. In fact, the persistence of some cells, including inflammatory cells and myofibroblasts, is detrimental, and apoptosis is needed to prevent chronic inflammation and excessive scarring, respectively [89,91]. ECM molecules regulate some of the inflammatory cell apoptosis after their activation. For example, lowemolecular weight hyaluronan promotes neutrophil and macrophage apoptosis, likely through CD44 [19,47]. However, apoptosis in neutrophils is primarily regulated by a constitutive signaling pathway that can be delayed by inflammatory signaling or by fibrinogen binding, but otherwise seems little affected by matrix interactions [91,92]. Apoptosis also participates in the wound remodeling phase, because the granulation tissue evolves into a relatively acellular scar tissue [89]. In this remodeling phase, apoptotic cell death eliminates many types of cells at the same time without causing tissue damage. Within the granulation tissue, the number of cells undergoing apoptosis increases around days 20e25 after injury, dramatically reducing wound cellularity after day 25 [89a]. This coincides with the release of mechanical tension after myofibroblast-mediated wound contraction, which triggers apoptosis of human fibroblasts and myofibroblasts, which suggests the importance of interstitial collagens, their receptors, and mechanotransduction in myofibroblast apoptosis [7]. This myofibroblast apoptosis may be required to both promote wound healing and prevent scarring. Indeed, a type of excessive, abnormal scar called a hypertrophic scar exhibits reduced fibroblast/myofibroblast apoptosis, resulting in excessive fibrosis and scarring [84]. This apoptotic failure in hypertrophic scars likely results from an overexpression of tissue transglutaminase, leading to increased matrix breakdown and decreased collagen contraction [2].

CELLeEXTRACELLULAR MATRIX INTERACTIONS DURING REGENERATIVE FETAL WOUND HEALING True tissue regeneration after injury rarely occurs in vertebrate species, but it occurs in specific instances, including fetal cutaneous wound healing, liver regeneration, and urodele amphibian limb regeneration. Unlike wound healing in normal adult animals, which is characterized by scarring, fetal cutaneous wounds heal without fibrosis and scar formation, leading to regeneration of the injured area. The contribution of celleECM interactions to regeneration in fetal healing is discussed subsequently (Fig. 2.3).

FIGURE 2.3 Comparison of particular celleextracellular matrix (ECM) interactions occurring in scar-forming adult healing versus those occurring during regenerative fetal healing. As shown in this diagram, unique subsets of ECM molecules are associated with scarring versus regenerative healing. As such, therapeutic alteration of ECM composition may allow physicians to modulate healing to promote tissue regeneration. Additional therapeutic approaches may be generated upon further investigation into the importance of additional celleECM interactions in scarring and regenerative responses. ED-A, extra domain A; TGF-1, transforming growth factor-1.

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Adhesion and Migration Fetal wounds, at least these relatively early in fetal development, heal without scarring, in contrast to most adult healing, which heals with at least some scarring [89]. One major difference between scarless fetal healing and adult healing is the lack of an inflammatory response before embryonic day (E)15e16 in mouse development, and the attenuated inflammatory response seen after E18 [7]. Indeed, the onset of scarring in fetal healing at later embryonic stages coincides with the substantial appearance of mast cells and neutrophils in fetal wounds [2,93,94]. Mast cells after E18 have much larger mast cell granules, which contain higher levels of proinflammatory cytokines that may then contribute to the increased inflammatory response after this stage of development [95]. This lack of neutrophils in early fetal wounds is also associated with the low expression of neutrophil adhesion molecules, decreasing extravasation coupled with low levels proinflammatory cytokines chemokines such as TNFa, IL-6, and CXCL8, and increased levels of antiinflammatory cytokines such as IL-10 [96]. Experimental increases in IL-10 in adult wounds reduced inflammation and promoted regenerative healing, providing important evidence for the role of inflammation in the neutrophil recruitment and scarring present in adult wounds [89,96]. Scarless fetal wounds also differ from scarring adult wounds in celleECM interactions owing to differences in the composition of the ECM molecules, the timing of their appearance after wounding, and their duration in the wound area. One crucial ECM molecule in fetal wound healing is hyaluronan, which appears to be necessary for the regenerative response, because its removal from fetal wounds promotes a healing response more similar to that of adults, and treatment of normally scarring wounds or wound organ cultures with hyaluronan decreases scarring [97e99]. Highemolecular weight hyaluronan is more abundant in fetal skin wounds than in adult wounds, where lowemolecular weight hyaluronan is more abundant; the latter possibly results from increased hyaluronidase activity in adult wounds [2,83]. Fetal fibroblasts also express higher levels of the hyaluronan receptors CD44 and RHAMM after injury, thus increasing receptoreligand interactions that promote fibroblast migration [98]. Tenascin C, fibronectin, and collagen levels also differ in adult and fetal wounds. Tenascin C is expressed at higher levels in fetal skin than in adult skin and is induced more rapidly and to a greater extent in fetal wounds, modulating cell adhesion to fibronectin and promoting migration within matrices containing fibronectin [27]. Fibronectin production increases more rapidly in fetal wounds, although the fibronectin produced does not contain the profibrotic EDA domains [2,27]. This increased expression of tenascin and fibronectin is associated with concomitant increases in the expression of integrins that serve as their receptors. In particular, a5 subunit, avb3, and avb6 integrins, which bind fibronectin and/or tenascin, are upregulated in the wounded fetal epithelium [100]. The combined rapid increases in fibronectin and tenascin, coupled with increased expression of their respective integrin receptors in epithelial cells, are likely important in facilitating cell migration and reepithelialization in fetal wounds. In contrast to the increased levels and/or rate of tenascin C and fibronectin production in fetal wounds, these wounds contain reduced levels of total collagen compared with adult wounds. However, fetal wounds contain a greater proportion of collagen III compared with collagen I than do their adult counterparts [84]. The observed changes in the collagen I/III ratio in fetal wounds and their relative lack of collagen deposition and fibrosis may result from changes in the deposition, organization, and cross-linking of collagen at the wound site or rapid turnover of these ECM components by protease-mediated degradation [84,98]. Related to increases in matrix turnover, fetal wounds have increased levels and activity of multiple MMPs, with decreased levels of their endogenous inhibitors, tissue inhibitor of metalloproteinases, ultimately promoting matrix degradation and turnover [84]. Not only does the resulting matrix degradation and turnover prevent fibrosis, it also likely facilitates cell migration by reducing matrix density and increases the generation of proteolytic matrix fragments that modulate various stages of wound repair.

Proliferation Increased levels of hyaluronan present during fetal wound healing likely decrease fetal fibroblast proliferation [101]. Fetal fibroblasts also exhibit decreased proliferation in response to growth factors compared with that of adult cells. For example, IGF-1, which induces ERK signaling, proliferation, and matrix synthesis in postnatal fibroblasts, induces proliferation to a much lesser extent and fails to induce significant ERK signaling or matrix synthesis in fetal fibroblasts [102]. Furthermore, whereas TGFb1 induces proliferation in postnatal fibroblasts, it does not do so in fetal fibroblasts, possibly because of the ability of TGFb1 to induce hyaluronan synthesis in fetal but not postnatal fibroblasts [103].

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Differentiation Fetal wounds have a decreased number of myofibroblasts, which appear in the wounded site earlier and remain a shorter time than in adult wounds [89,98]. Fetal fibroblasts produce more type III collagen and less type I collagen than adult cells, and the diameter and organization of the fibrils in the fetal wound are comparable to unwounded skin, whereas those of the adult wound exhibit a disorganization indicative of scarring [98,104]. Adult wound tissue is stiffer than fetal tissue, likely resulting from increased collagen I levels in adult wounds, facilitating myofibroblast differentiation that depends on stiff collagen matrices [2]. In contrast, fetal wounds have a decreased number of myofibroblasts, which appear in the wounded site earlier and are more transient than in adult wounds [89,98]. The low number of myofibroblasts in fetal wounds may result, at least in part, from a lack of collagen matrix stiffness [83]. Because myofibroblasts are themselves responsible for much of the collagen I production and tissue contraction in adult wound tissue, their relative absence in fetal wounds may be responsible for the reduced collagen I levels and lower contraction in fetal wounds [89]. Increased fetal hyaluronan may also prevent myofibroblast differentiation by increasing expression of TGFb3, which is antifibrotic, in contrast to TGFb1, which increases collagen I deposition and promotes scar formation [84,98]. Indeed, adult wounds treated with hyaluronan healed more rapidly with a significant decrease in TGFb1 levels, which suggests that the large amounts of hyaluronan in fetal wounds may thus explain, at least partly, the greatly reduced levels of TGFb1 in fetal wounds [83]. Downregulation of TGFb1 in adult wounds produces a decrease in scarring similar to that observed with hyaluronan treatment, whereas addition of TGFb1 to normally scarless fetal wounds induces a more scarring phenotype, with myofibroblast differentiation, wound contraction, and fibrosis [98,105]. Thus, hyaluronan-mediated inhibition of TGFb1 expression may be critical in scarless fetal healing. The relatively small amount of TGFb1 present during fetal wound healing may be regulated by inhibitory ECM molecules present in the injured area. One such inhibitor is the proteoglycan decorin, which is capable of binding TGFb1 and preventing receptor activation and is expressed to a greater extent in fetal wounds than in adult wounds [2]. Decreased activity of this growth factor, combined with low levels of expression in fetal wounds, likely results in decreased fibrosis, myofibroblast differentiation, and wound contraction, leading to regeneration rather than scarring.

Apoptosis Less is known about the apoptotic process in fetal wounds than in adult wounds, which makes it difficult to compare cellular apoptosis under these conditions. An investigation of apoptotic induction at very early time points after wounding in both scarless (E15) and scar-forming (E18) fetal mouse wounds found lower apoptosis in scarforming wounds than scarless wounds [106]. Some of these cells disappearing from scarless wounds may be myofibroblasts, because any myofibroblasts that differentiate from fetal fibroblasts, either in vivo or in vitro, disappear rapidly, perhaps owing to an altered rate of apoptosis in these wounds [106]. If changes in apoptotic efficiency indeed occur, they may result from the decreased contraction, and thus decreased mechanical tension, in fetal wounds, as well as altered collagen levels within the collagen matrix [7]. However, myofibroblasts may disappear from fetal wounds through dedifferentiation back to fibroblasts, which complicates the picture [89]. Perhaps apoptosis is not as critical in the healing of fetal wounds as in adult wounds, because leukocyte influx and myofibroblast differentiation appear to be minimal in fetal wounds, and thus may not require large numbers of cells to undergo apoptosis for regeneration to occur [7,89].

IMPLICATIONS FOR REGENERATIVE MEDICINE One primary goal of studies comparing differences in celleECM interactions, and thus changes in signaling, that accompany regenerative and nonregenerative healing is to determine which types of interactions promote and which inhibit tissue regeneration (for an example, see Fig. 2.3). After elucidating the functions of particular interactions, it may be possible to increase the regenerative response through (1) the induction of proregenerative ECM molecules or signaling events in the wounded area combined with (2) the antagonism of antiregenerativeescarring interactions or signaling events using specific inhibitors. This discussion of regenerative medicine will focus on possible strategies to promote regeneration in adult scarring wounds, thus causing adult wounds to resemble more closely fetal scarless wounds. Such an increased regenerative response would be particularly useful in treating wounds that heal with increased scar formation, such as keloids and hypertrophic scars.

REFERENCES

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Different types of approaches may be used to increase proregenerative ECM levels in the wounded area, including direct application of the matrix molecules themselves, the addition of agents that increase their expression, the addition of cells producing these types of ECM that have been prepared to minimize immunogenicity, the introduction of biomaterials modified to contain adhesive, proregenerative regions of these ECM molecules, or wound treatment with inhibitors of their proteolysis. Several different ECM molecules are present at higher levels in fetal wounds than in adult wounds, including hyaluronan, tenascin, fibronectin, and collagen III, which may have important roles in the regeneration process [2,27,98]. Thus, altering the levels of these molecules in a scarring wound may improve regeneration. In keeping with the substantial evidence supporting a role for hyaluronan in scarless healing, multiple types of biomaterials currently used to promote healing incorporate hyaluronan in the presence or absence of other matrix molecules or various cell types, although the antiscarring outcomes of these biomaterials vary [107]. One potential reason for the observed variability lies in the rapid degradation of hyaluronan caused by hyaluronidase activity in vivo, and a biomaterial containing modified, more hyaluronidase-resistant hyaluronan mimics improved healing in multiple contexts [99]. Other biomaterials have been used with some success, including those with substantial amounts of collagen III and “natural” scaffolds containing multiple matrix molecules and proteoglycans [99,108]. Combinations of hyaluronan with other matrix molecules, including tenascin, embryonic fibronectin, and/or collagen III, should mimic the fetal wound environment better and may lead to more regenerative healing. Alternatively, various matrix scaffolds could be combined with growth factors that promote healing and reduce scarring, using the matrix molecules to deliver growth factor more effectively to the wound and thereby better facilitate a regenerative response [76]. Alterations in the biomaterial formulation, such as adding molecules that bind and release growth factors or engineering growth factors that associate more effectively with biomaterials, could regulate the timing of growth factor release, allowing their release over a relatively long period to promote more effective healing [76]. These or other biomaterials may be useful for the delivery of proregenerative ECM molecules and/or growth factors to the injured area, thereby promoting healing and reducing scar formation. When attempting to promote regeneration, it is also imperative to inhibit events associated with scarring, including excessive ECM deposition, fibrosis, and contraction. During the adult healing process, these scarassociated processes are primarily controlled by the myofibroblast, a differentiated cell type that arises during the adult healing process, but which is largely absent throughout fetal wound healing. Therefore, inhibition of myofibroblast differentiation or function in combination with the addition of proregenerative molecules may facilitate a stronger regenerative response. Inhibition of differentiation could be accomplished by blocking factors that normally stimulate myofibroblast differentiation, such as TGFb1 and CXCL8, by preventing fibroblasteECM interactions that facilitate myofibroblast differentiation, such as EDA-containing fibronectin, and by the delivery of antifibrotic molecules such as TGFb3 and IL-10 [84,85,96]. It is also possible that the application of one molecule may promote more regenerative healing by affecting multiple parts of the healing process. For example, insulin interaction with its receptor affects multiple aspects of keratinocyte behavior, stimulating cell motility, increasing expression of the cell surface adhesion molecule integrin a3, enhancing secretion of the ECM molecule laminin332, and improving epidermal differentiation during wound healing [109]. Furthermore, we have shown that insulin stimulates the formation of regenerative rather than scarring matrix [110]. The surge in research regarding ECM molecules themselves and their interactions with particular cells and cell surface receptors has led to the realization that such interactions are many and complex, and that they are of the utmost importance in determining cell behavior during events such as wound repair and tissue regeneration. Therefore, the manipulation of specific celleECM interactions has the potential to modulate particular aspects of the repair process and thereby promote a regenerative response.

Acknowledgments We thank acgdesign.com for the construction of the figures presented in this article.

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3 Mechanisms of Blastema Formation and Growth in Regenerating Urodele Limbs David L. Stocum Indiana University-Purdue University, Indianapolis, IN, United States

INTRODUCTION The limbs of larval and adult urodele amphibians are unique among tetrapod vertebrates in their ability to regenerate from any level of the limb after amputation. Limb regeneration can be divided into two major phases: (1) formation of a blastema that resembles the early embryonic limb bud and (2) blastema redevelopment, which involves growth and redifferentiation. Blastema formation refers in this chapter to events leading to the establishment of an accumulation of undifferentiated cells under a thickening of the distal wound epidermis, the apical epidermal cap (AEC). Pattern formation, in which the spatial relationships of the structures to be regenerated are determined and specified, is a process that spans both phases. The ability to form a blastema after amputation is what distinguishes the limbs of urodeles from those of anuran amphibians, reptiles, birds, and mammals, and is the primary focus of this chapter. Blastema formation is a reverse developmental process realized partly by cell dedifferentiation in tissues local to the amputation plane [1] and partly by a contribution of muscle stem cells [2]. Blastema development is similar to that of the embryonic limb bud, with one major exception: blastema cell proliferation depends on an interaction between the limb nerves and the AEC, whereas proliferation of embryonic limb bud cells relies on an epithelialemesenchymal interaction among the counterpart of the AEC, the apical ectodermal ridge, and the subjacent mesenchymal cells. The musculoskeletal and skin tissues of the new limb parts derived from the blastema redifferentiate in continuity with their parent tissues, whereas blood vessels and nerves regenerate by extension from the cut ends of the preexisting blood vessels and axons, respectively. The growing blastema goes through several morphological/histological stages to attain a cone of cells that then broadens and initiates differentiation in proximal to distal and anterior to posterior directions, ending in distal bifurcations of the digits. If we were able to understand why some animals such as urodele amphibians are able to form a regenerationcompetent blastema after amputation whereas others such as adult anurans, birds, and mammals are not, it might be possible to design chemical approaches to inducing blastema formation in human appendages. At the least, such knowledge might improve our ability to deal with nonamputational injuries to musculoskeletal, vascular, and neural tissues. With this in mind, I review here what is known about blastema formation in the regeneration-competent limbs of urodeles and compare it with blastema formation in the regeneration-deficient anuran, Xenopus laevis.

BLASTEMA FORMATION Blastema formation in regenerating urodele limbs can be subdivided into three phases: (1) hemostasis and reepithelialization, (2) histolysis and dedifferentiation, and (3) blastema cell migration and accumulation (Fig. 3.1). The latter two phases overlap with one another. Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00003-5

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(A)

(B) AMP BLASTEMA FORMATION

AB BLASTEMA RE-DEVELOPMENT

MB LB

(C)

2FB 3FB 4FB

FIGURE 3.1 (A) Diagram of phases and stages of regeneration after amputation of a urodele limb. The two black lines indicate the two major phases of regeneration (blastema formation and blastema redevelopment) and the stages of regeneration after amputation (AMP). 2FB, 3FB, 4FB, finger bud stages; AB, accumulation blastema; LB, late bud; MB, medium bud. The colored lines indicate different subphases of blastema formation and redevelopment. Blue, pattern formation; green, blastema growth; orange, histolysis and dedifferentiation; white, hemostasis and reepithelialization; yellow, redifferentiation. (B) Longitudinal section of regenerating axolotl hindlimb 4 days after amputation through the midtibia-fibula. Arrow points to the thickening apical epidermal cap (AEC). The cartilage (C), muscle (M), and other tissues are breaking down in a region of histolysis and dedifferentiation (H/DD) under the wound epithelium. Magnification 10, light green and iron hematoxylin stain. (C) Longitudinal section of regenerating axolotl hindlimb 7 days after amputation through the midtibia-fibula. An accumulation blastema (AB) has formed by the migration of dedifferentiated cells under the AEC. Arrows mark the junction between the accumulation blastema and the still-active region of histolysis and dedifferentiation proximal to it. magnification 10, light green and iron hematoxylin stain.

Hemostasis and Reepithelialization After limb amputation or after making skin wounds in amphibians, vasoconstriction occurs and a thrombincatalyzed fibrin clot forms within seconds to protect the wound tissue and provide a temporary matrix from which repair or regeneration is initiated. An epithelium two to three cells thick covers the wound surface within 24 h after amputation, depending on limb size. The basal epidermal cells at the cut edge of the skin migrate as a sheet that is extended by mitosis of cells adjacent to the wound edges [3]. The fibrin clot contains significant amounts of fibronectin, which the epithelial sheet uses as a substrate for migration [4]. Within 2e3 days postamputation (dpa), the wound epidermis thickens to form the AEC. The basal cells and gland cells of the wound epidermis/AEC have secretory functions, as evidenced by their more extensive endoplasmic reticulum and Golgi network [5]. WE3, 4, and 6 are three secretory-related antigens expressed specifically by dermal glands and wound epidermis/AEC [6]. Two other antigens, 9G1 [7] and NvKII [8], are also specific to the wound epidermis, but their functions are unknown. The early wound epidermis has an important function in generating early signals for limb regeneration. Naþ influx in the amputated newt limb and Hþ efflux in the amputated tail of Xenopus tadpoles generate ionic currents across the wound epidermis essential for regeneration. Naþ influx is via sodium channels [9]. Hþ efflux in the amputated tail is driven by a plasma membrane adenosine triphosphatase (ATPase) in the epidermal cells [10] and is likely to be important for urodele limb regeneration as well, given that a gene encoding a v-ATPase was the most abundant clone in a suppressive subtraction complementary DNA library made from 4-dpa regenerating limb tissue in the axolotl [11]. Drug-induced inhibition of either Naþ or Hþ movements during the first 24 h or so after amputation results in failure of blastema formation [10,12]. Inositol triphosphate (IP3) and diacylglycerol (DAG) are the products of phosphatidylinositol bisphosphate (PIP2), which in turn is derived from inositol. IP3 synthase, a key enzyme for the synthesis of inositol from glucose-6-phosphate, is upregulated during blastema formation in regenerating axolotl limbs [13]. IP3 stimulates a rise in cytosolic Ca2þ that results in the localization of protein kinase C (PKC) to the plasma membrane, where PKC is activated by DAG and regulates transcription [14]. During blastema formation, there is a general

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39

downregulation of proteins involved in Ca2þ homeostasis, which suggests that IP3 might signal a rise in cytosolic Ca2þ in regenerating limbs to localize PKC to the plasma membrane [13]. Other studies have shown that IP3 is generated from PIP2 within 30 s after amputation in newt limbs [15] and that PKC rises to a peak by the accumulation blastema stage [16]. How these early signals are linked to the next phase of blastema formation, histolysis, and dedifferentiation, is unknown.

Histolysis and Dedifferentiation Histolysis is the loss of tissue organization resulting from the enzymatic degradation of the extracellular matrix (ECM). Dedifferentiation is the reversal of a given state of differentiation to an earlier state via nuclear reprogramming and loss of specialized structure and function. All of the tissues subjacent to the wound epidermis undergo intense histolysis (ECM degradation and tissue disorganization) for 1e2 mm, resulting in the liberation of fibroblasts, Schwann cells of the peripheral nerves, and skeletal cells [1]. Myofibers fragment at their cut ends and break up into mononucleate cells while releasing satellite cells (the stem cells that effect muscle regeneration). The liberated cells lose their phenotypic specializations and revert to mesenchymal-like cells with large nuclei and sparse cytoplasm that exhibit intense RNA and protein synthesis. Histolysis and dedifferentiation begin within 2e3 dpa in larval urodeles and within 4e5 days in adults. Mechanisms of Histolysis Degradation of tissue ECM is achieved by acid hydrolases and matrix metalloproteinases (MMPs) [17]. Acid hydrolases identified in regenerating urodele limbs include cathepsin D, acid phosphatase, b-glucuronidase, carboxyl ester hydrolases, and N-acetyl-glucosaminidase. Osteoclasts are abundant in the region of histolysis, where they degrade bone matrix via hydrochloric acid, acid hydrolases, and MMPs. Upregulated MMPs include MMP-2 and -9 (gelatinases), and MMP-3/10a and 10b (stromelysins) [18]. Macrophages are a major source of MMPs, particularly MMP-9 [19]. The basal layer of the wound epidermis is a source of MMP-3/10a and 10b in the newt limb, as well as of a novel MMP with low homology to other MMPs [20]. These MMPs are responsible for maintaining contact between the wound epidermis and the underlying tissues by preventing reassembly of a basement membrane. Chondrocytes are a source of MMP-2 and -9 in the newt limb, and these enzymes diffuse outward from the degrading skeletal elements [20]. The importance of MMPs to histolysis, and the importance of histolysis to the success of regeneration, are underscored by the failure of blastema formation in amputated newt limbs treated with an inhibitor of MMPs (GM6001) [21]. Histolysis continues to contribute blastema cells until the medium bud stage; then it ceases owing to the activity of tissue inhibitors of metalloproteinases (TIMPs) [18,22]. TIMP1 is upregulated during histolysis, when MMPs are at maximum levels, and exhibits spatial patterns of expression congruent with those of MMPs in the wound epidermis, proximal epidermis, and internal tissues undergoing disorganization. Mechanisms of Dedifferentiation Dedifferentiation is a complex and poorly understood process involving epigenetic reprogramming that suppresses the transcription of differentiation genes, while activating transcription of genes associated with stemness, reduction of cell stress, and remodeling internal structure. Inhibition of these transcriptional changes by actinomycin D does not affect histolysis, but it prevents or retards dedifferentiation, leading to regenerative failure or delay [23]. This suggests that at least part of the proteases involved in histolysis are not regulated at the transcriptional level, but that proteins effecting dedifferentiation are thus regulated. Dedifferentiated cells express a more limb budelike ECM in which the basement membrane is absent, type I collagen synthesis and accumulation are reduced, and fibronectin, tenascin, and hyaluronate accumulate [24e26]. The molecular details of transcriptional regulation during dedifferentiation are largely unknown. Degradation of the ECM by proteases would break contacts between ECM molecules and integrin receptors, leading to changes in cell shape and reorganization of the actin cytoskeleton that might activate epigenetic reprogramming. Stemness genes upregulated during blastema formation are msx1, nrad, rfrng, and notch [17]. Msx1 inhibits myogenesis [27] and its forced expression in mouse myotubes causes cellularization and reduced expression of muscle regulatory proteins [28]. Inhibition of msx1 expression in cultured newt myofibers by anti-msx morpholinos prevents their cellularization [29]. Newt regeneration blastema extract stimulates mouse myonuclei to reenter the cell cycle, cellularize, and reduce their expression of muscle regulatory proteins [30]. Nrad expression is correlated with muscle

40

3. BLASTEMA FORMATION AND GROWTH

dedifferentiation [31], and Notch is a major mediator of stem cell self-renewal [32]. A micro-RNA gene regulatory circuit has been identified in regenerating axolotl limbs and fish fins [33]. Three of the six transcription factor genes (klf4, sox2, and c-myc) used to reprogram mammalian adult somatic cells to induced pluripotent stem cells (iPSCs) [34,35] are upregulated during blastema formation in regenerating newt limbs, and also during lens regeneration [36]. The Lin 28 protein, the product of a fourth transcription factor gene used to derive iPSCs [35], also is upregulated during blastema formation in regenerating axolotl limbs [13]. Thus, transcription factors that reprogram fibroblasts to iPSCs may also gave a role in nuclear reprogramming during limb regeneration. The further molecular characterization of transcription factors, micro-RNAs, and changes in epigenetic marks via chromatin-modifying enzymes will be crucial for understanding the mechanism of dedifferentiation in regenerating amphibian limbs. The differential regulation of pathways that protect cells from stress and apoptosis also have a role in dedifferentiation. Proteomic analysis suggests that reduced metabolic activity, upregulation of pathways that accelerate protein folding or eliminate unfolded proteins (the unfolded protein response), and differential regulation of apoptotic pathways may largely prevent apoptosis [13], which is known to be minimal in regenerating limbs [37,38]. This idea is consistent with other studies on cultured chondrocytes, b cells, and Muller glia cells of the retina showing that cells dedifferentiate as part of a mechanism to combat apoptotic cell stress [13]. The molecular details of internal cellular remodeling are poorly understood. Dismantling of the phenotypic structure and function is most visible in myofibers, but the molecular details of the process are largely uninvestigated for any limb cell type. Two small molecules, one a trisubstituted purine called myoseverin and the other a disubstituted purine dubbed reversine, have been screened from combinatorial chemical libraries and have been found to cause cellularization of C2C12 mouse myofibers [39,40]. Myoseverin disrupts microtubules and upregulates genes for growth factors, immunomodulatory molecules, ECM remodeling proteases, and stress-response genes, which is consistent with the activation of pathways involved in wound healing and regeneration, but it does not activate the whole program of myogenic dedifferentiation in newt limbs [41]. Reversine treatment of C2C12 myotubes resulted in mononucleate cells that behaved like mesenchymal stem cells, i.e., they were able to differentiate in vitro into osteoblasts and adipocytes as well as muscle cells [42]. Myoseverin and reversine might be useful in analyzing the events of structural remodeling, and may have natural counterparts that can be isolated.

Differential Tissue Contributions to the Blastema Individual tissues of the limb make differential contributions to the blastema. In the axolotl limb, dermal fibroblasts represent 19%, and chondrocytes 6%, of the cells present at the amputation surface [43]. Dermal fibroblasts contribute nearly half of the blastema cells; fibroblasts of the periosteum and myofiber/nerve sheath interstitial connective tissue, Schwann cells, and myogenic tissue contribute the rest. Experiments transplanting green fluorescent protein (GFP) cartilage into a limb wound induced the formation of a supernumerary limb, which suggests that chondrocytes make no contribution to the blastema [44]. Transplants of other GFP-labeled tissues have shown that the redifferentiation of blastema cells is largely lineage-specific, i.e., they are constrained to reproduce their parent cell types. The exception is connective tissue fibroblasts, which after dedifferentiation are able to transdifferentiate at high frequency into cartilage [45e47]. Lineage-restricted redifferentiation is also reflected in stably maintained histone methylation patterns of parent cells [48]. Thus, although transcription factors that reprogram fibroblasts to iPSCs have a role in nuclear reprogramming during limb regeneration, other factors clearly ensure that dedifferentiated cells reverse their transcription programs only far enough to attain a state that can respond to proliferation and patterning signals while maintaining their phenotypic memory of origin. There are developmental and species differences in muscle contributions to the blastema. Muscle regenerates in larval and adult axolotl and larval newt limbs by Pax7þ satellite cells, whereas in adult newts, muscle regenerates via fragmentation of the ends of cut myofibers and dedifferentiation of the resulting mononucleate cells [49,50]. Whether the switch from muscle regeneration via satellite cells in newt larvae to dedifferentiation of myofiber fragments in adults is all-or-none or gradual is unknown.

Cell Cycling During Blastema Formation Tritiated thymidine (3H-T) labeling studies have shown that cells of amputated urodele limbs initiate cell cycle entry coincident with their liberation by histolysis. The pulse-labeling index reaches 10e30% during formation of the accumulation blastema [51,52]. However, the mitotic index is low, between 0.1% and 0.7% (average of about 0.4%, or 4 in

BLASTEMA FORMATION

41

1000 cells) in both Ambystoma larvae [53] and adult newt [54]. The low mitotic index during establishment of the accumulation blastema suggests that it forms primarily by accumulating dedifferentiated cells rather than their mitosis. The fact that cells readily enter the cell cycle during formation of the accumulation blastema but divide only infrequently suggests that a large proportion of dedifferentiating cells arrest in G2 (M51). Further indirect evidence for G2 arrest is the strong upregulation of the ecotropic viral integration factor 5 (Evi5) throughout blastema formation in regenerating axolotl limbs [13]. Evi5 is a centrosomal protein that accumulates in the nucleus during early G1 in mammalian cells and prevents them from prematurely entering mitosis by stabilizing Emi1, a protein that inhibits cyclin A degradation by the anaphase-promoting complex/cyclosome [55]. At G2, Emi1 and Evi5 are phosphorylated by Polo-like kinase 1 and targeted for ubiquitin-driven degradation, allowing the cell to enter mitosis. Thus, high levels of Evi5 during blastema formation may restrain cells from entering mitosis until they are fully dedifferentiated and present in enough numbers to form an accumulation blastema [13]. To test this hypothesis, it will first be necessary to determine the spatiotemporal expression pattern of Emi1 and Evi5. The hypothesis predicts that these proteins would be expressed at high levels in both wound epidermis and blastema mesenchyme, and that expression would decrease as the cells transited a normal cell cycle. The signals that drive liberated cells to enter the cell cycle have been studied in detail in myofibers of the regenerating newt limb. Entry into the cell cycle of muscle-derived blastema cells appears to be initiated by the muscle LIM protein (MLP), a member of the MARCKS family that has a role in muscle differentiation [56]; whether MLP initiates cell cycle entry of blastema cells derived from other limb tissues is unknown. Progression through G1 and S in cultured newt and mouse myoblasts and newt myofibers is promoted by a thrombin-activated factor present in the serum of all vertebrates tested thus far, including mammals, that deactivates the Rb protein [57]. Mouse myofibers do not respond to this factor. Newt blastema extract promotes DNA synthesis in both newt and mouse myofibers [30], which suggests that mouse myofibers lack an essential signal pathway ingredient that is supplied by newt blastema extract but not by serum. Although the thrombin-activated protein is both necessary and sufficient to stimulate the entry of myonuclei into the cell cycle, it is not sufficient to drive them through mitosis, and myonuclei arrest in G2. Cell cycle reentry is independent of myofiber cellularization, because cell cycleeinhibited myofibers implanted into newt limb blastemas break up into mononucleate cells [58]. However, mitosis appears to require mononucleate cell status. The mechanism of myofiber fragmentation into single cells is not known, nor is it known whether the thrombin-activated protein is also necessary to drive cells such as dedifferentiating fibroblasts and Schwann cells into the cell cycle or whether this is a feature unique to myofibers. Biochemical evidence suggests that the thrombin-activated factor may be a potent growth factor required in very small amounts [57].

Macrophages Have an Important Role in Blastema Formation Macrophages of the innate immune system are important mediators of wound repair in mammals via their bactericidal and phagocytic activities and their secretion of growth factors and cytokines that modulate inflammation and initiate structural repair by fibroblasts [17]. Studies emphasize the importance of the immune system, particularly macrophages, for the events of blastema formation during urodele limb regeneration [59e61]. Proinflammatory and antiinflammatory cytokines are upregulated during blastema formation in regenerating limbs of adult axolotls, coincident with a significant enrichment of macrophages, which produce both types of cytokines as well as MMPs. Macrophage depletion by liposome-encapsulated clodronate during blastema formation results in regenerative failure and scarring of the limb stump. The epidermis closes the wound but does not develop an AEC, and dermal scar tissue is interposed between the wound epidermis and underlying tissues [19]. By contrast, depletion after a blastema enters the growth phase only delays regeneration. These results suggest a central role for macrophages in resolving inflammation and the degradation of ECM. Macrophages have also been shown to have a role in removing senescent cells during urodele limb regeneration. In mammals, senescent cells accumulate in tissues with age. Few cells undergoing apoptosis are detected in regenerating urodele limbs [37]. Cell senescence is induced during blastema formation in regenerating urodele limbs, but senescent cells are cleared by macrophages and do not accumulate [62].

Blastema Cell Migration and Accumulation The AEC appears to direct the migration of blastema cells to aggregate beneath it [1]. This was shown by experiments in which shifting the position of the AEC laterally caused a corresponding shift in blastema cell accumulation, and transplantation of an additional AEC to the base of the blastema resulted in supernumerary blastema

42

3. BLASTEMA FORMATION AND GROWTH

formation. Nerve guidance of blastema cells to form eccentric blastemas appeared to be ruled out, because similar experiments on aneurogenic limbs resulted in eccentric blastema formation. The redirected accumulation of blastema cells in these experiments results from the migration of the cells on adhesive substrates produced by the eccentric AEC. Transforming growth factor (TGF)-b1 is strongly upregulated during blastema formation in amputated axolotl limbs [63]. A target gene of TGF-b1 is fibronectin, a substrate molecule for cell migration that is highly expressed by basal cells of the wound epidermis during blastema formation [13,64]. Inhibition of TGF-b1 expression by the inhibitor of SMAD phosphorylation, SB-431542, reduces fibronectin expression and results in failure of blastema formation [63]; this suggests that fibronectin provided by the AEC provides directional guidance for blastema cells.

BLASTEMA GROWTH There are two synergistic requirements for cells of the accumulation blastema to break G2 arrest, enter mitosis, and proliferate. The first is the production of mitogens via an interaction of the AEC with nerve axons. The second is that cells of the accumulation blastema derived from opposite positions on the limb circumference must migrate toward one another and interact. Unless these two conditions, along with vascularization, are met, the accumulation blastema will not persist and will regress.

The Apical Epidermal CapeNerve Interaction Neither denervation at the time of amputation nor deprivation of wound epidermis prevents histolysis and formation of blastema cells, but the blastema cells do not persist [65]. The 3H-T labeling index is the same as that of controls during formation of the accumulation blastema in both epidermis-free and denervated limbs, which suggests that neither the nerve nor the wound epidermis is required for DNA synthesis during this time [53,66e68]. The AEC of the accumulation blastema is invaded by sprouting sensory axons as it forms, whereas other sensory axons and motor axons make intimate contact with mesenchyme cells of the blastema [69]. Coincident with innervation of the AEC, the labeling and mitotic indices of the accumulation blastema rise as much as 10-fold (Fig. 3.2) and the blastema enters the phase of growth and patterning [51,52]. These increases do not take place in denervated or wound epidermisedeprived limbs. 3H-T pulse labeling studies indicate that the final cycling fraction of blastema cells is between 92% and 96% in larvae and over 90% in adults [70,71]. These results are consistent with the idea that blastema formation results from cell migration and aggregation, not mitosis.

(A)

3H-T

MI

(B)

90

4.0

30

0.4 0.0 AB

AB 3

FIGURE 3.2 Diagram of changes in tritiated thymidine ( H-T) labeling and mitotic index (MI) during blastema formation and growth, expressed as percentages of the total cell number on the ordinate. AB on the abscissa represents the accumulation blastema stage. The growth phase is to the right of the AB. (A) Before the accumulation blastema stage, the 3H-T labeling index is the same in control (green line) and in epidermis-free and denervated limbs (both represented by the red line). These indices in deprived limbs fail to rise in concert with the controls during blastema growth and an accumulation blastema does not form. (B) Before the accumulation blastema stage, the basal mitotic index of controls (green line) and epidermis-free limbs (red line) are nearly identical, but the MI does not increase with the controls during blastema growth. In contrast, the MI in denervated limbs (blue line) does not achieve the basal level and remains near zero. An accumulation blastema does not form in either denervated or epidermis-free limbs.

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When the growing blastema achieves a critical mass of cells, it becomes independent of the nerve for its differentiation and morphogenesis, but its individual cells remain nerve-dependent for mitosis [72,73]. Thus, a blastema denervated at the medium bud stage will form a morphologically normal but miniature regenerate owing to the lack of further mitosis. An established blastema mesenchyme stripped of its epidermis by chelation and denervated by implanting it in a dorsal fin tunnel such that its distal end becomes recovered by fin epidermis also forms a morphologically normal but miniature regenerate. Consistent with this result, the 3H-thymidine labeling and mitotic indices of epidermis-free newt limb blastemas cultured in the presence of dorsal root ganglia are reduced three- to fourfold [74]. If the mesenchyme is implanted completely into the fin tunnel, however, a miniature regenerate forms that is distally truncated [75]. This result suggests that the AEC has a role in proximodistal (PD) patterning in addition to proliferation. As in other vertebrate embryos, the amphibian limb bud requires signals from the AEC for outgrowth that are generated by a reciprocal interaction between mesenchyme and AEC. It becomes nerve-dependent for regeneration only after it has differentiated and become innervated [76]. Urodele limb buds rendered aneurogenic are AECdependent for growth but do not become nerve-dependent for regeneration [77,78], although this dependence can be instituted by allowing the limbs to become reinnervated [79]. These facts suggest a model for blastema growth in which the AEC supplies mitogens to subjacent blastema cells but requires factors supplied by the nerve to perform this function (Fig. 3.3). If this model outlined is correct, putative AEC and nerve factors should meet several criteria, in addition, to being expressed by the AEC and nerves. For AEC candidates, these are expression of their receptor in the blastema mesenchyme, loss of AEC expression by denervation, and the ability to support regeneration of denervated or AECdeprived limbs to digit stages. Neural factors should be transported from nerve cell bodies along limb nerve axons to the AEC where they bind to their receptor, denervation should prevent blastema cell mitosis by abolishing AEC factors, and they should support regeneration to digit stages in denervated limbs. Many factors expressed by the AEC stimulate blastema cell proliferation in vitro and in vivo. Fibroblast growth factor (Fgf)1, Fgf2, and the anterior gradient (AG) protein are expressed by the AEC in vivo [80,81], and mesenchymal cells express receptors for Fgfs and AG [81,82]. Fgf1 elevated the mitotic index of cultured blastema cells [83], and Fgf2 elevated the mitotic index of blastema cells in amputated limbs covered by full-thickness skin [84]. However, only two factors expressed by the AEC have been shown to be downregulated by denervation and to substitute for the nerve in supporting regeneration to digit stages. These are Fgf2 [85] and AG [81]. AG is involved in head development of the Xenopus embryo and has been the more thoroughly investigated of the two. It is strongly

AEC N

Stump

Blastema mes

FIGURE 3.3 Model of apical epidermal cap (AEC)eneural (N) interaction for mitotic growth of the blastema. The function of the AEC is to secrete mitogens into the subjacent blastema mesenchyme (mes) (green dashed line); this function depends on factors made by dorsal root ganglion neurons (solid yellow line). Dashed blue line indicates level of amputation.

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3. BLASTEMA FORMATION AND GROWTH

expressed in Schwann cells insulating the axons of regenerating newt limbs at 5 and 8 dpa, when initial dedifferentiation is under way. By 10 dpa, AG expression shifts to the gland cells of the AEC, coincident with formation of the accumulation blastema. Denervation abolishes AG expression, indicating that it is induced by axons. The AG gene supports regeneration to digit stages when electroporated into denervated newt limbs at 5 dpa. The receptor for AG is the blastema cell surface protein Prod1, a member of the Ly6 family of three-finger proteins anchored to the cell surface by a glycosylphosphatidylinositol linkage [86]. Conditioned medium of Cos7 cells transfected with the AG gene stimulates 5-bromo-20 -deoxyuridine incorporation into cultured blastema cells; this incorporation is blocked by antibodies to Prod1, which suggests that AG acts directly on blastema cells through Prod1 to stimulate proliferation [81]. The dependence on nerves for mitosis throughout blastema growth implies that this action is continuous, an implication that could be confirmed or refuted by examining expression patterns of AG in control and denervated limbs at successively later stages of blastema redevelopment. As would be predicted by the nerveeAEC model, AG is expressed in developing urodele limb buds and regenerating aneurogenic limbs [87]. Fgf2 expression in the AEC of aneurogenic limbs has not been examined. Factors expressed by DRG neurons that promote blastema cell proliferation in vitro include transferrin, substance P, Fgf2, and glial growth factor 2 (Ggf-2) [17,88]. Ggf-2 was reported to rescue regeneration to digit stages in denervated axolotl limbs when injected intraperitoneally during blastema formation [89], but this has not been investigated further. New nerve factor candidates are combinations of Fgf8 and bone morphogenetic protein. Both are expressed in DRG neurons and are detectable in peripheral limb nerve axons [90]. Furthermore, they can substitute for the nerve in inducing a supernumerary limb [91].

Interaction of Cells From Opposite Sides of the Limb Circumference Blastema cells also fail to persist unless they are derived from opposite positions on the limb circumference and interact. This was shown by experiments in which the normally asymmetrical (anteroposterior [AP] or dorsoventral [DV]) skin of the newt limb was made symmetrical by rotating a longitudinal strip of skin 90 degrees, grafting it around the circumference of an irradiated limb, and then amputating through the strip [92]. This led to failure of blastema cell mitosis, which indicates that mitosis requires an interaction between blastema cells with opposite positional identities. Normal regeneration ensued, however, when this requirement was met by grafting short longitudinal skin strips from two or more opposite points of the circumference. Lheureux [93] devised an experimental model to show the synergistic effect on mitosis of nerve-induced mitogen expression by the AEC with the interaction between cells of opposite positional identities. He made a wound on one side of an adult newt limb, deviated a nerve to the wound site, and juxtaposed a graft of skin from the opposite side of the limb to the skin of the wound. By themselves, nerve deviation or juxtaposing opposite positional identities resulted in the formation of blastema cells that failed to persist. Together, however, they stimulated the formation of a supernumerary blastema that grew and regenerated a complete limb. The Lheureux model was later used by another group to evoke supernumerary blastema formation in axolotl limbs under the name “Accessory Limb Model” [94]. Supernumerary limbs can be induced in the same way by reversing the AP or DV axis of the blastema with respect to the limb stump.

D Inject fgf8 baculovirus; wound

Inject shh baculovirus; wound

P

A

V FIGURE 3.4 Cross-section of axolotl stylopodium and result of injecting baculovirus constructs containing Fgf8 under posterior skin and Shh

under anterior skin, followed by wounding and nerve deviation. Fgf8 and Shh substitute for anterior and posterior skin, respectively, in evoking supernumerary limb formation. A, anterior; D, dorsal; P, posterior; V, ventral.

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The Lheureux system has been used to investigate several aspects of limb regeneration. One of these has been to define the signals passing between cells of opposite positional identity. Nacu et al. [95] injected the sonic hedgehog (shh) gene under anterior skin of the stylopodium and the fgf8 gene under posterior skin using a baculoviral vector system (Fig. 3.4). Shh and Fgf8 have been implicated in AP patterning of the limb bud and regeneration blastema [96,97]. A wound was then created in the skin and a nerve deviated to the site. The gene transfections substituted for cells of opposite positional identity, which indicated that Shh and Fgf8 signal opposite cells to drive some aspect of mitosis. The AEC does not persist in the absence of polar juxtaposition, which suggests that proliferating blastema cells may also be essential for AEC maintenance. Positional identity in the PD axis of the limb is associated with a proximal (low) to distal (high) gradient of cell adhesion [98e100]. Prod-1 is expressed in an opposite gradient; antibodies to Prod1, or its removal from the blastema cell surface by phosphatidylinositol-specific phospholipase C inhibit the recognition of adhesive differentials between distal and proximal blastemas [86]. These results suggest that Prod-1 has a role in integrating mitosis with patterning through its AG ligand [101]. The PD segments of the regenerating limb appear to be specified in a proximal to distal direction [102], but how the positional identities reflected in cell surface adhesion are generated is not clear. There is evidence that PD patterning of the chick limb bud switches from an extrinsic signaling mechanism to an intrinsic timing mechanism that is reflected in intrinsic changes in cell affinities in the distal limb bud mesenchyme [103]. Whether blastema patterning uses a similar mechanism is unknown but would be worth investigating.

List of Acronyms and Abbreviations AEC Apical epidermal cap AG Anterior gradient AP Anteroposterior DRG Dorsal root ganglion DV Dorsoventral ECM Extracellular matrix GFP Green fluorescent protein IP3 Inositol triphosphate MMP Matrix metalloproteinase PD Proximodistal PKC Protein kinase C Shh Sonic hedgehog TIMP Tissue inhibitor of metalloproteases

Glossary Aneurogenic limb A limb that develops without innervation. Blastema The collection of undifferentiated cells that forms after amputation of a limb. Dedifferentiation The process by which differentiated cells revert to progenitor cells. Fibroblastema A blastema composed of fibroblastic cells instead of dedifferentiated cells. Growth factors Proteins that promote cell proliferation and differentiation. Histolysis The breakdown of extracellular matrix to release cells. Satellite cell The progenitor cells that regenerate injured muscle. SMAD Proteins phosphorylated by activation of TGF-b/BMP receptors that participate in activating or repressing transcription. Stem cell Undifferentiated cells that maintain or repair tissue structure. Transdifferentiation The differentiation of one cell type into another cell type.

Acknowledgments Research from this laboratory was supported by the W.M. Keck Foundation and the US Army Research Office (Grant number W911NF07-10176).

References [1] Thornton CS. Amphibian limb regeneration. In: Brachet L, King TJ, editors. Advances in morphogenesis, vol. 7. New York: Academic Press; 1968. p. 205e44. [2] Morrison JI, Loof S, He P, Simon A. Salamander limb regeneration involves the activation of a multipotent skeletal muscle satellite cell population. J Cell Biol 2006;172(3):433e40. [3] Hay ED, Fischman DA. Origin of the blastema in regenerating limbs of the newt Triturus viridescens. An autoradiographic study using tritiated thymidine to follow cell proliferation and migration. Dev Biol 1961;3:26e59.

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[4] Repesh LA, Furcht LP. Distribution of fibronectin in regenerating limbs of the adult newt Notophthalmus viridescens. Differentiation 1982;22: 125e31. [5] Singer M, Salpeter MM. Regeneration in vertebrates: the role of the wound epithelium in vertebrate regeneration. In: Zarrow M, editor. Growth in living systems. New York: Basic Books; 1961. [6] Tassava RA, Castilla M, Arsanto J-P, Thouveny Y. The wound epithelium of regenerating limbs of Pleurodeles waltl and Notophthalmus viridescens: studies with mAbs WE3 and WE4, phalloidin, and DNase 1. J Exp Zool 1993;267(2):180e7. [7] Onda H, Tassava RA. Expression of the 9G1 antigen in the apical cap of axolotl regenerates requires nerves and mesenchyme. J Exp Zool 1991;257(3):336e49. [8] Ferretti P, Brockes JP, Brown RA. A newt type II keratin restricted to normal and regenerating limbs and tails is responsive to retinoic acid. Development 1991;111(2):497e507. [9] Borgens RB, Vanable Jr JW, Jaffe LF. Bioelectricity and regeneration: large currents leave the stumps of regenerating newt limbs. Proc Natl Acad Sci USA 1977;74(10):4528e32. [10] Adams DS, Masi A, Levin M. Hþ pump-dependent changes in membrane voltage are an early mechanism necessary and sufficient to induce Xenopus tail regeneration. Development 2007;134(7):1323e35. [11] Gorsic M, Majdic G, Komel R. Identification of differentially expressed genes in 4-day axolotl limb blastema by suppression subtractive hybridization. J Physiol Biochem 2008;64(1):37e50. [12] Jenkins LS, Duerstock BS, Borgens RB. Reduction of the current of injury leaving the amputation inhibits limb regeneration in the red spotted newt. Dev Biol 1996;178(2):251e62. [13] Rao N, Jhamb D, Milner DJ, Li B, Song F, Wang M, et al. Proteomic analysis of blastema formation in regenerating axolotl limbs. BMC Biol 2009;7:83. [14] Lodish H, Berk A, et al. Molecular cell biology. New York: Freeman, W.E. and Co; 2008. [15] Tsonis PA, English D, Mescher AL. 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Identification and characterization of newt rad (ras associated with diabetes), a gene specifically expressed in regenerating limb muscle. Dev Dynam 2001;220(1):74e86. [32] Lundkvist J, Lendahl U. Notch and the birth of glial cells. Trends Neurosci 2001;24(9):492e4. [33] King BL, Yin VP. A conserved micro RNA regulatory circuit is differentially controlled during limb/appendage regeneration. PLoS One 2016. https://doi.org/10.1371/journal.pone.0157196. [34] Takahashi K, Tanabe K, Ohnuki M, Ichisaka T, Tomoda K, Yamanaka S. Induction of pluripotent stem cells from adult human fibroblasts by defined factors. Cell 2007;131(5):861e72. [35] Yu J, Vodyanik MA, Smuga-Otto K, Antosiewicz-Bourget J, Frans JL, Tian S, et al. Induced pluripotent stem cell lines derived from human somatic cells. Science 2007;318(5858):1917e20. [36] Maki N, Suetsugu-Maki R, Tarui H, Agata K, Del Rio Tsonis K, Tsonis PA. Expression of stem cell pluripotency factors during regeneration in newts. Dev Dynam 2009;238(6):1613e6. [37] Mescher A, White GW, Brokaw JJ. Apoptosis in regenerating and denervated, nonregenerating urodele forelimbs. Wound Repair Regen 2000;8(2):110e6. [38] Atkinson D, Stevenson TJ, Park EJ, Reidy MD, Milash B, Odelberg SJ. Cellular electroporation induces dedifferentiation in intact newt limbs. Dev Biol 2006;291(1):257e71.

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C H A P T E R

4 The Molecular Circuitry Underlying Pluripotency in Embryonic and Induced Pluripotent Stem Cells Rachel H. Klein, Paul S. Knoepfler University of California Davis, Davis, CA, United States

INTRODUCTION Multiple criteria are employed to characterize cellular pluripotent potential, including (1) the expression of molecular markers, in particular, transcription factors known to regulate embryonic stem cell (ESC) potency and selfrenewal; (2) the absence of molecular and morphological markers defining specific lineages that are typically referred to as “differentiation-associated genes”; and (3) the ability to form all three embryonic germ layers including ectoderm, endoderm, and mesoderm upon induction of differentiation in vitro or in vivo. Upon injection into immunocompromised mice, ESCs and induced pluripotent stem (iPS) cells will rapidly and reproducibly form teratomas containing differentiated cells from the three germ layers. Ultimately, the most rigorous pluripotency assay, but one not currently amenable to studying human cells, for ethical reasons, is that upon implantation of ESCs into blastocysts there is a subsequent contribution of these cells to all tissue types of the adult chimeric animal. Although some stem-like cells have been isolated from many organisms, including zebrafish, dogs, chickens, mice, rats, and humans, to date only stem cells from a few species have demonstrated this ultimate and most stringent capability associated with pluripotency by which they can be defined concretely as ESCs. In this review, we describe the mechanistic details of the molecular circuitry that regulates the maintenance of the pluripotent state at the level of signal transduction, chromatin dynamics, and transcription factor control. We also discuss the reprogramming of somatic cells into iPS cells that possess essentially all ESC-like properties, and how the processes that govern their production and maintenance relate to ESC maintenance. Analyses of the rewiring of the molecular circuitry involved in going from a somatic to pluripotent state have reinforced the identities and roles of core pluripotency mechanisms.

GROUND STATE AND PRIMED EMBRYONIC STEM CELLS HAVE UNIQUE SIGNALING NETWORKS UNDERLYING PLURIPOTENCY The initial protocols employed to isolate and maintain murine ESCs involved plating inner-cell mass cells onto a feeder cell layer of embryonic fibroblasts in a medium containing serum proteins [1,2]. The complex mixture of exogenous factors released by fibroblasts into the medium maintains ESCs in their pluripotent state and allows for the undifferentiated self-renewal and proliferation of these cells. Despite the complex composition of fibroblastconditioned medium containing serum components, several key growth factor signaling pathways essential for subsequent pluripotency have been identified and extensively characterized, including leukemia inhibitory factor (LIF)/signal transducer and activator of transcription 3 (Stat3), and bone morphogenic protein (BMP) signaling. Upon removal of the feeder cells, or medium conditioned by the feeder cells, ESCs spontaneously differentiate into all three germ layers of the developing organism. More recent protocols have been developed and used Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00004-7

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alternative media options, including 2i and 3i media that allow for feeder-independent growth of mouse ESCs (mESCs) [3]. Development of these media was based on the idea that inhibition of intrinsic differentiation signals could maintain pluripotency, similar to the application of exogenous factors (from fibroblast feeders and serum). For instance, 2i medium uses an inhibitor of mitogen activated protein kinase (MAPK)/extracellular signale regulated kinase (ERK) (MEK) combined with an inhibitor of glycogen synthase kinase 3b (Gsk3b) and does not require LIF for stem cell maintenance [4]. Human ESCs (hESCs) can be grown on feeder cells in media conditioned by fibroblasts or in chemically defined media [5]. Interestingly, studies with these human cells revealed striking differences in signaling pathway requirements for stem cell maintenance compared with mESCs, which led to the hypothesis that despite the high levels of conservation in development across widely divergent species, pluripotent stem cell maintenance is at least partially a unique process in different organisms. This line of thinking has been challenged by more recent findings that distinguish two main states of ESCs: naive and primed. It was found that under commonly used derivation and culture conditions, mESCs are maintained in the naive state whereas hESCs acquire and maintain the primed state [6]. Primed mESCs, which are derived from the mouse embryo after implantation and are typically called epiblast stem cells (EpiSCs), share many of the same signaling pathway requirements as do hESCs grown under standard conditions [7]. Several articles reported the development of media and conditions that allow for culturing of naive hESC [8,9] and found in this context that the pathway requirements are much more similar to those required by naive mESC, including the dependency on LIF. These findings suggest LIF signaling is a key aspect of naive pluripotent stem cell self-renewal across species. Thus, apparent differences in human and murine ESCs may be largely a function of the developmental state related to how the cells are cultured, which signaling pathways are modulated in the process, and thus which cell fateerelated transcription factors are activated or repressed downstream.

INDUCED PLURIPOTENT STEM CELLS An additional type of pluripotent stem cells, iPS cells, can be derived from human or mouse somatic cells through related protocols mainly based on exogenous expression of a select group of pluripotency-related transcription factors. In 2006, Yamanaka and colleagues reported that by introducing four transcription factors necessary for ESC self-renewal including Oct4, Sox2, Klf4, and c-Myc into the genome of mouse fibroblasts, some cells underwent complete reprogramming to a state of pluripotency [10]. Further work including studies by Yu et al. found that Klf4 and c-Myc can be substituted by Nanog and another transcription factor Lin28 [11], or that the entire reprogramming process in humans and mice can be accomplished through the expression of a single cluster of microRNAs (miRNAs), mir302/367 [12]. These results indicate that multiple combinations of transcription factors or other regulators can reprogram cells to the same primitive developmental state [11]. A year later, Yamanaka and other teams reported successful human somatic cell reprogramming to make iPS cells [13]. These human iPS cells possess all of the hallmarks of hESCs in their functional abilities to differentiate into all cell types of an organism [10]. Importantly, these distinct pluripotent cells also require the same signaling pathways to maintain their undifferentiated state and respond appropriately to growth factors and cytokines eliciting specific lineages [14]. Reprogramming methods have been developed including the use of nongenetic methods that do not leave a genomic fingerprint via episomal vectors, nonintegrating virus, messenger RNAs (mRNAs), proteins, and even in an entirely chemical cocktail [15e17]. Although the reprogramming protocols are well-established, large fluctuations remain in efficiency and the probability that any given cell will become an iPS cell. Even among a clonally selected somatic cell population infused with the same copy number of reprogramming factors, heterogeneity abounds, with a minority of cells reprogramming within a few weeks while others require longer periods [18]. Part of the answer to the heterogeneity may reside in the concept of nongenetic heterogeneity and random fluctuations in protein expression levels among a clonal population of cells; heterogeneity in any number of transcription factors, signaling proteins, or epigenetic marks may place a cell in a state either relatively amenable or resistant to proper manipulation by the reprogramming factors.

LEUKEMIA INHIBITORY FACTOR AND BONE MORPHOGENIC PROTEIN SIGNALING PATHWAYS REGULATE MOUSE EMBRYONIC STEM CELL SELF-RENEWAL mESCs derived from permissive strains can maintain their undifferentiated state with a combination of signaling from LIF and BMPs (Bmp4) [19,20]. LIF receptor activation leads to receptor dimerization with gp130 subunits and subsequent tyrosine phosphorylation and nuclear localization of the transcriptional activator Stat3 (Fig. 4.1) [21].

LEUKEMIA INHIBITORY FACTOR AND BONE MORPHOGENIC PROTEIN SIGNALING PATHWAYS REGULATE MOUSE EMBRYONIC

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FIGURE 4.1 Signaling circuitry regulating mouse and human embryonic stem cell (ESC) pluripotency. The Wnt pathway is a highly conserved regulator of pluripotency and is active in both mouse and human ESCs. Different signaling pathways are required to maintain pluripotency, depending on the state of the stem cells and species. Naive mESCs require leukemia inhibitory factor (LIF) and bone morphogenic protein (BMP) signals to maintain self-renewal, whereas primed mESCs require transforming growth factor b (TGFb) and fibroblast growth factor (FGF) signaling. Naive human ESCs (hESCs) have been derived using various media formulations, but in general they require LIF, FGF, TGFb, and WNT signaling. Primed hESCs depend on the activity of TGFb and FGF signals. These pathways ultimately function at multiple levels to maintain the pluripotent state by inhibiting differentiation and feeding into the core transcriptional regulatory circuitry of ESCs. ALK, anaplastic lymphoma kinase; EpiSC, epiblast stem cells; ERK, extracellular signaleregulated kinase; GSK, glycogen synthase kinase; MAPK, mitogen activated protein kinase; MEK, MAPK/ERK; mESC, mouse ESC; STAT, signal transducer and activator of transcription; TCF, transcription factor

Whereas LIF also activates the phosphatidylinositol-3-kinase/protein kinase B and MAPK signaling pathways in mESCs, only Stat3 activation is required for pluripotency [22]. One of the key downstream targets of LIF/Stat3 is the transcription factor Myc (Fig. 4.1) [23]; high levels of Myc expression even in the absence of LIF signaling allows for self-renewal of mESCs [23], which suggests that the main downstream target of the LIF/Stat3 signaling pathway in mESC is c-Myc.

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BMPs are transforming growth factor b (TGFb) superfamily members that bind to type I TGFb receptors anaplastic lymphoma kinase (ALK)1, ALK2, ALK3, or ALK6 (Fig. 4.1). Upon ligand binding, type I receptors form heterodimers with type II receptors, which recruit and phosphorylate receptor activated SMADs 1, 5, and 8 (R-SMADs). Serine/threonine phosphorylation of R-SMADs allows association and complex formation with SMAD4, which can subsequently enter the nucleus and initiate transcription (Fig. 4.1) [24]. Among the targets of BMP signaling is Inhibitor of Differentiation 1 (Id1), a key factor that can maintain ESC self-renewal when overexpressed even in the absence of BMP signaling [19]. In addition, chromatin immunoprecipitation (ChIP)-sequencing (Seq) studies revealed that Smad1 colocalizes with octamer binding protein 4 (Oct4), SRY-box 2 (Sox2), Nanog, and Stat3 at a number of pluripotency gene targets [25]. Whereas ERK has generally been considered to promote differentiation of mESCs (it promotes Kruppel-like factor 4 (Klf4) ubiquitination and degradation [26] and reduces Nanog’s activating ability [27], and inhibitors of upstream factor MEK are used to maintain mESC pluripotency), it has also been shown to have an important role in maintaining genomic stability, promoting proliferation, and suppressing apoptosis in mESCs [28]. This suggests a more complex role for this factor in stem cell dynamics.

TRANSFORMING GROWTH FACTOR b AND FIBROBLAST GROWTH FACTOR SIGNALING PATHWAYS REGULATE HUMAN EMBRYONIC STEM CELL SELF-RENEWAL As discussed earlier, hESCs cultured on fibroblast feeder cells with serum exist in a more primed state, with similarities to mouse EpiSCs. LIF/STAT3 signaling does not stimulate self-renewal in this state; rather, TGFb/activin and fibroblast growth factor (FGF) signaling are required to maintain hESC pluripotency [14,29]. In contrast to naive-state mESCs, BMP signaling actually promotes hESC differentiation [30,31]. TGFb and activins account for the second branch of the TGFb superfamily of ligands, and binding to receptors ALK4, ALK5, and ALK7 triggers serine/threonine phosphorylation of the C-terminal region of SMAD2 and SMAD3, which also dimerize with SMAD4 to allow nuclear entry and transcription (Fig. 4.1) [24]. FGFs function through tyrosine receptor dimerization upon ligand binding and subsequent activation of phosphorylation events in the MAPK cascade (Fig. 4.1) [32]. FGF signaling can further phosphorylate both BMP and TGFb-mediated R-SMADs at the “linker” domain of the proteins. This phosphorylation has been associated with signal termination, because linker phosphorylation allows recognition of SMAD proteins by the ubiquitin ligase SMURF1 [33,34]. Polyubiquitination of the SMAD proteins by SMURF1 leads to subsequent degradation of the SMAD proteins and termination of the signal. Hence, an intricate balance of antagonistic signaling inputs may exist in the maintenance of hESC pluripotency. Among their definitive roles in proliferation and survival, FGF signals may also act to inhibit differentiation promoting BMP signals in hESCs by promoting degradation of any active SMAD 1/5/8 proteins [33]. Alternatively, FGF signals may also fine-tune the amount of active TGFb-mediated SMAD 2/3 proteins to produce the proper threshold of activity necessary to maintain pluripotency, because an excess of TGFb/activin signaling can lead to definitive endoderm differentiation of ESCs [35]. Studies demonstrating the necessity of these pathways to maintain self-renewal have followed two strategies. First, small molecule inhibition of TGFb/activin receptors results in the rapid differentiation of hESCs even in fibroblastconditioned medium, which illustrates the necessity for TGFb signals to maintain pluripotency [29]. Second, defined medium with select growth factors and cytokines has been developed to substitute fibroblast-conditioned medium, which contains a diverse milieu of undefined components. These studies revealed that TGFb or activin plus FGF-2 at defined concentrations is a necessary component for self-renewal, and removal of either of these factors results in differentiation of ESCs [14,36]. For both mouse and human iPS cells, the same media and growth factor pathway activities are required for culture and maintenance of their pluripotency as for the respective comparable types of ESCs, which further supports the important nature of these pathways in pluripotency.

WNT SIGNALING CONTRIBUTES TO MAINTENANCE OF PLURIPOTENCY IN MOUSE EMBRYONIC STEM CELLS AND TO THE NAIVE HUMAN EMBRYONIC STEM CELL STATE Although activation of LIF and BMP signaling are sufficient to maintain mESC self-renewal, the highly conserved Wnt pathway is also an important modulator of these signaling pathways; it can contribute to the maintenance of mESC pluripotency [37e39]. In the presence of Wnt ligand, a receptor complex forms between receptors Frizzled

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and Lrp5/6. This complex recruits and sequesters Axin and Gsk3b, releasing their inhibitory interaction with b-catenin, which subsequently can accumulate in the nucleus, where it serves as a coactivator to activate Wnt-responsive genes (Fig. 4.1) [40]. b-Catenin also interacts with chromatin remodeling factor Brg1 to remodel chromatin during gene activation in response to Wnt signaling [41]. b-Catenin is important for maintaining stem cell identity and inhibiting neuronal differentiation, because mESCs that lack this factor show decreased expression levels of stem cell genes Dppa4, Dppa5, and Rex1, and increased expression of Oct6, a marker of neuronal differentiation [42]. Functional studies demonstrating the importance of Wnt signaling have employed small molecule inhibitors of Gsk3b, which destabilizes b-catenin. Inhibition of Gsk3b results in increased Wnt activity, and cells cultured in the presence of Gsk3b inhibitors have increased propensity to maintain their pluripotent state even under differentiation conditions [19,37]. Furthermore, the role of Wnt ligands in supporting stemness has been demonstrated in experiments that show that Wnts secreted by feeder cells or Wnt-conditioned media maintain pluripotency in mESCs [38,39]. Extensive cross-talk has been documented between the LIF and Wnt signaling pathways; when LIF is removed from culture media, mESCs reduce levels of nuclear b-catenin [43]. In addition, Wnt signaling molecules Wnt3a, Wnt5a, and Wnt6 can upregulate Stat3 signaling (downstream of LIF), and upregulation of Wnt signaling can compensate for low levels or the complete absence of LIF; LIF presence can also compensate for the loss of b-catenin [38]. Because b-catenin knockout mESCs have been shown to be able to self-renew in the presence of LIF, it appears not to be absolutely essential to the maintenance of pluripotency; rather, it acts in concert with other factors to enhance stem cell self-renewal. This idea is supported by the finding that Wnt signaling is also important in reprogramming differentiated cells into iPS cells, because Wnt3a can improve reprogramming efficiency [44]. Transcription factor 3 (Tcf3) is a transcription factor and binding partner for b-catenin. It has been shown to bind to many pluripotency genes and act as a repressor. It is also commonly found in complex with Oct4, Sox2, and Nanog, where it is thought to have a role in tempering the level of activation by these factors at pluripotency genes [45,46]. Upon stimulation of Wnt signaling by Wnt3a in mESCs, b-catenin enters the nucleus and binds to Tcf3, promoting its phosphorylation and degradation, thereby relieving the repression of at least a subset of Tcf3 targets and promoting stem cell self-renewal [47,48]. Seemingly conflicting reports exist on the role of Wnt signaling in promoting differentiation of mESCs and in maintaining pluripotency. These differences can potentially be explained by the vast number of transcriptional regulators that can associate with b-catenin. In addition to the TCF family, b-catenin can bind with p300, cyclic adenosine monophosphate response element binding protein (CBP), the histone methyltransferase Mll1, and numerous transcription factors of the SOX, SMAD, FOXO, and nuclear receptor families [49]. Some work suggested that although b-catenin can form a complex with p300 and with CBP, the b-catenin/CBP complex is specifically essential for mESC maintenance, whereas p300 is dispensable, and that the complex specificity can determine the activity of b-catenin [50,51]. An interesting open question is how complex specificity itself is controlled. The mechanisms of Wnt signaling action on cell fate appear to be highly time- and context-dependent. For example, Wnt signaling activated early in mESC differentiation leads to the promotion of mesoderm development toward heart; however, the same Wnt signaling activated at a later stage of differentiation actually inhibits this fate [52]. Many of the conflicting ideas about Wnt function in differentiation and in pluripotent cells may also arise from cross-talk and interactions of Wnt with different signaling pathways that are selectively active either in ESCs or in a lineage differentiation context influenced by temporal factors. The role of Wnt signaling in hESCs has not been fully elucidated, partly because most work has focused on hESCs under standard culture conditions, which promote a more primed cell state. Studies of cells under these conditions indicated that Wnt signaling might not be essential for hESC maintenance. However, a study looking specifically at hESCs in the naive state concluded that whereas Wnt signaling was not required for the expression of pluripotency factors in this context, it contributes to cell proliferation and colony-forming ability, and that blocking Wnt signaling led to the acquisition of a metabolic and transcriptional profile more similar to hESC in the primed state, which suggests that the role of Wnt signaling in naive hESCs is to prevent transition to the primed state [53].

THREE TRANSCRIPTION FACTORS, OCTAMER BINDING PROTEIN 4, SRY-BOX 2, AND NANOG, FORM THE CORE PLURIPOTENCY TRANSCRIPTIONAL NETWORK The previously discussed growth factor signaling pathways that regulate pluripotency could do so through a number of different mechanisms, as outlined earlier. The most widely accepted model is that they function coordinately to inhibit differentiation-associated factors and induce expression of pluripotency-related transcription

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factors. Among the many transcription factors identified and characterized as having critical roles in stem cells, a core set of three pluripotency transcription factors has been shown to form the backbone of the pluripotency gene regulatory network. These three, OCT4, SOX2, and NANOG, act together downstream of the key signaling pathways and function in part through positive feedback loops to maintain each other’s expression as well as that of the wider pluripotency gene expression program. The three are thought to be central to transcriptional regulation of ESC identity because of their essential roles during early development, including in the inner cell mass (ICM), and their ability to maintain the ESC state [6,54e56]. Disruption of Oct4 in knockout embryos and ESCs results in the inappropriate differentiation of ICM and ESC to trophectoderm, whereas Nanog mutants develop into extraembryonic endoderm [6,55,56]. Sox2 loss-of-function mutants also divert to trophectoderm [54]. Notably, the phenotype of mESCs overexpressing Oct4 resembles that of Nanog loss of function, forming embryonic endoderm, whereas cells with Nanog overexpression are highly resistant to differentiation [55]. These findings suggest a higher level of complexity in core pluripotency transcription factor functional interactions where they do not always simply work together coordinately to promote a blanket state of pluripotency and in which they can sometimes be antagonistic in specific contexts. Furthermore, they also highlight the need for exquisite fine-tuning of gene expression circuitry to promote pluripotency without completely blocking eventual differentiation capacity. Nevertheless, genome-wide analysis has revealed that these three transcription factors generally form an autoregulatory network and do so by binding to each other’s promoter regions and enhancing each other’s expression [57]. Furthermore, these factors regulate the expression of large numbers of downstream genes governing aspects of differentiation, cell cycle, and self-renewal [57]; and as mentioned earlier, these core factors are integral to the reprogramming process to make iPS cells.

MYC LINKS CELL SIGNALING TO PLURIPOTENCY GENE REGULATION In addition to core pluripotency transcription factors, c-Myc was also found by Yamanaka initially as one of the four main reprogramming cocktail factors. Whereas c-Myc was subsequently determined not to be essential for reprogramming, it can potentiate efficiency of the process by one or two orders of magnitude [58]. In stem cells, MYC factors have a broader role in regulating many unique properties associated with stemness and pluripotency, including specific states of metabolism, cell signaling, cell cycle dynamics, and the epigenetic landscape. Because MYC can regulate all of these cellular functions (even if in distinct ways) in somatic and progenitor cells, where it is also expressed, it is probably not an actual “pluripotency factor” in the narrowest sense of that term. Three MYC family members have been identified and characterized: MYC, MYCN, and MYCL. In stem cells, there is a high level of redundancy between MYC and MYCN; however, knockout of both causes a loss of pluripotency and promotes differentiation [59]. Reportedly, loss of Mycl causes no detectable defects in stem cells and does not appear to be able to compensate for loss of the other two MYC family members [60]. MYC represents one connection point between cell signaling and the core pluripotency transcriptional network. Stem cell maintenance requires active Wnt signaling and repression of high levels of MAPK signaling. When MYC is expressed, it acts to suppress MAPK signaling by transcriptionally upregulating the ERK inhibitors, dual-specific phosphatases (DUSP2/7); it also acts to promote Wnt signaling through upregulation of WNT receptors and suppression of WNT antagonists. In mESCs, LIF is also required to activate the JAK/Stat3 pathway with c-Myc as a key downstream target that is activated by LIF [23]. MYC has been shown to act as both a transcriptional activator and a repressor in hESCs. In ChIP experiments, MYC is primarily associated with active histone modifications, but it also has an important role in promoting pluripotency through repressing the expression of differentiation genes including the HOX gene clusters in hESCs [61]. This is consistent with findings that loss of MYC does not have a strong effect on the expression of core pluripotency genes; the most dramatic effects instead manifest as upregulation of early differentiation genes [59]. Similarly, during reprogramming, MYC may act to suppress differentiation-specific gene expression [62]. In the repressive context in hESCs, MYC interacts with MIZ1 and repressive chromatin-modifying complexes including DNMT3A, and histone deacetylases (HDACs) [61]. The interaction between MYC and MIZ-1 appears to be antagonistic: whereas the two co-bind many targets, their effects on transcription are often in opposite directions [61]. This opposing relationship may serve to balance the contradictory forces of stem cell self-renewal and differentiation; when external signals alter the balance of MYC and MIZ-1, one force can dominate and either promote commitment to differentiation or maintain pluripotency.

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MYC also influences stem cell potency in part through regulating the cell cycle, where it acts to promote S-phase progression [59]. Among MYC targets are numerous cyclins and cyclin-dependent kinase enzymes, as well as the miRNA cluster miR-17-92 [63], which targets and suppresses cell cycle inhibitors including those in the retinoblastoma family [25,64]. Because ESCs have unique cell cycle properties, it has been suggested that the effect on the cell cycle and proliferation upon loss or downregulation of MYC could trigger ESC differentiation [59]. Numerous sources of evidence indicate that whereas MYC is frequently found at gene promoters with the core pluripotency factors, some of its mechanisms of action are unique compared with other transcription factors. In fact, ChIP-microarray (ChIP-chip) and ChIP-Seq studies generally point to unique modules of gene targets for MYC compared with core pluripotency factors such as OCT4, SOX2, or NANOG. MYC also has an important role in releasing transcriptional pausing of RNA polymerase II and the transcription machinery at gene promoters, a function that is unique from core pluripotency transcriptional regulators [65]. MYC has been found to interact with a number of different epigenetic modifying enzymes in both activating and repressing conditions. These include histone acetyltransferases (HATs), HDACs, DNA helicases RUVBL-1 and -2, which are essential for ESC morphology, and histone demethylase LSD1 [66]. In addition to a role in repressing lineage specific and differentiation factors, MYC has been shown to regulate global levels of euchromatin in a number of cell types, an uncommon function for a transcription factor. This may relate to the proposed model in which MYC acts as a general activating factor at certain types of gene promoters that are already primed for gene expression by binding specific transcription factors. In this context, MYC has been shown to interact with a number of HAT complexes to promote histone H3 and H4 acetylation. Although MYC is not generally considered a pioneer factor that can bind and remodel chromatin, it has an important and early job during reprogramming to promote euchromatin by regulating global levels of histone modifications. MYC frequently binds to regions where OCT4, SOX2, or KLF4 are already bound, and then increases local euchromatin formation and transcriptional activity [67]. Thus, overall, the contributions of MYC to pluripotency are pleiotropic and mediated by a number of distinct epigenetic and transcriptomic mechanisms. MYC’s importance in both stem cells and cancer further points to specific downstream pathways shared among these cell types. For instance, inhibition of differentiation-related genes, maintenance of rapid cell cycling, and an elevated metabolic state are also features of tumorigenesis, and tumorigenesis and reprogramming appear to be related phenomena [68].

A SPECIFIC EPIGENETIC PROGRAM HELPS MAINTAIN PLURIPOTENCY The importance of chromatin modification states in regulating pluripotency is evidenced by the fact that perturbations of the expression level or function of many different epigenetic enzymes can impair stem cell self-renewal and lead to the loss of pluripotency, events that are followed by differentiation [69e72]; these enzymes are necessary for specific chromatin structural and functional states. In addition, modulation of epigenetic enzyme activity can promote cellular reprogramming. For example, the use of HDAC inhibitors, or the H3K9 methyltransferase inhibitor G9a, makes reprogramming more efficient, implicating H3K9 methylation in the process [73,74]. Other histone marks and histone-modifying enzymes are implicated as well. For example, the polycomb group complex polycomb repressive complex 2 (PRC2) is responsible for the H3K27me3 mark that acts to repress gene expression; in ESCs, this mark is found in many differentiation genes and prevents premature expression (Fig. 4.2) [75]. Loss of PRC2 complex components causes a dramatic reduction in H3K27me3 and impairs pluripotency by upregulating differentiation genes [69]. PRC2 also has a role in maintaining bivalent domains, characterized by the presence of both active H3K4me3 and repressive H3K27me3 marks together in discrete chromatin domains in developmental genes that allow for rapid activation upon commitment of the pluripotent cell to differentiation (Fig. 4.2). Heterochromatin, marked by H3K9me3, is generally less common in ESCs than in differentiated cells. Consistent with this, the H3K9 methyltransferase G9a is important for early differentiation and development, and has been shown to recruit DNA methyltransferase enzymes Dnmt3a and Dnmt3b to downregulate pluripotency genes including Oct4 (Fig. 4.2) [76,77]. Conversely, enzymes that remove the H3K9 methylation marks are important for maintaining pluripotency. These include Jumanji domain enzymes Jmjd1a (Kdm3a) and Jmjd2c (Kdm4c), which are upregulated by Oct4 and function to remove H3K9 methylation to maintain the pluripotency of ESCs [71]. A number of interesting differences exist between DNA methylation levels and patterns in pluripotent cells compared with differentiated cells. For example, in looking at cytosine guanine (CpG) methylation, differentiated cells frequently have enrichment of methylation in sites where there is a high density of CpG dinucleotides. In contrast, in pluripotent cells, higher methylation levels are observed in regions with lower concentrations of CpG dinucleotides [78]. These low-density CpG regions that are methylated in pluripotent cells are associated with

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FIGURE 4.2 Chromatin dynamics of pluripotency and stem cell differentiation. In pluripotent cells, lineage genes are maintained in two

states: repressed by polycomb complexes and DNA methylation or poised for rapid activation based on histone modifications mediated by polycomb and trithorax complexes. Pluripotency genes are maintained by octamer binding protein 4, SRY-box 2, and Nanog complex binding, as well as recruitment of active chromatin-modifying complexes. Upon differentiation, a specific subset of lineage genes is activated in each cell type by lineage-specific transcription factors that recruit activating chromatin complexes. Pluripotency genes are repressed by DNA methylation and heterochromatin associated histone marks. HAT, histone acetyltransferase; OSKM, gene expression of octamer binding protein 4, SRY-box 2, Kruppel-like factor 4, and MYC; OSN, octamer binding protein 4, SRY-box 2, and Kruppel-like factor 4; PRC2, polycomb repressive complex 2; TrX, trithorax; Wdr5, WD repeat domain 5.

bivalent chromatin marks [79] and are thought to function by being able to respond rapidly and dynamically to signals that modulate chromatin remodeling, leading to efficient activation of genes with low-density CpG methylation in their promoters. There are three DNA methyltransferase enzymes: DNMT1, DNMT3A, and DNMT3B. DNMT1 is mainly involved in maintaining methylation over DNA replication events, whereas DNMT3A and DNMT3B modulate de novo methylation events. Deletion of any one Dnmt enzyme in mice results in embryonic or neonatal lethality [80e82], which indicates an important role in development, but mESCs deleted for one or all three Dnmt enzymes retain their pluripotency, despite losing DNA methylation globally. This suggests that in mice, DNA methylation is not required to maintain pluripotency, at least in vitro [83]. In hESCs grown under standard primed conditions, loss of both DNMT3 enzymes does not affect stem cell self-renewal, but loss of DNMT1A results in cell death [84]. These findings may indicate differences in the role of DNA methylation in maintenance between human and mouse ESCs, or it may reflect the fact that hESCs are usually grown under more primed conditions, whereas mESCs are maintained in a more naive state. DNMT3A and 3B are also not strictly required to generate iPS cells from differentiated cells [85]. In addition to the three DNMT enzymes, DNMT3L is a catalytically inactive DNMT protein that is highly expressed in mESCs and hESCs [86]. Its high expression in this context led to its inclusion in the 24 factors originally tested by Yamanaka for their reprogramming ability in differentiated cells. Although it does not possess enzymatic activity in itself, DNMT3L has been shown to dimerize with DNMT3A and significantly increase its enzymatic

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activity [87]. DNMT3L can also interact with the N terminal tail of histone H3, and several studies have suggested that it can discriminate among different H3K4 methylations [88,89]. DNMT3L knockdown alters the methylation landscape in ESCs and affects differentiation, but it does not appear to have a strong effect on pluripotency gene expression or stem cell self-renewal [90]. DNA demethylation can occur through a series of steps in which the 5mC is converted to 5-hydroxymethylcytosine through an oxidation reaction and then further modified to 5-formylcytosine and finally 5-carboxylcytosine, which can be removed by the cell through base excision repair [91]. The methylation intermediate 5fC is often found in gene regulatory elements, including poised enhancers, and in regions associated with p300 binding [92], which suggests a regulatory function for these methylation intermediaries. The main steps of this demethylation reaction are carried out by the TET family of enzymes. TET family members Tet1 and Tet2 are strongly expressed in mESCs and decrease in expression during cell differentiation. Loss of Tet1 reduces the expression of important pluripotency genes including Klf4, Nanog, and Esrrb, among others [93,94]. In addition, the expression of Tet1 and Tet2 is regulated by Oct4 [95]. During reprogramming, Tet1 has a role in removing methylation at the promoter and enhancer of Oct4 and interacts directly with Nanog [96]. Loss of Tet1 reduces the efficiency of iPS cell induction, whereas early activation of Tet1 increases reprogramming efficiency [96]. Although TET proteins appear to have important roles in modulating DNA methylation in ESCs, whether they are required to maintain pluripotency is not clear: Tet1 knockout mice survive and are fertile, and mESCs from these mice display no defects [97]. Less work has been done on TET proteins in hESCs, but knockdown of TET2 did not affect pluripotency markers although it skewed the differentiation of hESCs toward neuroectoderm at the expense of mesoderm and endoderm [98]. The chromatin in ESCs is characterized as being more open and containing more euchromatin relative to heterochromatin, compared with more differentiated cell types. Therefore, it is unsurprising that histone modifications enriched in euchromatin and associated with more active transcriptional states have also been shown to have an important role in maintaining the pluripotent state by promoting euchromatin. Histone acetylation neutralizes the charge affinity between DNA and histone proteins, leading to more open chromatin that is found at active chromatin domains [99]. That a variety of HDAC inhibitors consistently increase reprogramming efficiency [74] also indicates that open, active chromatin is an important feature of pluripotency. Trithorax group complexes are responsible for the H3K4me3 mark that is strongly enriched at active promoters, and core trithorax complex member Wdr5 is most highly expressed in pluripotency [70]. Wdr5 expression decreases as differentiation progresses, and it can aid in reprogramming [70]. It is also a downstream target of Oct4, Sox2, and Nanog [70]. WD repeat domain 5 protein interacts directly with OCT4 and chromatin remodeler INO80 to increase activation of self-renewal genes [100]. In addition to the interaction of cell signaling molecules with transcription factors, there are also extensive interactions between cell signaling and chromatin itself. For example, JAK signaling results in phosphorylation of tyrosine 41 on histone H3, leading to reduced affinity of heterochromatin protein 1a binding at pluripotency genes and promoting gene expression [101]. The chromatin remodeling factor Brg1 acts in Stat3 target genes in pluripotent cells [102]. The embryonic stem cell Brg1/hBrm associated factor complex, an SWI/SNF chromatin remodeling complex that contains Brg1 as a subunit and has a crucial role in mESCs, has also been shown to be an integration point for multiple signaling pathways and transcriptional regulators. Not only does it associate with Oct4, Sox2, and Nanog, it binds with Stat3 and Smad1, linking chromatin remodeling to both the LIF and BMP signaling pathways [102e104].

MICRORNAS INTEGRATE WITH CELL SIGNALING AND TRANSCRIPTION FACTORS TO REGULATE STEM CELL PROLIFERATION AND DIFFERENTIATION Noncoding RNAs represent an additional level of regulation of gene expression and have an important role in stem cell maintenance as well as cellular differentiation. MicroRNAs, a class of noncoding RNAs, are short single-stranded RNA molecules that are usually 20e30 nucleotides in length and function by targeting a specific set of mRNA transcripts, interfering with their transcription. The most common mechanism for miRNA processing involves a series of cleavage steps mediated by enzymes Drosha and Dicer [105]; the resulting miRNA is incorporated into a protein complex known as the RNA-induced silencing complex [106]. Relatively few miRNAs are expressed in ESCs compared with differentiated cells [107e109]. Despite this, studies have identified several clusters of miRNAs that are highly enriched in ESCs. For instance, the miR302 cluster is selectively expressed in both mESCs and hESCs compared with differentiated cells, whereas mESCs also express the

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miR290 cluster and hESCs express the miR17 family, and miRNAs from the chromosome 19 miRNA cluster and miR-371-373 cluster (thought to be orthologous to mir290 in mESC) [110e112]. Core pluripotency transcription factors Oct4, Sox2, and Nanog can bind the promoters of most of these ESC-enriched miRNAs [113]. The importance of these miRNAs to ESC maintenance was demonstrated by deleting miRNA processing enzymes Drosha and Dicer. Dicer knockout resulted in embryonic lethality in mice [114], which indicated an important role for miRNA in development. In addition, in mESCs, knockout of Dicer or the Dicer binding partner DiGeorge syndrome critical region gene 8 caused proliferation defects and prevented proper differentiation; however, stem cell self-renewal appeared to be preserved [115e117]. Similarly, in hESCs, knockdown of DICER or DROSHA caused reductions in cell proliferation [118]. These results indicate that miRNAs function in ESCs to regulate the cell cycle: in particular, the shortened G1-S transition that uniquely characterizes stem cells. More broadly, it is hypothesized that pluripotency is facilitated by short cell cycle gap phases and that gap phases represent specific windows of opportunity for stem cells to initiate the differentiation process. Another interesting aspect of miRNA is their ability to contribute to reprogramming during iPS cell formation. Expression of miR290 cluster members or several other miRNAs can increase the efficiency of reprogramming [119e121], whereas knockdown of Dicer or Drosha reduces efficiency [119]. Reportedly, reprogramming was completely inhibited by the deletion of the miR302 cluster in human fibroblasts [122]. ESC-expressed miRNAs have also been shown to influence cell signaling; miR290 members can repress Wnt signaling inhibitor Dkk1 in mESC [123], and mir302 targets Lefty, an inhibitor of the TGFb/Nodal signaling pathway in hESCs [124].

CHROMATIN STRUCTURE DETERMINES REGULATORY ACTIVITY OF TRANSCRIPTION FACTOR BINDING TO PLURIPOTENCY GENES Two-dimensional views of chromatin (such as strictly linear perspectives on transcription factor control of nearby target genes) do not reflect its unique three-dimensional (3D) structure accurately. The major roles of 3D chromatin structure in pluripotency have come into focus. Regulation of chromatin structure and the formation of chromatin loops are essential aspects of pluripotency; the proper expression of members of the cohesin and mediator complexes is required for stem cell maintenance, and the preexisting chromatin structure in somatic cells has been shown to have a strong influence on the reprogramming efficiency of these cells into iPS cells [125]. In fact, the expression of transcriptional factors that regulate reprogramming and their binding to gene promoters of pluripotency targets such as Oct4 is not enough on its own to activate gene transcription; the proper chromatin structure must also be in place. In the case of OCT4 in hESCs, it has been shown that in addition to the binding of transcription factors such as OCT4 and NANOG, activation requires a downstream enhancer to be in close proximity to the promoter, and that this structure depends on the mediator and cohesin protein complexes [126]. The intersection of transcription factor binding and chromatin structure is also evident in long-range interactions that are present at the Nanog locus. In mESCs and iPS cells, it has been shown that mediator and cohesin are responsible for approximately 40% of the observed long-range genomic interactions with the Nanog locus, and that Med1 and Smc1a interact directly with pluripotency and reprograming factors Oct4, Sox2, Klf4 [127]. A separate study showed that tethering Nanog to a region of the genome in mESCs was sufficient to induce a number of chromatin loops between pluripotency genes, which artificially created a Nanog binding site [128]. In addition, Klf4 is required for cohesin complex recruitment to Oct4 and for the resulting loop between enhancer and promoter [129]. Together, these results suggest that pluripotency and reprograming factors are involved in mediating chromatin looping and in the recruitment of cohesin and mediator complexes to gene targets where chromatin looping has essential roles in pluripotency. From some of these studies, it is also evident that chromatin looping at pluripotency genes tends to occur before gene activation. For example, during the reprogramming process, it is striking that pluripotency-specific genomic interactions of the Nanog locus occur several days before activation of Nanog expression [127]. These data support a model in which reprogramming factors bind and act to recruit cohesin and mediator complexes to restructure the chromatin, forming loops that place enhancers and promoters in proximity, and group coexpressed genes together, before these factors then can activate gene expression.

CONCLUSIONS The molecular basis of pluripotency is a complex coordination of extracellular and environmental factors with intracellular signal transduction and transcriptional regulation via specific epigenomic events. We have seen significant leaps in understanding of how signaling cascades converge upon core transcriptional circuitry to coordinate

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maintenance and induction of pluripotency. Furthermore, understanding of the intricate mechanisms through which signaling pathways create the multitude of cell types and tissue lineages remains paramount to understanding basic human development as well as to manipulating and controlling lineage specification for the purposes of regenerative medicine. The rapid advent of induced pluripotency by reprogramming differentiated cells into an embryonic state through cocktails of transcription or chemical factors has opened significant doors to the concept of personalized regenerative medicine. Understanding the fundamental mechanisms of this process may ultimately provide us with unprecedented control to reprogram somatic cells into iPS cells or directly to any desired cell type for the purpose of cell transplantation. Toward that goal, new technologies such as clustered regularly interspaced short palindromic repeats (CRISPR)/CRISPReassociated protein 9 gene editing and pluripotent stem celle based organoids are likely to accelerate new discoveries further.

List of Acronyms and Abbreviations BMP Bone morphogenic protein EpiSC Epiblast stem cell FGF Fibroblast growth factor GSK3b Glycogen synthase kinase 3b HDAC Histone deacetylase hESC Human embryonic stem cell iPS cell Induced pluripotent stem cell LIF Leukemia inhibitory factor MEK mitogen activated protein kinase/extracellular signaleregulated kinase mESC Mouse embryonic stem cell OCT4 Octamer binding protein 4 PRC2 Polycomb repressive complex 2 SOX2 SRY-box 2 TCF3 Transcription factor 3 TGFb Transforming growth factor b

Acknowledgment The authors thank Ali Brivanlou and Harvir Singh, who were the authors of the previous edition of this chapter, as some of their work in that previous version of the chapter has been carried over to the current version.

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C H A P T E R

5 Scarless Wound Healing: From Experimental Target to Clinical Reality Alessandra L. Moore1,2, Clement D. Marshall1, Allison Nauta1,3, Hermann P. Lorenz1, Michael T. Longaker1 1

Stanford University School of Medicine, Palo Alto, CA, United States; 2Brigham and Women’s Hospital, Boston, MA, United States; 3Oregon Health and Sciences University, Portland, OR, United States

INTRODUCTION Though scarring is the normal outcome of wound healing in adults, little is talked about its long-term consequences. For instance, because skin and other organs scar rather than regenerate after injury, several physiologic sequelae often occur. After bowel surgery, the visceral and parietal peritoneum develop dense scar tissue known as adhesions, which cause postoperative bowel obstructions [1]. After traumatic injury to soft tissue, ligaments, and tendons, scarring can cause contractures that limit movement and cause functional restriction [2]. Scarring in the nervous system results in loss of function because neuronal connections are destroyed [3]. In the cornea, scarring limits visual acuity [4]. Cardiac myocytes are also severely affected by scarring, as seen by the number of arrhythmias associated with impaired conduction after infarct [5]. The only exceptions to the rule of fibrosis after injury in adults are found in bone, oral mucosa, and hepatic tissue, which are capable of partial regeneration [6,7]. The most significant phenotype in skin scar formation can be found in burn injuries [8,9]. Approximately 500,000 patients in the United States undergo medical treatment for burn injuries annually; over one-third of patients have burns that exceed 10% of the total body surface area (American Burn Association Burn Incidence Fact Sheet, 2016). In addition, many of these patients are children, a population that is particularly vulnerable to the negative physical and psychological effects of scarring; 10,000 children are permanently disabled each year from burn injury and fibrosis [8e12]. Perhaps the most dramatic consequence of burns are hypertrophic scars, which afflict up to 70% of burn patients in their lifetime. These dense, fibrous scars can limit range of motion severely and can negatively affect quality of life through pain, itching, and irritation rendered from the scar tissue itself [13]. All of this, of course, does not account for the long-term psychological impact caused by delayed reintroduction to society after hospitalization, partly from disability but also from issues related to confidence and cosmetic appearance [13]. As mentioned, not only normal scars cause problems for adult and pediatric patients. Issues with underhealing and overhealing also cause significant morbidity. Nonhealing chronic ulcers and excessive fibroproliferative scarring are the most common examples of dysfunctional wound healing, each of which is prevalent in certain pediatric and adult populations. Often, dysfunctional wound healing will cause pain, functional restriction, and severe psychological outcomes [14,15]. In patients with chronic illnesses, dysfunctional wound healing usually falls into the category of “underhealing”. In this situation, wounds fail to heal for numerous reasons, including infection, impaired blood flow, severe malnutrition, and inadequate wound care. These patients have become an increasing concern, particularly as the population ages and more health care resources are allocated to treat chronic diseases and their associated complications. Diabetic patients are a dramatic example of the burden of chronic wounds on society. The following statistics, obtained from the Centers for Disease Control and Prevention’s 2014 National Diabetes Statistic Report, illustrate the magnitude of diabetic nonhealing wounds: 29 million people in the United States have diabetes, which makes diabetic patients approximately 10% of the US population. In addition, 65% of diabetic patients have at least one other

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co-occurring illness. In 2010, 73,00 nontraumatic lower limb amputations were performed in patients with diabetes. A total of 23% of all patients with diabetes have foot problems ranging from numbness to amputations. From 25% to 50% of all hospital admissions for diabetes are for nonhealing ulcers, which are the cause of most nontraumatic extremity amputations performed in the United States each year [16]. On the other extreme, excessive healing is also a serious clinical burden. Pathological scarring causes hypertrophic scars and keloids. Whereas hypertrophic scars eventually abate with management and time, keloids continue to grow as a type of soft tissue tumor that may persist for life after excision [17]. These scarring processes cause functional impairment and symptoms such as burning, itching, and pain. Ultimately, hypertrophic scars and keloids are difficult to treat medically or surgically, particularly because uniform treatment algorithms do not exist [18]. As is noted from these examples of normal and dysfunctional wound healing, the burden of wound healing is substantial in the United States and on patients. As such, research in the field is urgently needed to help patients regain functional abilities and limit morbidity after injury. Here, we discuss the normal mechanism of adult wound healing and explore the remarkable ability of fetal skin to regenerate. Current and future therapeutics will be discussed, with the hope of highlighting the potential of regenerative healing.

ADULT SKIN Anatomy of Adult Skin Adult skin is made up of two layers, the epidermis and dermis. The epidermis is a type of keratinizing stratified squamous epithelium that has five distinct layers, each characterized by the level of keratinocyte maturation. Keratinocytes originate from a thin sheet of stem cells in the basal epidermis and migrate to the surface over 4 weeks to become soft keratin, which eventually sloughs off. Epidermal appendages, which are epithelial structures that extend intradermally, are an important source of cells for reepithelialization in skin wound healing. Epidermal appendages include sebaceous glands, sweat glands, apocrine glands, and hair follicles. Appendages can extend deep into the dermis or through the dermis into the subcutaneous tissue/hypodermis. The hair follicle is an important appendage composed of an external outer root sheath attached to the basal lamina that is contiguous with the epidermis. The hair follicle also contains a channel with a hair shaft. Together, the hair follicle and its attached sebaceous gland are called the pilosebaceous unit. The base, or bulb, of the hair follicle contains a committed but proliferating progenitor cell population, as well as the matrix encasing the dermal papilla. The dermal papilla contains specialized mesenchymal cells. From this region, the hair shaft and its channel will grow. Sweat glands (or eccrine glands) produce sweat, a mixture of water and salts, which functions to cool the body by evaporation. The sweat gland is distinct in that it contains a coiled intradermal portion that extends into the epidermis by a relatively straight distal duct. These glands, too, have an important role in normal skin, as well as physiologic homeostasis. Below the epidermis lie two distinct layers of the dermis: the more superficial papillary layer and the deeper, more fibrous reticular layer. The papillary dermis is highly vascular, sending capillaries (dermal papillae) superficially. The reticular dermis contains densely packed collagen fibers and tends to be less vascular, except where sweat glands and hair follicles traverse through the tissue. Reticular dermis is also rich in elastic fibers and contains some macrophages, fibroblasts, and adipose cells that participate in skin wound healing [19] (Fig. 5.1).

Adult Wound Healing and Scar Formation Adult wound healing is traditionally described as a sequence of temporally overlapping phases: inflammation, proliferation, and remodeling. Disruption of the vascular network within cutaneous wounds results in platelet aggregation and the formation of a fibrin-rich clot, which protects from further extravasation of blood or plasma. Aggregation of platelets initiates the coagulation cascade [20e22]. In addition to providing hemostasis, platelets modulate fibroblast activity through degranulation and secretion of multiple cytokines and growth factors, such as platelet-derived growth factor (PDGF), platelet factor 4, and transforming growth factor b1 (TGF-b1). These growth factors and cytokines remain elevated throughout the process of normal wound healing [23,24]. Largely under the influence of platelet-derived inflammatory molecules, neutrophils and monocytes initiate their migration to the wound. Because of the high concentration of neutrophils in circulation, these cells are the first to enter the area of injury and quickly reach high concentrations, becoming the most dominant influence in early wound healing. Neutrophils primarily produce degradative enzymes and phagocytose foreign and necrotic material, but they also produce vascular endothelial growth factor (VEGF), tumor necrosis factor a, interleukin 1 (IL-1),

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ADULT SKIN

Hair shaft Stratum Corneum

Pore

Pigmented epithelial cells

Stratum Gingivatum

Epidermis

Stratum Spinosum Stratum Basale*

Dermis Erector pili muscle

Hypodermis

Sweat gland Sebaceous gland*

Blood and lymph vessels Hair follicle

Dermal papilla*

Fat cells Hair follicle bulge*

Adipose tissue stroma*

Normal skin anatomy. Normal skin anatomy is composed of two layers, the epidermis and dermis, with a layer of subcutaneous fat known as the hypodermis below. Contained within the dense array of epithelial cells and extracellular matrix are many specialized structures including sebaceous glands, sweat glands, and hair follicles. In wound healing, normal tissue homeostasis, and tissue engineering, skin stem cells have an important role. Areas where resident adult stem cells are located are indicated by an asterisk. Not pictured are interfollicular epidermal stem cells.

FIGURE 5.1

and other growth factors that assist in wound healing [25]. Interestingly, studies show that neutrophil infiltration is not essential to normal healing, which demonstrates one of many redundancies in the repair process [26]. The level of inflammation in a wound ultimately depends on the presence or absence of infection. In the presence of infection, neutrophils continue to be active in high concentration throughout later stages of wound healing, leading to further inflammation and fibrosis [22]. In the absence of infection, neutrophils greatly diminish activity by day 2 or 3, as monocytes increase in number in response to both extravascular and intravascular chemoattractants. Monocytes differentiate into macrophages that bind to the extracellular matrix (ECM), which induces phagocytosis and allows for debridement of necrotic cells and fractured structural proteins. During the late inflammatory phase, tissue macrophages release cytokines and scavenge dead neutrophils, then elevate macrophages as the dominant leukocyte in the wound bed. In contrast to neutrophils, studies on tissue macrophage and monocyte-depleted guinea pigs demonstrated macrophages to be essential to normal wound healing through their stimulation of collagen production, angiogenesis, and reepithelialization [27]. However, like neutrophils, if macrophages persist in the wound environment, the result is excess scar formation. Under these circumstances, macrophages produce cytokines that activate fibroblasts to deposit excessive amounts of collagen [28] (Fig. 5.2). The presence of macrophages in the wound marks the transition between the inflammatory phase and the proliferative phase of wound healing, which begins around day 4e5 after injury. Granulation tissue begins to form and is described as a loose network of collagen, fibronectin, and hyaluronic acid embedding a dense population of macrophages, fibroblasts, and neovasculature. During the deposition of granulation tissue, macrophages and fibroblasts move into the wound space as a unit while new blood vessels sprout from adjacent exposed endothelium [22]. The rate of granulation deposition depends on many factors, including the interaction between fibronectin and fibroblast integrin receptors [29]. Fibroblasts in the wound bed deposit collagen and a proteoglycan-rich provisional matrix, a process that is stimulated by TGF-b1 and TGF-b2 in adult wounds. Studies have shown that exogenous administration of these molecules leads to increased collagen and inflammatory cells at the wound site, which potentially implicates this subfamily of growth factors as a source of overactive scar formation [30,31]. During the proliferative phase of wound healing, which occurs from approximately day 5 to day 14 after wounding, collagen is deposited more densely in the wound. Once a threshold level of collagen is reached, collagen synthesis and fibroblast accumulation are suppressed by an unknown negative-feedback mechanism [32]. The balance

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FIGURE 5.2 Inflammatory cell recruitment to the site of tissue damage. Therapeutic intervention aimed at dampening the immune response could target any of the steps along the pathway of inflammatory cell recruitment. (A) Leukocytes in blood vessels adjacent to the site of tissue damage emigrate through the vessel wall by diapedesis and (B) migrate to the site of tissue damage in response to chemotactic signals. Inflammatory cells activate resident fibroblasts and attract other bone marrowederived cells to the wound, where the repair outcome is (C) scar formation. After acting at the wound site, the activated repair cells disperse, differentiate, or (D) apoptose, thus ending the repair response.

of collagen synthesis and degradation is controlled by collagenases and tissue inhibitors of metalloproteinases. When this negative feedback does not occur appropriately, pathological scars form with the deposition of densely packed, disorganized collagen bundles [22]. Reepithelialization begins in the first 24 h after wounding, with the goal of creating a new protective, natural skin barrier. Initially, basal keratinocytes at the border of the wound, which under normal circumstances are linked together by desmosomes, detach from each other and the ECM and migrate laterally to fill the void in the epidermis. Through this process, keratinocytes are exposed to serum for the first time. As they are subjected to new and increased levels of inflammatory cytokines and growth factors, keratinocytes are signaled to further migration, proliferation, and differentiation [33]. Neovascularization also occurs during the proliferative phase and is influenced by multiple cytokines, circulating endothelial progenitor cells, and the ECM [34]. The formation of new blood vessels can also be stimulated by lactic acid, plasminogen activator, collagenases, and low oxygen tension [22]. New blood vessels in the wound bed bud and grow at an exceptional rate during skin wound healing, eventually giving the wound bed adequate oxygenation and its characteristic pink hue. Once apoptotic pathways become active again as the granulation tissue matures, angiogenesis eventually desists and the numerous new blood vessels reduce to only a few major tributaries [35]. The maturation phase of wound healing consists of collagen remodeling beginning during the second week of healing. At this point, fibroblasts differentiate into myofibroblasts, which are characterized by greater expression of smooth muscle actin and the ability to contract wounds. Fibroblasts eventually decrease in number from the proliferative phase, and the scar tissue becomes less vascular and paler as vessels involute [36]. Scar tissue finally gains tensile strength as collagen cross-links increase during remodeling, largely owing to the action of myofibroblasts. Maturation involves the replacement of the initially randomly oriented types I and III collagen by predominantly type I collagen organized along the lines of tension. Eventually the process completes; however, scar tensile strength will never reach that of unwounded skin [37] (Fig. 5.3).

Fibroproliferative Scarring Fibrosis is defined as “the replacement of the normal structural elements of the tissue by distorted, nonfunctional, and excessive accumulation of scar tissue” [38]. Many medical problems are linked to excessive fibrosis, in addition to a number of deaths each year in the United States and across the globe [37,39]. As previously

ADULT SKIN

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(A)

(B)

(C)

FIGURE 5.3 Classic stages of wound repair. There are three classic stages of wound repair: inflammation (A), new tissue formation (B), and

remodeling (C). (A) Inflammation. This stage lasts until about 48 h after injury. Depicted is a skin wound at about 24e48 h after injury. The wound is characterized by a hypoxic (ischemic) environment in which a fibrin clot has formed. Bacteria, neutrophils, and platelets are abundant in the wound. Normal skin appendages (such as hair follicles and sweat duct glands) are still present in the skin outside the wound. (B) New tissue formation. This stage occurs about 2e10 days after injury. Depicted is a skin wound at about 5e10 days after injury. An eschar (scab) has formed on the surface of the wound. Most cells from the previous stage of repair have migrated from the wound and new blood vessels now populate the area. The migration of epithelial cells can be observed under the eschar. (C) Remodeling. This stage lasts for a year or longer. Depicted is a skin wound about 1e12 months after repair. Disorganized collagen has been laid down by fibroblasts that have migrated into the wound. The wound has contracted near its surface and the widest portion is now the deepest. The reepithelialized wound is slightly higher than the surrounding surface and the healed region does not contain normal skin appendages. Reproduced with permission from Gurtner GC, Werner S, Barrandon Y, Longaker MT. Wound repair and regeneration. Nature 2008; 453(7193):314e21.

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mentioned, excessive fibroproliferative scarring occurs when the mechanisms of wound healing fail to respond to inhibitory cues at any one of the phases of wound healing. In essence, abnormal scar formation is an excess accumulation of unorganized collagenous ECM. Although the appearance of scars is often unpredictable, several factors influence the severity of scarring. These include not only genetics but also tissue site, sex, race, age, magnitude of injury, and wound contamination. Generally, skin sites with thicker dermal tissue tend to heal with thicker scars. Also, estrogen is believed to promote scarring; premenopausal women have denser scar tissue than do postmenopausal women. Patients with darkly pigmented skin are also more prone to thicker scarring, as are young people. Larger, deeper, and more contaminated wounds tend to produce larger resultant scars, as well [40e44]. Keloids Over months to years, some scars will develop benign but locally aggressive tumors known as “keloids” [45]. This extreme example of fibroproliferative scarring is characterized by “cauliflower” nodules that extend well beyond the area of original injury and continue to grow, sometimes into significant masses [45]. Symptoms can range from irritation and pain to severe disfigurement and functional restriction owing to the appearance, size, color, and location of lesions [17,18,28,39,46e48]. In addition to being disfiguring, keloids tend not to regress. Instead, they can continue to grow slowly over many years, and their growth rate tends to correlate with symptoms. The most common areas affected by keloids are upper-body sebaceous areas, whereas the extremities are less commonly involved [49]. Aside from being involved in areas with a high density of sebaceous glands, tension may have a role in their development over time [45]. Histologically, keloids are characterized by thick, large, closely packed bundles of disorganized collagen. In addition, the scar tissue is colonized by mast cells, eosinophils, plasma cells, and lymphocytes with a stark absence of macrophages [45]. Unlike hypertrophic scars, keloids have nodules in the mid to deep dermis containing densely packed collagen with a few myofibroblasts. Mucin is deposited focally in the dermis, and hyaluronic acid expression is confined to the thickened, granular/spinous layer of the epidermis. Finally, an amorphous and unidentified mucopolysaccharide surrounds the dense collagen bundles in keloidal scars that is not found in other scar tissue types [18]. Interestingly, in keloids, more than one mechanism of dysfunctional healing may be at play. Fibroblasts in keloids tend to respond abnormally to wound healing growth factors, secreting collagen, elastin, and fibronectin with little response to inhibitory compounds. Moreover, the ability for fibroblasts to proliferate in keloids is increased. In addition, dysfunctional apoptosis has been observed as keloidal wounds enter the remodeling phase of wound healing [50]. It is possible that this is partly the result of central scar ischemia and ongoing release of VEGF, hypoxia inducible factor 1a, and other growth factors that promote proliferation [17,18,28,39,46e48]. Finally, although a number of therapies exist for the treatment of keloids, including surgical excision, laser therapy, cryotherapy, and chemotherapy, keloids remain remarkably resistant and recur over time [51e54]. Because of these features, it is possible that keloid tumor cells are mutated skin progenitor cells which are known to display a similar resilience [55] (Fig. 5.4A). Hypertrophic Scars The incidence of hypertrophic scars, another type of fibroproliferative scar, is significantly higher than keloids. This is partly because of the large number of young adults with initially hypertrophic scars that eventually regress with age (35%) [56]. Hypertrophic scars are raised, similar to keloids; however, they are usually no more than 4 mm from the surface of skin. Visually, hypertrophic scars are erythematous or brown-red in color, but they can become pale over time. Also, hypertrophic scars will sometimes regress spontaneously. Yet, also unlike keloids, which are predominantly over skin containing sebaceous glands, hypertrophic scars usually occur on extremity joints such as the elbows and knees [48]. Histologically, hypertrophic scars are characterized by the expansion of collagen bundles that are fine, wellorganized, and parallel to the epidermis. Hypertrophic scars do not have the nodal collagen bundles seen in keloids, but they have islands of hypercellularity in the deep dermis that upon closer inspection contain aggregates of fibroblasts, collagen, and neovasculature. Mucin is absent and hyaluronic acid is a major component of the papillary dermis [18,49]. The purported mechanism of hypertrophic scar formation is increased and ongoing collagen deposition with decreased collagenase activity. In addition, fibroblasts in hypertrophic scars are more likely to assume a myofibroblast phenotype, which may contribute to collagen deposition and scar contracture without addressing scar remodeling. Because 70% of burn victims are affected, initial wound depth may have a significant impact on the development of hypertrophic scars, because they appear to be most common in patients with delayed healing or

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ADULT SKIN

(A) KELOID SCAR Single or Small

Large or Multiple

Radical therapies

Surgery and Adjuvant Therapy

Non-Surgical Monotherapy

Effective Effective

- Cryotherapy - Laser therapy - Compression therapy - Silicone gel sheets - Bleomycin, 5-FU - Cosmetic camouflage

- Surgery and corticosteroids - Surgery and radiation

- Corticosteroids - Cryotherapy - Laser therapy - Bleomycin, 5-FU

Effective

Conservative therapy and longterm follow-up

Ineffective

Consider Non-surgical Monotherapy

Repeat symptomatic non-surgical polytherapy or perform mass reduction surgery

Conservative therapies and long-term follow-up

- Compression therapy - Silicone gel sheets - Cosmetic camouflage

Effective

Ineffective

Recurrence

Recurrence

Repeat, Conservative therapies, and long-term follow-up

Mass Reduction Surgery

Symptomatic Nonsurgical Polytherapy

- Compression therapy - Silicone gel sheets - Cosmetic camouflage

(B) HYPERTROPHIC SCAR

Without Contracture If ineffective, repeat algorithm

Linear or small

Wide or large Partial Surgical Contracture Release

Nonsurgical Polytherapy - Silicone gel sheets - Corticosteroids - Compression dressings - Laser therapy - Cosmetic camouflage

Severe Contracture

Mild Contracture

- Skin grafts - Local flap coverage

Complete Surgical Resection

Recurrence

Effective

Conservative therapy with long-term follow-up

Effective

Surgery with adjuvant therapy

Repeat surgical correction

Ineffective

Conservative therapy with long-term follow-up

- Surgery and corticosteroids - Surgery and radiation

Effective

Recurrence

Nonsurgical Polytherapy

FIGURE 5.4 Scar reduction strategies: algorithms for (A) keloids and (B) hypertrophic scars. Modified from Ogawa R. The most current algorithms for the treatment and prevention of hypertrophic scars and keloids. Plast Reconstr Surg 2010; 125(2):557e68.

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a prolonged inflammatory phase. In deeper and larger wounds, fibroblasts assume a myofibroblast phenotype, which renders them less capable of reducing type I collagen in scars during the remodeling phase [56] (Fig. 5.4B).

Underhealing: Chronic Skin Ulcers Many types of chronic nonhealing dermal ulcers exist, such as pressure ulcers, diabetic lower-extremity ulcers, and venous stasis ulcers. Besides causing pain and disfigurement, chronic ulcerations often lead to infection as well as amputation. In patients with chronic ulcerations such as those who have diabetes, often the single most important life-saving measure is limb salvage to preserve mobility and function [57]. In those who are immobile, pressure ulcers predominate, as seen by their commonality among debilitated or institutionalized patients, those with spinal cord injuries, and patients with cerebrovascular infarcts. The total cost of wound care per hospitalization for pressure ulceration is estimated to be $130,000, a figure expected to grow as the population ages [58]. Interestingly, nonhealing wounds have distinct histologic hallmarks too, as do keloid and hypertrophic scars have their own set of unique histologic hallmarks. In chronic ulcers, there is excessive neutrophil infiltration, and wounds seem to remain in the inflammatory phase. Likely, the abundance of neutrophils is responsible for the lingering inflammation seen in chronic ulcers. As neutrophils release enzymes, such as collagenase (matrix metalloproteinase [MMP] 8), connective tissue is digested as fast as new matrix is deposited [59]. Neutrophils also release elastase, an enzyme known to destroy PDGF and TGF-b, which are important growth factors for normal wound healing [60]. The environment of chronic ulcers is also known to contain an abundance of reactive oxygen species that damage healing tissue [61]. The result of this nonhealing mechanism is a wound that not only fails to heal but perpetuates its existence.

FETAL SKIN Development of Fetal Skin The skin’s superficial layer, the epidermis, is derived from surface ectoderm, whereas the dermis is generated from mesenchyme. The epidermis starts as a single layer of ectodermal cells covering the embryo, which begins to emerge at gestational day 20 in humans. In the second month, cell division takes place, at which time the periderm (epitrichium) emerges as a thin superficial layer of squamous epithelium overlying the basal germinative layer. Over the next 4e8 weeks, the epidermis becomes highly cellular. New cells are produced in the basal germinative layer and are continuously keratinized and shed to replace cells of the periderm. Together, these cells are part of the vernix caseosa, a greasy white film that covers fetal skin. In addition to desquamated cells, the vernix caseosa contains sebum from sebaceous glands. With sebum and desquamated cells mixed together, this slippery substance serves as a protective barrier during gestation and facilitates passage through the birth canal at delivery. Replacement of the periderm continues until the 21st week, at which point the periderm has been replaced by the stratum corneum. Through a series of stages of differentiation, the epidermis stratifies into four layers by the end of the fourth month, including the stratus germinativum (derived from the basal layer), the thick spinous layer, the granular layer, and the most superficial stratum corneum. By the time these four layers emerge, interfollicular keratinization has begun and the epidermis has developed buds that become epidermal appendages. Melanocytes of neural crest origin have also invaded the epidermis, synthesizing melanin pigment that can be transmitted to other cells through dendritic processes. By the end of the 21st week, the fetal epidermis has many of the components that it will maintain into adulthood. After the 21st week, the dermis begins to mature from a thin and cellular structure to a thick and fibrous one. By 24 weeks of gestation, fetal skin matures and thickens to become histologically distinct from its embryonic beginnings [62].

Fetal Scarless Wound Repair Whereas adult wounds heal with fibrous tissue (scarring), early-gestation fetal skin wounds heal scarlessly. Fetal wounds are capable of healing with restoration of normal skin architecture and preservation of both tissue strength and function. This observation has been confirmed in multiple animal species, including mice, rats, rabbits, pigs, sheep, and monkeys. The mechanisms responsible for fetal scarless wound healing are intrinsic to fetal tissue and are independent of environmental or systemic factors such as bathing in sterile amniotic fluid, perfusion with fetal serum, or the fetal immune system [63e65]. To support this point, studies in which human fetal skin

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that was transplanted subcutaneously into the dorsolateral flank of athymic mice healed without a scar further suggested that the scarless wound phenotype depends on characteristics intrinsic to fetal tissue [66]. Ultimately, scarless fetal wound repair outcomes depend on two factors: the gestational age of the fetus and the size of the wound. Excisional wound healing studies performed on fetal lambs showed that at a given gestational age, larger wounds healed with an increased incidence of scar formation. Likewise, the frequency of scarring increased with increasing gestational age [67]. Since the publication of these studies, transitional periods have been found for humans (24 weeks’ gestation) [68], rats (between gestation days 16.5 and 18.5) [64], and mice [69]. Extensive research has been dedicated to determining the culprit for the shift to adult wound healing. Eventually, instead of depositing bundles of ECM in a normal basket-weave pattern, organisms begin to heal breaches in skin integrity with scarring. Interestingly, as fetuses develop and enter the early period of scar formation, the wound phenotype has been described as a “transition wound.” At this point, the repair outcome is tissue that retains the reticular organization of collagen characteristic of normal skin but is devoid of epidermal appendages [70]. The skin does not truly regenerate, but the dermis does not form a scar. In addition, although the fetal ECM was once thought to be inert, evidence suggests that it is a dynamic structure that has a pivotal role in cellular signaling and proliferation. The fetal ECM is now known to be a reservoir of growth factors essential to development [71]. The fetal ECM also has a structural protein composition different from that of adult skin. For example, the collagen content of the ECM changes as the fetus ages, starting with a relatively high type III to type I collagen ratio and shifting to the type I collagen-predominated adult phenotype. Another structural difference between fetal and adult ECM can be found in the hyaluronic acid content. Hyaluronic acid, the negatively charged, extremely hydrophilic, nonsulfated glycosaminoglycan of the ECM, has been shown to be in higher concentration in the ECM during rapid growth processes such as cellular migration and angiogenesis. In vitro studies show that hyaluronic acid can cause fibroblasts to increase the synthesis of collagen and noncollagen ECM proteins [72]. During adult repair, hyaluronic acid initially increases dramatically, and then decreases from days 5 to 10, after which the concentration remains at a low level. Interestingly, this hyaluronic acid profile is not the case in fetal wound ECM, in which the hyaluronic acid level remains high. Similar to the concentration of type III collagen, the ECM hyaluronic acid content decreases from the fetal to the postnatal period [63]. The concentration of other substances such as decorin, fibromodulin, lysyl oxidase, and MMPs further sets the fetal ECM apart from the adult ECM [71]. These substances are proteoglycan ECM modulators that have a role in the development and maturation of collagen. Lysyl oxidase cross-links collagens whereas MMPs degrade collagen. In addition, decorin content and the expression of lysyl oxidase and MMPs increase as fetal tissue matures. Fibromodulin modulates collagen fibrillogenesis and has been shown to bind and inactivate TGF-bs. Fibromodulin decreases with gestational age, paralleling the shift from scarless fetal wound healing to scarring adult repair [73] (Fig. 5.5).

REGENERATIVE HEALING AND SCAR REDUCTION THEORY Targeting the Inflammatory Response Initial research into the mechanisms responsible for scar formation led investigators to focus on the inflammatory phase of wound healing as a target for reducing the incidence and magnitude of scar formation. This choice was based on observations that regenerative wound healing is replaced by scarring as the immune system develops in the embryo [65]. Interestingly, many studies have shown that reduction of inflammation in postnatal skin wounds correlates with reduced scarring [74]. One example of reduced inflammation and scarring can be found in enhanced healing in mice devoid of mothers against decapentaplegic homolog 3 (Smad3), a protein known to transduce TGF-b signals. These mice exhibited more rapid reepithelialization and decreased inflammation (blunted monocyte activation) [43]. In another example, PU.1 null mice devoid of functional neutrophils and macrophages heal wounds over a time course similar to that of their wild-type counterparts but exhibit scar-free healing similar to embryonic wound healing [75]. These two studies support the contention that the inflammatory response may be deleterious to normal wound repair by contributing to increased fibrosis. Experiments performed on athymic mice [74] and experiments involving antisense downregulation of connexin43, a protein involved in gap junctions and inflammation, support these findings [74,76]. Furthermore, other studies have provided evidence that wound inflammatory cells from the circulation

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5. SCARLESS WOUND HEALING

FIGURE 5.5 Wound histologic sections at 48 h after injury in embryonic day 17 fetuses. India ink was injected at the time of injury. (A) Lowpower magnification reveals complete reepithelialization and a mild increase in the number of inflammatory cells present (arrows) (hematoxylin and eosin; original magnification, 100; bar ¼ 100 mm). (B) High-power view shows India ink (arrowheads) collected around regenerating hair follicles (arrow) (eosin; original magnification, 400; bar ¼ 25 mm). (B) Mallory trichrome shows a fine reticular dermal collagen pattern. A white arrow shows a hair follicle (Mallory trichrome; original magnification, 400; bar ¼ 25 mm). Reproduced with permission from Colwell AS, Krummel TM, Longaker MT, Lorenz HP. An in vivo mouse excisional wound model of scarless healing. Plast Reconstr Surg 2006; 117(7):2292e96.

produce signals that either directly or indirectly induce collagen deposition and granulation tissue formation, which increase scarring [77]. Although this research points to the inflammatory phase of wound healing as a cause of scar tissue formation, studies have provided evidence that the inflammatory phase and scarring might not be as directly linked as previously believed. Cox-2, an enzyme involved in prostaglandin production, is a mediator of inflammation. Two studies

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show conflicting evidence regarding the effect of Cox-2 inhibition; one study reported decreased scar formation and the other claimed no difference in wound healing or scar formation [78,79]. Likewise, a study that transiently induced neutropenia in mice accelerated wound closure but failed to show a difference between collagen content in neutrophil-depleted wounds compared with wild-type controls [80]. Although other possible mediators of scar formation exist, the inflammatory response remains a major target for ongoing research aimed at preventing or reducing the appearance of scars. As Stramer et al. illustrated, many points exist at which interventions could dampen the inflammatory response. The first target could be leukocytes, at any point as they migrate (1) through the vessel wall from the bloodstream, (2) from outside the vessel to the wound, or (3) as they transmit a signal to fibroblasts, inducing the fibrotic response. A second target could be the fibroblasts themselves, and interventions could be designed to block the action of these cells as they respond to leukocyte signaling [81]. An interesting development in the scar theory was the discovery of heterogeneous fibroblast populations, with certain fibroblasts being responsible for the entirety of scar collagen production. Engrailed-1, a gene for a homeobox protein, was used to trace a lineage of fibroblast appearing at the transition from regenerative to scarring phenotype in the dorsal skin of fetal mice, and was also found to deposit all skin scar collagen in adults. This discovery may have uncovered the elusive element behind scar formation and may offer a specific cellular target to achieve regenerative healing. Also, along with the discovery of Engrailed-1 positive fibroblasts was the discovery of specific inhibitors of these fibroblasts, offering already a potential therapy to achieve reliable scar reduction. Through targeting scar fibroblasts rather than inflammatory cells, perhaps other phases of wound healing, such as the proliferative and remodeling phases, may also be investigated having potential therapeutic benefit [82].

Cytokines and Growth Factors Transforming Growth Factor b Superfamily Though there is an abundance of knowledge pertaining to the TGF-b pathway, the development of clinical tools targeting this growth factor have largely arrested. The TGF-b superfamily includes several proteins, the most important are which are TGF-b1, TGF-b2, and TGF-b3; all of them have been shown to influence adult wound healing [83]. These cytokines are secreted by keratinocytes, fibroblasts, platelets, and macrophages, which act to influence their own and other cell populations to migrate into the wound bed [84]. The TGF-b superfamily has also been implicated in matrix remodeling and collagen synthesis [85]. Partly, this comes from evidence that TGF-b1 activates myofibroblast differentiation to influence wound contraction and the synthesis of collagen and fibronectin in granulation tissue [86]. Investigators have compared the TGF-b isoform profiles of fetal and adult skin. Results showed that injured fetal epidermis contains a greater amount of TGF-b3, derived from keratinocytes and fibroblasts, and less TGF-b1 and TGF-b2, derived from degranulating platelets, monocytes, and fibroblasts, compared with healing adult skin [87,88]. Since this cytokine profile was discovered, isoforms TGF-b1 and TGF-b2 have generally been thought to be pro-fibrotic, whereas TGF-b3 is thought to support scarless healing [84]. Discovery of the relative ratios of these isoforms prompted experiments aimed at mimicking the embryonic profile, using antibody neutralization of TGF-b1 and TGF-b2 and treatment with exogenous TGF-b3 [89]. Although antibody neutralization of TGF-b1 or TGF-b2 has no effect on wound healing, experiments showed that antisense RNA knockdown of TGF-b1 does in fact reduce scar formation [90]. Likely, the length of time that TGF-b1 is neutralized over the course of wound repair influences scarring, with longer neutralization needed for greater influence on scar reduction [84]. In addition, the coordinated actions of TGF-b1 and TGF-b3 in wound healing may be more complicated and necessary than originally hypothesized, because alterations in the ratio of TGF-b1 and TGF-b3 fail to produce significant results [91]. Interestingly, clinical trials with recombinant TGF-b3 failed to demonstrate diminished scarring. Avotermin, which received acclaim and investment in early clinical trials, failed to meet its Phase III end points [92]. Since then, Avotermin and its alternative were abandoned, perhaps highlighting the importance of integrating TGF-b with tissues, cells, or other cell signaling molecules. Now, with the expansion of wound healing biomaterials, the TGF-b superfamily is being targeted as a therapeutic to be delivered via biomimetic scaffolds [93]. Engineered as well as dermal composite scaffolds are emerging as potential therapies, with some promising in vivo studies [94]. Connective Tissue Growth Factor Connective tissue growth factor (CTGF), like many other growth factors in wound healing, is capable of influencing fibroblast differentiation and inflammatory cell migration [95]. It is also considered profibrotic by a mechanism related to TGF-b, and through its influence on fibroblasts to deposit scar ECM matrix [95], as evidenced by its

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increased expression in scleroderma patients [96]. CTGF is a TGF-b target gene that is activated by Smad proteins after TGF-b binds to and activates its receptors. Once activated, this cysteine-rich matricellular protein binds to integrin and proteoglycan receptors activating pathways such as wingless type (Wnt), bone morphogenic protein, VEGF, and TGF-b [97], which highlights the redundancy in many important wound healing signaling cascades. Adult fibroblasts have higher expression of CTGF than fetal fibroblasts, which makes this an attractive therapeutic target. Studies on fetal fibroblasts stimulated by TGF-b show increased expression of CTGF, suggesting that scarless fetal repair may be partially a result of lower CTGF expression [98]. Following this evidence, CTGF was found to have a positive vulnerary effect in diabetic wound healing, accelerating the time to wound closure in diabetic foot ulcers [95]. In a rabbit ear model, hypertrophic scarring improved using messenger RNA (mRNA) antisense inhibition of CTGF [99]. Like TGF-b, CTGF may be involved in too many overarching signaling cascades to influence wound healing by solitary treatment with recombinant or inhibitory compounds, although its predominant temporal influence in wound healing may be limited to the early inflammatory phases [97]. Vascular Endothelial Growth Factor There are four isoforms of VEGF, VEGF A through D. Keratinocytes, fibroblasts, and macrophages all produce VEGF, which is thought to be one of the main regulators of angiogenesis and vasculogenesis in wound healing. VEGF acts through two receptors in endothelial cells, VEGF-R1 and VEGF-R2. In adult wound healing, VEGF increases over time and has been associated with angiogenesis [71]. Also, direct and indirect targets of VEGF can affect wound healing [100,101]. However, through studies on fetal rats, scarless healing has an increase in VEGF expression three times higher than what is observed in late-gestation fetal wounds [73]. This work suggests that increased VEGF expression is partially responsible for the accelerated wound healing that occurs early in gestation, but ultimately its role may not be in regulating scarless wound repair. Fibroblast Growth Factors Embryonic wounds contain lower levels of fibroblast growth factors (FGFs), growth factors involved in skin morphogenesis [88]. The expression of FGFs, including keratinocyte growth factors 1 and 2, increases as the fetus ages, which suggests that these growth factors are profibrotic, similar to CTGF and TGF-b [102]. Many isoforms have been studied, including FGF-5, which doubles in expression at birth, FGF-7, which multiplies more than sevenfold at birth, and FGF-10, which doubles at the transitional period [71]. In general, downregulation of the FGF isoforms occurs during scarless wound healing, whereas the opposite is true during adult wound healing, which suggests that FGF upregulation is likely partially responsible for scar formation or at least managing cell differentiation and proliferation [71]. Studies suggest that FGFs, as well as Wnts, also have a role in organizing and differentiating embryonic organs [103e106]. Moreover, FGFs may be stored in the ECM by binding to proteoglycans, becoming a reservoir during tissue injury, when serum proteases cleave FGFs into functionally active isoforms and begin their signaling cascade [104]. In injured murine skin, FGF-9 initiates wound-induced hair follicle neogenesis through a feedback loop including two different fibroblast subtypes, the Wnt2a/b-catenin pathway, and FGF-10 [107]. Although studies in electroporation may provide an example of dermal appendage regeneration through ECM maintenance [108], these studies have not been verified, which makes FGF-9 the only single-cell signaling molecule capable of restoring some normal skin architecture during adult skin healing. Platelet-Derived Growth Factor Like FGF, PDGF has been identified as a profibrotic growth factor. Adult wounds contain high amounts of PDGF, whereas this growth factor is virtually absent in embryonic wounds. However, this is largely the result of inhibition of platelet degranulation in embryonic wounds [88]. Experiments involving the administration of PDGF to fetal wounds show that this growth factor induces scarring through increased inflammation, fibroblast recruitment, and collagen deposition [109]. Wingless Type Signaling Wnts are a critical regulatory cell signaling molecule existing in the form of secreted glycoproteins. Wnts, which are usually tethered to a lipid moiety such as a palmitate tail, usually have limited diffusion capacity and function in local tissue environments [110]. In embryos, Wnts have a major role in axial differentiation. In adults, Wnt responsive cells are found in all major tissue types and origins, which suggests Wnts continue to have important roles in adult tissue organization [111]. In adult skin, Wnt-responsive cells can be found in the epidermis and dermis. By expert opinion, Wnt signaling cascades in adult stem cells are thought to be responsible for ongoing cell fate, spatial

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recognition, and cell polarity, the importance of which is often understated [111]. In addition, specific Wnt proteins such as Wnt2a may be involved in the cell-to-cell communication necessary for specialized dermal structures to develop, such as new hair follicles in wounded adult skin, as was previously discussed [107]. Wnt expression in fetal tissue is usually higher than in adult tissue at baseline and often gradated [111,112]. With wounding, the expression of Wnts in fetal dermal tissue, as with many other growth factors and inflammatory cells, changes very little [112]; instead, the overwhelming number of organizational growth factors and Hox gene products leads to scarless tissue regeneration. In adults, Wnts increase during skin wound repair. However, their role in skin wound healing is less certain because Wnts (1) function mostly as a signal transducer and (2) are difficult to study [112]. Ultimately, Wnts are a promising although enigmatic choice for studying wound healing. Their only foreseeable barriers in clinical translation are appropriate dosing and delivery modalities, because Wnts also have a role in cancer development [106,107]. Interleukins The interleukins (ILs) are a class of cytokines involved in activating the inflammatory cascade; they are frequently targeted as wound healing adjuvants. IL-8 stimulates neovascularization and attracts neutrophils. IL-6 is produced by adult fibroblasts in response to stimulation by PDGF, and then activates macrophages and monocyte chemotaxis. With an insult to skin integrity, IL-6 and IL-8 rapidly increase expression to facilitate recruitment of circulating inflammatory cells to the wound. This elevated expression is maintained over 72 h during adult repair but is suppressed after 12 h during scarless fetal repair. Early fetal fibroblasts express lower levels of both IL-6 and IL-8 than do their adult counterparts at baseline and in response to PDGF stimulation. Therefore, these proinflammatory cytokines are thought to promote scarring. In addition, studies on the administration of IL-6 to fetal wounds induces scarring, which further supports this theory [65,113]. However, IL-10 is thought to be antiinflammatory, based on its antagonism of IL-6 and IL-8. Evidence for this comes from fetal embryonic day (E)15 skin in IL-10 knockout mice grafted to syngeneic adult mice. Incisional wounds on these skin grafts showed scar formation, whereas similar wounds in 15-day gestation wild-type skin grafts on adult wild-type mice healed scarlessly. These results suggest that IL-10 is essential for scarless fetal healing owing to its ability to dampen the inflammatory response [65,113]. In a supporting study, administration of an IL-10 overexpression adenoviral vector reduced inflammation and induced scarless healing in adult mouse wounds [114].

CURRENT THERAPEUTIC INTERVENTIONS No currently available therapy can induce postnatal regenerative healing. Although many therapeutic interventions are used to reduce scar formation, research has not adequately demonstrated efficacy or safety for many of these treatments, because of small treatment groups and a lack of well-designed studies. However, the following treatments are used clinically to reduce scarring symptoms and scar formation.

Topical and Intralesional Corticosteroid Injections Corticosteroids are commonly used to treat symptomatic scars; triamcinolone is the most commonly used agent. The mechanism of action is multifactorial. The inflammatory response is globally decreased, which secondarily decreases collagen synthesis and increases collagen degradation. Corticosteroids also inhibit fibroblast proliferation and TGF-b1 and b2 expression by keratinocytes [115,116]. Although 50e100% symptom improvement has been reported, studies are limited by a lack of appropriate controls and poor design [115,117]. The use of corticosteroids is limited by reported adverse consequences in 63% of patients. These effects include delayed wound healing, hypopigmentation, dermal atrophy, and scar widening [117]. Based on successful studies combining corticosteroid injections with 5-fluorouracil therapy and laser therapy, polytherapy may be the best method with which to use steroids, because lower dosing and fewer adverse effects occur [117,118]. Overall, an agent with global reduction in cell protein synthesis and proliferation is not ideal for achieving tissue regeneration.

5-Fluorouracil In the treatment of fibroproliferative scars, 5-fluorouracil (5-FU) has been used since the 1990s, with some promising clinical results [119]. This therapy, which represents a middle ground between the low side-effect profile of

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silicone gel sheeting and the morbidity of surgery, has emerged as a promising alternative for patients who wish to avoid surgery or who have failed conventional treatment algorithms. 5-FU has gained some negative attention because of its use as a chemotherapeutic agent. When given systemically, it has potent side effects though a longestablished safety profile. Its mechanism of action involves inhibiting the synthesis of the pyrimidine thymidine, thus inhibiting DNA synthesis in dividing cells and inevitably causing apoptosis [119]. In glaucoma and other proliferative diseases (such as keloids), 5-FU offers promising results when administered locally. When used alone, 45e78% of patients saw improvement in their scar appearance. However, when combined with triamcinolone, the efficacy of local administration improved scar appearance in 50e96% of patients [119]. It may also be combined with surgery to reduce fibroproliferative scar recurrences. Unfortunately, most clinical trials recruit small numbers or are confined to populations inherently at risk for fibroproliferative scarring [120,121]. Further research analyzing the efficacy of 5-FU in scarring is needed, but given its relatively benign local side effects and demonstrated benefits, it should remain an alternative for patients with recalcitrant scarring disorders [18,119,121e123].

Imiquimod Originally marketed as a treatment for verruca, actinic keratosis, and basal cell skin cancer, imiquimod may be used in low-morbidity settings for the treatment of hypertrophic scars and keloids. This Toll-like receptor-7 agonist functions by stimulating dermal inflammatory cells to secrete interferons, ultimately recruiting activated leukocytes to skin when applied topically. There, the effect is mostly mediated by the immune system. This medication represents an exciting, minimally toxic, and targeted method of immune modulation with few systemic side effects, and mostly blistering or skin irritation as the major local side effect. A few small, randomized controlled trials evaluated the efficacy of imiquimod in the treatment of acute surgical incisions in the breast [124] and in minor dermatologic surgeries [124]. However, the most trials failed to show a difference in scar appearance and potentially may have worsened outcomes by potentiating inflammation [124]. In the case of already formed keloids and hypertrophic scars, imiquimod could be tested in select patients wishing to try nonsurgical therapies, but ultimately this product does not appear to have a role in acute wound healing.

Laser Therapy In the 1990s, pulsed dye laser therapy emerged as a potential treatment for fibroproliferative and acute scars [125,126]. Like other disorders of aberrant proliferation, the mechanism proposed involves destruction of new blood vessels to limit, at a minimum, the erythematous appearance of some scars [126]. The idea behind targeting fibroproliferative scars with laser treatment comes from the principle that vascularity is partially responsible for their erythematous appearance. Although some case studies showed positive results in the treatment of keloids, the ischemic mechanism of keloid progenitor cell differentiation and proliferation is concerning [127]. Laser therapy has relatively few adverse effects (hyperpigmentation in 1e24% of patients and transient purpura in some). However, more research is needed to support its efficacy.

Bleomycin Another potent chemotherapeutic being used in dermal fibroproliferative disorders is bleomycin. This antibiotic has profound antibacterial, antiviral, and antitumor activity through DNA strand scission [128]. The exact mechanism of action is not entirely understood, but it is generally accepted that bleomycin induces cell death by forming complexes with iron and other metals that generate free radicals that eventual cleave both double- and singlestranded DNA, as well as degrade RNAs [128]. In resistant hypertrophic scars, keloids, and warts, bleomycin has been used off-label by dermatologists via intradermal injections [129]. Similar to other off-label use products, there are few clinical trials testing its efficacy [129]. However, if monitored closely with an experienced practitioner, it may be effective and safe. Side effects from treatment with bleomycin may be as minor as skin irritation, to as severe as skin necrosis and eschar formation [129]. When giving systemically, bleomycin notoriously causes pulmonary and skin fibrosis in humans and mice [130]. As such, it is often used to induce both conditions in an effort to create mouse models of pulmonary fibrosis and scleroderma [130]. Like other chemotherapeutic agents exploited in wound healing, development of a targeted therapy would be more attractive.

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Silicone Gel Sheets Maintaining tissue hydration in the base of wounds has long been known to improve outcomes in acute and chronic wound healing. Silicone, in either sheets or gels, emerged in the 1980s as a potential vulnerary agent by providing a hydrating barrier to open wounds. For decades, silicone has been used in deeper tissues, in the form of breast implants or as an interface for matrices used in hernia repair and large tissue defects; which highlight its safety [131,132]. Silicone gel sheets are hypothesized not only to hydrate wounds but also to inhibit collagen deposition and downregulate TGF-b2. Its development came from a series of studies that implicated dehydration of deep dermal fibroblasts as the mechanism for inducing scar collagen production. This therapy has been studied for both treatment and prophylaxis of excessive scarring [133]. Initial studies showed conflicting results in terms of efficacy [134,135], requiring further study. A review emphasized the weak evidence surrounding the use of silicone gel sheets in the treatment of keloids and hypertrophic scars [136]. However, silicone gel sheets will likely continue to be used because they are a noninvasive treatment with few adverse effects. In clinical practice, silicones are usually found as a component to a complex matrix or as a topical therapeutic dressing for chronic wounds [132]. Although the efficacy is unknown, they are safe to continue using and investigating.

Pressure Dressings and Negative-Pressure Wound Therapy Despite being in clinical use since the 1970s [137], pressure dressings have not been validated by experimental trials to be efficacious in prophylaxis or in the treatment of scars [138]. These treatments may be efficacious in reducing the appearance of a scar if used in polytherapy, but further investigation is warranted. Pressure earrings have been used in earlobe keloid excisions but have not been shown to eliminate recurrence. Pressure is also used in the form of negative-pressure wound therapy. Most surgeons are familiar with this technique, in which a vacuum is applied to an open wound (chronic or acute) to activate mechanotransducers in cells and potentiate cell proliferation and eventual wound closure by secondary intention [139]. The application of negative pressure in wound therapy appears to be ever increasing, with promising results as a treatment for acute wounds [140], chronic wounds [141], burns [142], and for the closure of large contaminated incisions [143]. Many modifications have been made to the vacuum system, such as the application of silicone sheets or foam pads, which ultimately make minor improvements in the overall technology.

Radiation Therapy Radiation therapy can also be used as an adjunct to surgical excision in the treatment of keloids. Mechanistically, it is thought to decrease collagen production by reducing fibroblast proliferation and neovascular bud formation. Radiation therapy is most effective for recurrent keloids if a single dose is given within 24 h of surgical excision, but studies have attempted short courses of radiation after healing with positive results, as well [144]. Radiation treatment appears to decrease recurrence rates after surgical excision from 45e100% to 16e27% of patients [145e147]. The magnitude of difference from these studies clearly highlights their limited power. One limitation of radiation therapy is that standardized dosing, fractionation, time period, and frequency after surgical procedures has not been developed. Reish et al. reported good results in treating recurrent keloids after surgery with 300e400 Gy in three to four fractions or 600 Gy in three fractions [138]. However, like many other studies in wound therapy, large, randomized, controlled studies are required to evaluate the role of radiation further in wound healing. Ideally, therapy that may cause oncogenesis should be avoided. Given the morbidity of radiation, this is not a popular method to reduce scar tissue formation.

Cryotherapy Cryotherapy has been studied in conjunction with surgical excision to treat keloids and hypertrophic scars. Many of these studies are limited by small sample size and poor controls, but the largest study reported a 79.5% response rate with 80% reduction in scar volume [148,149]. Cryotherapy is thought to decrease collagen synthesis and mechanically destroy scar tissue. Side effects include hypopigmentation and depressed atrophic scar formation [150]. This therapy can be used as an adjunct to surgery or as monotherapy, but studies question its efficacy in large recalcitrant keloids [151].

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Extracellular Matrix Substitutes and Scaffolds As the extracellular matrix is researched, it is becoming increasingly apparent that once regarded “inert” structural proteins such as collagen, vimentin, fibronectin, hyaluronic acid, proteoglycans, and glycosaminoglycans are in fact functional and can regulate cell growth [152]. With this discovery, scaffolds mimicking the ECM targeting acute or dysfunctional wound healing have emerged at a rapid rate. The most basic ECM scaffolds are made of single ECM compounds such as collagen or hyaluronic acid, at times integrated in bilayers with synthetics. Examples include Hyalomatrix, Promogran, and Biocol [152]. Acellular dermal matrices, human amniotic membranes, and porcine intestinal mucosa, which once were used for large incision closures and hernias, are also emerging in the market as wound healing adjuvants. These, too, can be combined with a number of additional therapies, from impregnated pharmaceuticals to cultured stem cells [153]. Generally, results are promising in the treatment of chronic and diabetic wounds [152]. However, large randomized controlled trials have not been performed comparing different ECM substitutes.

Tension Offloading An important field of study in wound healing focuses on mechanical forces. There are multiple examples of pressure, suction, or stretch that result in rapid cell proliferation, as is seen with vacuum-assisted wound closure and pregnancy [139,154]. In adult wound healing, tension initiates a signaling cascade leading to the proliferation of local cells typically in a symmetric pattern. Both keratinocytes and fibroblasts have mechanosensors imbedded in their cell membrane. These sensors, with other molecules that bind to the ECM such as integrins, are ion channels that open only when stretched [155]. In the case of fibroproliferative wounds, this may be an inciting factor leading to aberrant cell growth [139]. In fetal wound healing, the loss of dermal architecture leads to the deposition of actin protrusions that will contract and recruit local cells to close gaps in tissue and eventually regenerate lost structures [156]. This is perhaps the most stark difference in wound healing between prenatal and postnatal organisms, in which similar filaments and structures are involved but ultimately are used in entirely different ways [156]. In surgical patients, tension can have a severe impact on wound healing. On the back, chest, sternum, and tibia, incisions naturally stent open and can be predisposed to prolonged healing, infection, and dehiscence whereas areas with low tension and redundant tissue, such as the eyelid, heal with minimal scarring. A clinical trial tested an external tension offloading device known as “embraceÒ ” in acute scar revision [157]. Patients achieved significantly diminished scarring with its use.

Surgery Remodeling is a process that can last 1e2 years. During this time, scars can lose their dark pigmentation, flatten, soften, and contractures can lessen. Because scars can often behave unpredictably, surgery is usually reserved until after this period has passed. There are many options for surgical treatment for scarring, including excision with direct closure, local skin flap coverage, or more extensive vascular flap coverage. These treatments are generally considered before surgical treatment or as an adjunct. For fibroproliferative scars, surgical treatment is usually a simple excision, with or without flap closure, depending on the size of the resulting defect. In chronic wounds, burns, and pressure ulcerations, surgery may consist of creating local or pedicled tissue flaps, split-thickness skin grafts, or repeated surgical debridements to encourage healthy wound healing.

FUTURE THERAPEUTIC INTERVENTIONS Growth Factors and Cell Signaling Molecules As mentioned earlier, targeting individual cell signaling molecules has not translated to effective clinical treatments. In the case of recombinant TGF-b, clinical trials were arrested when the treatment failed to affect wound healing [92]. Monotherapy with solitary ECM components also usually leads to some treatment effect, but none that regenerates normal dermal architecture. Studying the signals influencing cell polarity and differentiation, however, has been more fruitful. FGF-9 is being developed as a therapy for hair loss [107] whereas its target, Wnt2a, may prove to be important in generating new hair placodes. WNt3a has already proved to be valuable in regenerating tissues, and in a mouse model of tissue regeneration, full-thickness excisional wounds taken from the ears closed whereas

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placebo treatments did not [158]. It is likely that if signaling cascades are targeted specifically via receptor agonists and antagonists, polytherapy will be needed to circumvent redundancies in the inflammatory pathway and signals leading to collagen deposition. These few examples represent the most promising monotherapies currently in use.

Targeting Gap Junctions and Connexins Gap junctions are hypothesized to function during wound repair by transferring injury signals from cell to cell, coordinating the inflammatory response, mediating wound closure, and regulating scar tissue formation in response to injury [46,159e161]. Many connexins (Cx) are present in the skin; the most extensively studied connexin is Cx43, which is expressed in both the epidermis and dermis [76]. Cx43 has decreased expression at the wound edge by the first 1 or 2 days after injury [159]. During wound repair, increased phosphorylation of Cx43 by protein kinase C may cause decreased gap junctional communication through a decrease in unitary channel conductance. This inhibition then initiates the injury-related response by the involved cell, which recognizes injury via reduced cell-to-cell communication [162]. When Cx43 antisense oligonucleotides were applied to mouse skin wounds, decreased inflammatory cell infiltration, fibrotic tissue deposition, and accelerated wound healing were observed [163]. Other studies showed that transient inhibition of Cx43 decreases scarring after burn injury in wild-type mice and increases reepithelialization after burn injury in human diabetic patients [163,164]. To further support these data, Cx43 knockouts have accelerated wound closure [165] and decreased collagen type I synthesis in the presence of chemicals that uncouple communication between cells. Interestingly, these treatments did not affect the levels of collagen type I mRNA [46]. Based on these data, the application of lithium chloride, a substance known to enhance signal propagation through gap junctions, produced the opposite effect: enhancing the deposition of granulation tissue, increasing open wound closure time, and increasing scar [166]. Given the strong correlation between Cx inhibition and improved wound healing, other therapies aimed at blocking signal transduction from cell to cell are under investigation. For example, a group at the Medical University of South Carolina synthesized a membrane permanent peptide containing a sequence designed to inhibit interaction of the ZO-1 protein with Cx43. This peptide, known as ACT1 peptide, decreases the rate of channel organization in gap junctions [167]. Through further investigation, researchers found that this peptide interacts with more than one portion of Cx43 and enhances cutaneous wound healing through decreased inflammation and scarring [160]. The advantage of this novel protein is that overall Cx43 expression is not altered. Moreover, the expression of other genes is not directly altered, unlike with antisense therapy and gene knockdown modalities. As with the TGF-b superfamily, several commercial companies are attempting to develop Cx-related scar reduction therapies. These therapies include Cx43 antisense-based gene therapy and ACT peptide bioengineering [167], which are in the early stages of testing and will not be available for some time.

Other Drugs and Biologics Additional strategies increase the expression of intrinsic antiscarring molecules at the wound site, including fibromodulin, hyaluronic acid, and hepatocyte growth factor [168e170]. Other approaches include treatment with inhibitors of MMP [171], inhibitors of procollagen C-proteinase [172], inhibitors of dipeptidyl peptidase IV enzymes [173], as well as treatment with angiotensin peptides [174]. Adenovirus-p21 overexpression has also been linked to scar reduction [175].

Stem Cells True skin regeneration at sites of injury has not been accomplished by single moleculeespecific therapy. As such, regenerative repair may require a cell-based approach. Stem cell therapy, with the ability to differentiate cells into various necessary cell types, is a promising approach to regenerative repair (Fig. 5.6). Embryonic Stem Cells Embryonic stem (ES) cells are those that may be isolated from the embryo and possess both pluripotency and the capacity for unlimited self-renewal. The discovery of ES cells was the result of work in the 1970s involving transplanting embryonic cells into ectopic sites, which inevitably resulted in teratomatous tumor formation [177]. The discovery of specific embryonic cells that self-renew indefinitely in culture and can generate an entire organism

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FIGURE 5.6 Stem cellebased therapies to promote scarless wound healing. Representation showing general principles of two cell-based therapeutic methodologies: (1) application of stem celleconditioned media, and (2) direct application of stem cells to the wound bed. The poor survivability of mesenchymal stem cells (MSCs) transplanted to the wound bed has prompted the development of other novel therapies that take advantage of the paracrine mechanisms of action of these cells. Application of conditioned media from umbilical cord bloodederived MSC culture is one such example [178]. Reproduced with permission from Leavitt T, Hu MS, Marshall CD, Barnes LA, Lorenz HP, Longaker MT. Scarless wound healing: finding the right cells and signals. Cell Tissue Res 2016; 365(3):483e93.

or specific cell types led to the concept of the totipotent ES cell [179,180]. It was immediately recognized that these cells held the potential to be used for regenerating dysfunctional organs and tissue. Equally obvious were the potential ethical problems related to using cells from human embryos. Whereas other types of stem cells have been used successfully in clinical therapies, such as hematopoietic stem cell transplantation for hematologic diseases and malignancies, therapies involving ES cells remain entirely experimental. In an early attempt at tissue regeneration using ES cells, Fraidenraich and colleagues injected wild-type ES cells into mice predisposed to death by cardiac failure. The remarkable result of this experiment was that the mice were rescued from their lethal phenotype through a mechanism that appeared to involve signaling factors released by the ES cells [181]. A potential barrier to using therapeutic ES cells in humans is the host immune response against these foreign cells. Immunocompetent mice mount a robust immune response against human ES cells, leading to limited ES cell survival, although this response can be mitigated with the use of immunosuppression [182]. Another difficulty with using ES cells in human therapy is the formation of teratomas; this is unsurprising because the origin of the ES cell’s discovery was in the study of teratoma-forming embryonic cells. Undifferentiated ES cells transplanted into ectopic sites, such as the heart, consistently form teratomas [183]. However, advances in the directed differentiation of ES cells toward specific fates should make teratoma formation less of a concern [184]. Improvements in the purifying and characterizing ES cells have allowed for trials testing their efficacy in human diseases. In a phase I/II clinical trial, patients with macular degeneration were treated with retinal pigment epithelium cells derived from human ES cells and with immunosuppression [185]. Several patients experienced improvements in visual acuity up to 12 months after treatment, a result that would not be expected in the absence of treatment. Human oligodendrocyte precursor cells derived from human ES cells showed promise in treating spinal cord injury in preliminary rodent experiments [186]. However, human spinal cord injury patients treated with similar cells in a phase I trial showed no improvement after treatment [187]. These results highlight that promising results in animal models are not easily translated to human applications, and that novel stem cell therapies must be tested rigorously in human trials. In addition to macular degeneration and spinal cord injury, diabetes, myocardial infarction, and Parkinson disease are among human diseases that are the subjects of ongoing clinical trials involving ES cells [187]. Mesenchymal Stem Cells Mesenchymal stem cells (MSCs) are nonhematopoietic bone marrow stromal cells that were initially isolated based on their ability to adhere to plastic culture plates. These cells are unique in that they differentiate into

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mesenchymal lineages such as cartilage, fat, muscle, and bone [188]. MSCs are a heterogeneous group of cells that have had populations isolated not only from the bone marrow but also from adipose tissue and amniotic fluid. Based on their ability to expand in vivo and differentiate into multiple mesenchymal tissue types, these cells are thought to be an ideal source of autologous stem cells used for promoting wound healing and/or scar-reducing therapies [189]. MSC therapy avoids rejection and the ethical and moral concerns associated with ES cell therapies. MSCs may affect wound healing and tissue regeneration through many different avenues. Once transplanted, these cells migrate to the site of injury or inflammation, where they may stimulate the proliferation and differentiation of resident progenitor cells, secrete growth factors, participate in remodeling [188,190], and modulate the immune and inflammatory responses in the wound bed [191]. A wealth of clinical data attest to the safety of bone marrowederived MSCs, and emerging data support adiposederived mesenchymal cells as possessing a similar safety profile to bone marrowederived MSCs [192,193]. Therefore, MSCs could be used to affect various pathways involved in wound healing, including inflammation, aging, and cellular senescence. Several examples of human wound healing investigations exist using MSCs. The first was a small trial using a fibrin glue vehicle in both acute and chronic wounds. This study demonstrated that topical application of autologous passage 2e10 bone marrowederived MSCs, combined with fibrin spray, allowed acute surgical wounds and chronic lower extremity ulcers to heal faster. The wound healing speed increased in a manner directly proportional to the number of cells applied [194]. A larger study evaluated patients with various nonhealing wounds. Bone marrowederived MSCs were applied with a dermal scaffold to wounds, with or without autologous skin grafts. Results showed accelerated healing in wounds treated with MSCs [195]. A limitation of this study was that the cells were not passaged and flow cytometry was not used for isolation. MSCs are known to represent only 0.001% of nucleated cells in the bone marrow; therefore, the cell population used in these experiments was likely heterogeneous and may have contained tissue macrophages that would also assist in wound healing [188]. In a randomized controlled trial in which patients with critical limb ischemia received injections of either allogeneic circulating MSCs or control solution, no difference in outcomes was seen [196]. In a later randomized trial, 24 patients with nonhealing ulcers were randomized to receive autologous cultured MSCs or control treatment [197]. Those treated with MSCs experienced greater symptom improvement and ulcer healing compared with the control group. Several small, nonrandomized human pilot studies suggested improved healing rates in patients with limb ischemia after treatment with bone marrowederived mononuclear cells [198e205]. Designed to expand on these results, the JUVENTAS trial is perhaps the most prominent clinical investigation to date based on bone marrowe derived cell therapy [206]. This trial included 160 patients with critical limb ischemia thought to be nonrevascularizable. Subjects received either injections of bone marrowederived MSCs or placebo injections. There was no significant difference between groups in rates of amputations or quality of life measures. These results confirm the importance of verifying promising preliminary results with well-designed, randomized trials. Adipose-derived stromal cells (ASCs) are a potentially promising source for therapeutic cells because they are easily isolated from a patient’s own fat tissue and exhibit osteogenic and adipogenic activity in vivo and in vitro [207]. Data from animal models indicate that autologous and allogeneic ASCs can regenerate tissue after injury [208]. Two noncontrolled human trials with small numbers of patients demonstrated improved wound healing after the administration of ASCs [209,210]. However, these results must be interpreted with caution because of the absence of control groups. In a small randomized human trial, a product containing ASCs and fibrin glue accelerated closure of perianal fistulae, which share similarities with nonhealing skin wounds [211]. However, a larger follow-up study showed no improvement with ASC administration compared with control treatment [212]. Further research is needed to characterize MSCs and their niches. As purification and enrichment techniques improve, the role of MSCs in wound healing will gain clarity. Defining the direct role of MSCs in wound repair, as well as their effects on other cells, will guide their future therapeutic potential. Epidermal Stem Cells As mentioned previously, the epidermis in humans is a dynamic structure undergoing constant renewal. Epidermal turnover is estimated to take place over 60 days in humans, a process that requires a continuous supply of differentiated cells. Epidermal stem cells have a high capacity for self-renewal, as evidenced by the number of daughter cells that undergo terminal differentiation into keratinocytes [213]. Several stem cell niches are present in the epidermis. The best-characterized are the interfollicular epidermal stem cells and the hair follicle bulge region.

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The interfollicular epidermis (IFE) is the region of epidermis located between hair follicles. Under normal homeostatic conditions, IFE stem cells defined by expression of the gene Lrig1 of the basal layer of the epidermis divide at a steady rate to provide new keratinocytes to populate the epidermis [191,214]. These cells may also contribute to the growth of hair follicles and sebaceous glands [214] Whereas the IFE can receive contributions of cells from other structures, such as the hair follicle, it is also capable of repairing and renewing itself after injury in the absence of these other cells [191]. A separate population of epidermal stem cells defined by embryonic expression of Lgr6 represents a primitive stem cell population that establishes all lineages of the skin, including cells of the hair follicle, sebaceous gland, and interfollicular dermis. In postnatal life, Lgr6-expressing cells reside above the hair follicle bulge and contribute to maintaining the sebaceous gland and the IFE [215]. The hair follicle is a complex structure with several distinct regions whose cellular composition has been exceptionally well-studied. Each regions appears to contain a unique population of hair follicle stem cells (HFSCs). It is clear that HFSCs can repair the hair follicle itself after injury. However, an unresolved question is to what extent HFSCs are critical for regenerating the injured epidermis. The first hair follicle stem cells to be discovered were those residing in the hair follicle bulge region. These cells are characterized by expression of the genes Krt15, Lgr5, and Gli1 [191,216]. Initial experiments showing that these cells are present in the epidermis after a scratch injury suggested that they may participate in epidermal repair [217]. However, subsequent experiments showed that the presence of bulge cells in the epidermis is short-lived. It is likely that the contribution of these cells to long-term skin repair is minimal, and that their main role is to regenerate the hair follicle [191]. The junctional zone of the hair follicle is located above the bulge and adjacent to the sebaceous gland. It contains a complex population of stem cells that generally express Lrig1 but otherwise have different gene expression profiles and roles in regeneration [191,214]. Those that express Lgr6 in the embryo form the hair follicle, sebaceous gland, and interfollicular dermis. Postnatally, these cells contribute to repair of IFE and hair follicles. Because these Lgr6-positive cells are capable of forming several skin structures, they may represent a primitive skin stem cell [215]. Given the important role of epidermal stem cells in skin formation and regeneration after injury, the possibility of using them in a therapeutic role after injury is intriguing [216,218]. Preliminary experiments in animal models have yielded some promising results, such as epidermal stem cells transplanted onto rat wounds [216,218]. However, a major limiting factor in the therapeutic use of epidermal stem cells is their scarce availability and the difficulty in obtaining them. In patients most in need of new keratinocytes, burn victims, there may not be enough autologous cells left in certain situations to make this a viable clinical tool. At this point, the therapeutic potential for epidermal stem cells is largely theoretical, but research continues to develop at a rapid pace [219].

Induced Pluripotent Stem Cells Difficulties inherent in deriving and using human ES cells led to interest in generating pluripotent cells from other sources. In a landmark paper, Takahashi and colleagues described the transformation of adult dermal fibroblasts into induced pluripotent (iPS) cells using a specific combination of transcription factors [220]. Generation of iPS cells has been achieved from other human cell types, such as keratinocytes [221]. Using similar processes, fibroblasts may be transformed directly into functional, differentiated cell types such as neurons and cardiac myocytes [222,223]. The potential to derive iPS cells from adult tissue avoids the logistical and ethical problems associated with ES cells. Theoretically, iPS cells may also be used in an autogeneic fashion, avoiding the issue of immune rejection. Translating the potential of iPS cells into human therapies has been challenging. In the only human trial to date involving the therapeutic use of iPS cells, sheets of retinal pigment epithelium cells derived from iPS cells were implanted into the retina of a patient with macular degeneration. Although the patient reportedly experienced improved vision, the trial was halted owing to concerns about mutations detected in the iPS cells [224]. Advances in molecular biology have the potential to refine and improve methods for generating therapeutic iPS cells. For example, the ability to edit specific genetic sequences using clustered regularly interspaced short palindromic repeatseCas nuclease 9 may allow for the generation of iPS cells with specific traits that could increase their utility in specific disease processes [225]. In addition, the ability to generate functional organoids from iPS cells may prove to be useful for replacing dysfunctional organs in humans. Various groups have used iPS cells to generate organoids resembling cerebral cortex [226], intestine [227], and kidney [228], among others [229] (Fig. 5.7).

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FIGURE 5.7 Potential therapies for reducing scar formation during wound repair. To manipulate wound repair to become more regenerative than scar forming, strategies include the use of biomimetic scaffolds, manipulation of the mechanical environment (for example, negative-pressure wound therapy to increase healing) or the electrical environment, the administration of small molecules, the use of gene therapy approaches, and the use of cell-based strategies (including administration of epithelial stem cells). All of these elements have been demonstrated to have an effect on in vitro and in vivo models of wound healing as single-agent therapies. In theory, many of these elements could be combined to recreate a receptive environment (or “soil”) to promote regeneration. Combining these with the appropriate stem cells (or “seed”) will undoubtedly alter the result of wound healing in humans. Reproduced with permission from Gurtner GC, Werner S, Barrandon Y, Longaker MT. Wound repair and regeneration. Nature 2008;453(7193):314e21.

PERSPECTIVE Our understanding of adult stem cell populations, the complexity of embryonic development, and the redundancy of the adult skin wound healing mechanism has revealed that targeting a single-cell signaling cascade will not be sufficient to recreate scarless wound healing in adult mammals. Rather than the depth of our understanding leading to radical innovations, research has revealed a complex web of interactions among tissue types, structural proteins, developmental and highly evolutionarily conserved growth factors, external forces, and cell-to-cell interactions. On an optimistic note, the number of products for improved wound healing, both acute and chronic, has exploded, giving patients new options and hopes for improved outcomes and wound closure. Unfortunately, however, skin regeneration remains elusive. With the US population aging and accumulating comorbidities, new solutions are needed that combine scar treatment strategies to prevent the cost of wound healing complications from increasing. Perhaps most exciting, our

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understanding of stem cell niches and engraftment issues has improved to the point where autologous and allogenic stem cell therapy may become a therapeutic reality. If scarless wound healing cannot be accomplished in healthy adults by activating appropriate “self” signals, it is possible that engineered and cultured grafts may provide the necessary leap to accomplish regenerative healing.

List of Abbreviations 5FU 5 fluorouracil ASC Adipose-derived stromal cells CTGF Connective tissue growth factor Cx Connexin E Embryonic day ECM Extracellular matrix ES Cell embryonic stem cell FGF Fibroblast growth factor HFSC Hair follicle stem cell IFE Interfollicular epidermis IL inTerleukin iPS Induced pluripotent stem MMP Mixed metalloproteinase MSC Mesenchymal stem cell PDGF Platelet-derived growth factor Smad3 Mothers against decapentaplegic homolog 3 TGF-b Transforming growth factor b VEGF Vascular endothelial growth factor Wnt Wingless type

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6 Progenitor and Stem Cell Heterogeneity: Using Big Data to Divide and Conquer Melanie Rodrigues, Paul A. Mittermiller, Jagannath Padmanabhan, Geoffrey C. Gurtner Stanford University, Stanford, CA, United States

INTRODUCTION Traditional analysis of cells has relied on pooling of RNA or protein from hundreds of thousands of cells and displaying aggregate or average expression of replicate samples. These tools, which include polymerase chain reaction (PCR), microarrays, and western blotting, have been powerful in unfolding transcriptional networks, signaling cascades, and metabolic pathways, to advance our knowledge of disease and therapy. However, the underlying assumption with these techniques is that the population average represents the dominant biological state within the population. This assumption is flawed in most cases because the activity of each cell of the population is not reflected by the population average. Importantly, population averages are unable to capture the activity of rare but critical cells such as stem cells, or transiently amplifying cells such as progenitor cells (Fig. 6.1A). Interrogation of these rare but critical cells could prove to be key regulators of disease progression and therapy. With the evolution of single-cell technologies such as high-throughput sequencing (HTS), it is possible to evaluate single cells with a high degree of comprehensiveness. These technologies provide information about transcript expression, gene fusions, mutations, or single-nucleotide polymorphisms in individual cells. Outlier cells are no longer considered errors in measurement by default, but can be tested for the presence of a unique function. It is also possible to detect whether a cell population is homogeneous or if cells display heterogeneity by clustering the single-cell data into subpopulations of cells that exist in metastates. These subpopulations can then be tested for functional relevance in tissue homeostasis, repair, and disease (Fig. 6.1B). In homeostasis, cellular subpopulations function in a stable yet adaptable population equilibrium [1,2]. For example, in the skin, epithelial cells that display a high turnover rate are maintained by distinct subpopulations of self-renewing epithelial stem cells [3]. Similarly, in the bone marrow, the transcriptional, epigenetic, and functional heterogeneity of hematopoietic stem cells determines their cell cycle potential and differentiation ability [4,5]. A second level of complexity arises in tissue injury and repair. After tissue injury, various cell types need to be activated in spatiotemporal concert to bring about healing. In wound repair, for example, multiple cell types within the epidermis, dermis, hypodermis, and circulation must coordinate to bring about healing [6]. Single-cell technologies have revealed heterogeneity within several of the cell types involved in this repair process [7]. However, in tissues such as the heart and the brain, resident cells are still largely considered homogeneous owing to limited access to these tissues in the normal human state. Cellular heterogeneity also has a critical role in disease and consequences for how the diseased tissue is diagnosed and treated. In cancer biology, tumors display intertumor heterogeneity in which genetic and phenotypic variations are observed between individuals with the same type of tumor, or between tumors in different tissue types of the same individual. However, tumors also exhibit intratumor heterogeneity in which cells within the same tumor exhibit differences in gene expression, cellular morphology, motility, proliferation, metabolism, metastatic potential, and recurrence potential. Alterations in these cellular characteristics can be determined using next-generation HTS

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FIGURE 6.1 Population assays versus single-cell assays. In population-averaged techniques such as traditional microarray analysis, messenger RNA (green) from all cells (black circles) are pooled together and the aggregate expression level is reported. When studying heterogeneous populations, such as in cancer and stem cell biology, this approach can lead to considerable loss of information (A). Using single-cell analysis, it is possible to determine whether the population of cells can be further subdivided into subpopulations that are distinct from each other (B).

aided by computational analysis. These technologies have also made it feasible to detect mutational burden and the temporal order of mutations within tumors [8]. Traditional population assays cannot detect cellular heterogeneity because they obscure the response of individual cells and cellular subpopulations. Population assays also require hundreds of thousands of cells to be analyzed, which makes it difficult to study cellular heterogeneity in relatively small samples that are hard to obtain. In comparison, single-cell technologies can determine accurately and comprehensively differences at the genomic,

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transcriptional, translation, or epigenetic level. These technologies can assay rare cells. In addition, they allow for analysis of cells in an unbiased manner without the use of markers a priori. Although single-cell technologies are associated with several advantages, it is necessary to determine the need for resolution at the single-cell level. This is critical considering that the human body is composed of approximately 37.2 trillion cells [9] and each cell generates large datasets of information that must be analyzed computationally and mathematically. Once the need for single-cell analysis is established, it is essential to understand whether the biologic question requires a broad net to be cast to capture information about cellular behavior, or whether a targeted approach is required to reveal only certain cellular features with accuracy and precision. Thus, accuracy, precision, and comprehensiveness determine which technique is selected to answer the biologic question. Accuracy is the measure of certainty or validity that the measured value is close to the true value [10]. It is estimated by comparing the measurements generated by the single-cell technique with a reference standard technique (such as quantitative PCR for transcriptomic analysis). Precision is the ability to replicate or reproduce a measurement. In transcriptomics, high precision is associated with a narrow distribution of gene expression values [10]. Comprehensiveness or sensitivity is the amount of information obtained per cell. In transcriptomics, comprehensiveness refers to the total number of genes that can be detected [10,11]. Advanced single-cell technologies such as RNA sequencing (RNA seq) can detect more than 5000 genes per cell [12]. In addition to these three criteria, considerations such as cost-efficiency, sample size effects, and false discovery are also important for deliberation when comparing and developing HTS techniques [11,13]. This chapter describes the rapidly evolving single-cell technologies that have been used to study cellular heterogeneity. It explains how cellular subpopulations are discovered without prior bias. Finally, it addresses the impact of single-cell technologies on biology, regenerative medicine, and cellular therapy.

SINGLE-CELL ISOLATION To study heterogeneity and gather data on a single-cell level, cells need to be isolated with accuracy. Most genomic and transcriptional analyses assay nuclear content from individual cells and do not require the isolated cells to be alive. On the other hand, proteomic and metabolic measurements require live cells, which makes single-cell isolation a more challenging process. The classic method of isolating single cells involves manual pipetting followed by serial dilutions. Manual pipetting works only on microliter volumes, is cumbersome, and is difficult to scale. This has led to the use of robots such as the Mosquito HTS, which ensures repeatable nanoliter pipetting irrespective of the viscosity of the cell suspension or environmental conditions [14]. Automated liquid handlers can also scale experiments with ease, converting a 96-well assay into a 1536-well format [14]. In situations in which cells need to be isolated directly from a tissue sample or a surface, laser-capture microdissection can be used [15]. This technique requires a trained histologist to isolate the cells of interest accurately based on morphological properties or changes in fluorescence. To increase the speed of this technique, an automated laser microdissection technique called laser-enabled analysis and processing has been developed [16]. However, laser microdissection can be cumbersome, it may rupture surrounding cells, and it may isolate individual cells incompletely. For these reasons, this technique is not frequently used to obtain cells for HTS. For decades, fluorescence-activated cell sorting (FACS) has been the default technique to sort cells into welldefined populations. It has become the most widely used technique to isolate single cells for high-throughput studies. FACS can rapidly sort single cells at a high level of purity into 96- or 384-well plates at 10,000 cells per second, a pace that is impossible to match manually [17]. FACS systems work by passing cell suspensions through a small nozzle (70-100 mm in diameter) that creates a continuous stream of droplets, each droplet of which contains a single cell. Electrically charged plates are then used to deflect the droplets containing cells of interest into a microwell plate based on the physical, chemical, or optical properties of the cells that are enhanced using antibodies conjugated to fluorescent probes [15]. Sorting single cells through FACS requires a priori knowledge of surface markers for specific cells and can be difficult to perform in scenarios in which there are limited cells, such as early-stage reproduction and characterization of stem cells (Fig. 6.2). Moreover, the high flow rate and pressure can potentially damage delicate or large cells such as adipocytes as they move through the machine [18]. Most important, FACS does not allow for tracking an individual cell over time, which makes it difficult to determine the dynamic behavior of a cell. Microfabrication, precision engineering, and rapid prototyping techniques have led to the development of microfluidic devices that allow for sorting low cell volumes on miniaturized devices with rapid throughput [19]. These devices contain thousands of integrated fluid channels, valve control surfaces, and reservoir chambers to separate individual cells [2]. In addition, they reduce the size of the sorting equipment and eliminate biohazardous aerosols

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FIGURE 6.2 Enriching single cells using fluorescence-activated cell sorting (FACS). FACS can reliably test up to 15 proteins allowing for determination of signaling pathways and enrichment of cells based on surface markers. However, surface markers for enrichment are selected a priori and it is important to determine the best surface marker that can reliably enrich cells for therapeutic use.

[20]. Microfluidic technologies can be either active or passive. Active systems use external fields such as acoustics, electric, magnetic, or optical to displace individual cells into microchambers. The commercially available dielectrophoresis array, for example, uses a nonuniform electric field to exert forces on cells suspended in a liquid and trap them into “cages.” This allows for the separation of cells of interest from complex, heterogeneous samples. Passive systems, on the other hand, use inertial forces, filters, and adhesion mechanisms to sort cells [18,20,21]. The commercially available C1 chip from Fluidigm Corporation (South San Francisco, CA) is an example of a passive system in which hydrodynamic traps in a microfluidic channel allow for the isolation of a single cell from a heterogeneous sample. The cell is then subjected to RNA isolation and complementary DNA (cDNA) synthesis within the trap. This system can be scaled to accommodate many traps (up to 800 for the C1 chips), but the largest disadvantage is that the cell pool must be homogeneously sized because the microfluidic traps are sized for an average cell. If the cells are too large, they will rupture before reaching the trap. If the cells are too small, the traps will capture doublets or multiplets [22]. Droplet microfluidics is an emerging field of microfluidic technology in which active or passive systems are used to capture cells within discrete micrometer-sized aqueous (micro)droplets. The flow rate of the aqueous fluid determines the number of droplets that are unoccupied by cells. Each occupied droplet acts as a miniaturized reaction vessel and allows thousands of single cells to be analyzed in parallel every second. Whereas the droplets allow for fluorescent imaging and PCR analysis, they can also keep the cells live for a few hours to enable the analysis of dynamic cellular behavior such as enzyme kinetics or the response to drugs, antibiotics, biologic, and environmental factors. The commercially available 10X platform for example, uses droplet technology for single-cell genome sequencing and RNA seq [23]. In the RNA seq platform, cells are run through a microfluidic chip and single cells are individually encapsulated within gel beads. Each bead has a unique bar code to identify the source cells. The cells then undergo reverse transcription to create bar-coded cDNA. All cDNA from a single cell share the same bar code. The cDNA is then fragmented, amplified, and sequenced. Full sequences are reconstructed from the short-read sequences and expression levels for single cells are determined based on the number of transcripts expressed with the same unique bar code. This technology has been employed to investigate chimerism in immune cell populations, screening for Clustered Regularly Interspaced Short Palindromic Repeats interference, and testing single nucleotide polymorphisms in noninvasive prenatal testing, among other applications [23e25].

ACQUIRING SINGLE-CELL DATA Once individual cells are collected, single-cell analysis can be performed. The flow of biological information in cells is from DNA to RNA to protein, with DNA containing the genetic information of all cellular organisms, with RNA functioning as the messenger of this information and proteins acting as the working force to determine the phenotype of the organism. Random mutations bring about changes to the DNA. Injury and disease bring about changes in the transcribed RNA and protein. Sometimes there are modifications in the phenotype with no changes to the DNA sequence owing to changes in gene expression. These alterations, called epigenetic changes, can be

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influenced by various conditions including age, the cellular microenvironment, or disease states. This section describes the various high-throughput techniques developed to study changes in the DNA, RNA, protein, or epigenetics of individual cells.

Single-Cell Genomics Next-generation DNA sequencing allows for the comprehensive study of minute amounts of DNA from an individual cell. This technology has uncovered the evolutionary history of cells, genomes of unculturable microorganisms, and genetic mosaicism in normal physiology, disease, and cancer [26,27]. Genetic mosaicism is the presence of two or more populations of cells with different genotypes within the same individual, developed from a single fertilized egg. Although it has long been known that cancer is a mosaic disorder, in somatic tissues, early studies of genetic mosaicism were limited to abnormalities in skin development such as epidermolysis bullosa and ichthyosis. However, the advent of single-cell DNA sequencing unearthed variations in the chromosome, copy number, and single nucleotide sequences in a variety of tissue types, and mosaicism is now seen in a diverse range of clinical disorders that can be gonadal, somatic, or gonosomal [28]. To obtain single-cell genomic data, single cells are isolated and subjected to whole-genome amplification. This can be achieved through either PCR amplification or isothermal methods such as multiple displacement amplification [18]. The isothermal methods have greater coverage or comprehensiveness compared with the PCR-based methods; however, they display lack of uniformity or accuracy and precision [26]. Therefore, most single-cell genomic approaches use a hybrid method that consists of a limited isothermal amplification step followed by PCR amplification. These hybrid techniques include displacement degenerate oligonucleotide-primed (DOP)-PCR and multiple annealing and looping-based amplification cycles (MALBAC) and differ based on whether degenerate primers or random primers are used for amplification. DOP-PCR and MALBAC have uniform coverage that results in greater sensitivity and accuracy [29,30]. Once the genome is amplified, it is subjected to single-cell exome sequencing or entire-genome sequencing [31]. An important consideration while sequencing is that false variants increase as the size of the genome region increases. Furthermore, errors can be introduced in any stage, including the single-cell isolation and whole-genome amplification steps. Hence, it is imperative to develop tools that differentiate technical aberrancies and noise. The quality metrics include visual conformation of isolated cells as well as quantification of the whole-genome amplification product.

Single-Cell Transcriptomics Whereas genomics allows for the identification of genetic alterations within cells, transcriptomics allows the investigator to understand changes in the function of cells. Single-cell transcriptional studies emerged with the integration of single-cell quantitative polymerase chain reaction (qPCR) into microfluidic platforms, allowing for massively parallel qPCR reactions on a small chip [32]. The Biomark chip from Fluidigm Corporation, for example, provides a highly sensitive platform that allows for probing of individual cells for the expression of 96 select genes, providing a readout in less than 24 h after sample collection. This platform enables cDNA volumes to be detected that are 1000 times less than those required for traditional qPCR reactions. In this method, a single cell is sorted by FACS into a well of a 96-well plate. The cDNA conversion step is combined with a low-cycle reversetranscriptase (RT)-PCR step that preamplifies the cDNA with the select 96 primers before it is loaded onto a microfluidics chip. The microfluidics chip is subjected to qPCR in a Biomark machine (Fluidigm Corporation) where the cDNA from each cell is amplified by each of the 96 primers, which results in 9216 data points (Fig. 6.3A). This microfluidic-based qPCR technology generates highly accurate and precise gene expression data and has been used to determine heterogeneity among a wide range of cell types including fibroblasts, adipose stromal cells, glioblastomas, and hematopoietic stem cells [2,33e36]. The main downside of this method is that the number of genes that can be evaluated depends on the size of the microfluidics chip [37]. HTS of whole transcriptomes through RNA seq enables profiling of the transcriptome. This technology enables the detection of absolute levels of gene expression, gene fusions, single nucleotide variants, insertions, deletions, and weakly expressed genes. Unlike hybridization methods such as microarrays, RNA seq does not require prior knowledge about the organism (species or transcript-specific probes), which enables the identification of both known and novel transcripts. It also provides a signal-to-noise advantage over microarrays by eliminating cross-hybridization, nonideal hybridization, and DNA contaminants. DNA contaminants are negated by unambiguously mapping DNA

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FIGURE 6.3 Single-cell transcriptomics. Schematic of high-throughput, microfluidic chip-based, single-cell transcriptional analysis demonstrates how a single cell is sorted by fluorescence-activated cell sorting (FACS) into each well of a 96-well plate that has been preloaded with reverse transcriptaseepolymerase chain reaction (RT-PCR) reagents. A low-cycle RT-PCR preamplification step creates complementary DNA (cDNA) for each individual cell. Single-cell cDNA is then loaded onto the microfluidics chip along with the primer-probe sets for 96 gene targets and quantitative polymerase chain reaction is performed in the Biomark machine, leading to 9216 data points per chip (A). Schematic of single-cell RNA sequencing demonstrates that cells first need to be isolated and enriched by techniques such as FACS. Microfluidic technologies are then used to isolate single cells. Reverse transcription of messenger RNA is performed to produce cDNA that is amplified through PCR. The amplified cDNA is sheared to reduce the length of the sequences and the fragments are sequenced to the number of desired reads. Finally, the sequences are reconstructed and matched to a known library and the copy numbers of transcripts are determined (B).

sequences to unique regions of the genome. Based on its sensitivity and range of expression, there has been an overwhelming interest in using RNA seq on single cells to determine heterogeneity within cell populations, identify rare cells, and characterize poorly defined cells. Although individual protocols vary slightly, the overall processing of single-cell RNA seq is similar. First, template switching by reverse transcription of messenger RNA (mRNA) is performed to produce cDNA that is amplified through PCR. The amplified cDNA is sheared to reduce the length of the sequences and the fragments are

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sequenced to the number of desired reads. Subsequently, the sequences are reconstructed and matched to a known library and the copy numbers are determined [37]. This method allows sequencing of all mRNA molecules, resulting in an unbiased approach to evaluating the transcriptome and allowing for the discovery of novel transcripts or splice variants (Fig. 6.3B) [38e40]. Despite the comprehensiveness of the technique, single-cell RNA seq is associated with a myriad of technical issues that result in uncertainty of the generated data. Most microfluidic platforms require several thousands of cells to be loaded in highly concentrated solutions as starting material before the cells are distributed into individual traps or microdroplets. In most cases, these cells are freshly sorted by FACS to obtain a homogeneously sized population. Although tested for viability before loading into the microfluidic chamber, it is unlikely that all the FACS-sorted cells are live at the start of the assay. In addition, microfluidic separation of cells may cause cell rupturing or the microchambers may capture doublets or multiplets. Several technologies do not allow for the visualization of entrapped cells within the traps or microdroplets before cDNA synthesis. If doublets are captured, it is difficult to distinguish the data obtained from a single cell. Similarly, if the cells are ruptured and only part of the RNA becomes reverse transcribed, it is difficult to separate these data from those of an intact single cell. Many sequencing technologies use spike-in RNA controls (either external RNA controls consortium or custom) to determine the quality and success of the library constructed. These spike-ins provide transcripts at a known sequence, length, and concentration and serve as a quality control for the isolated single-cell RNA. Although used as controls, in several cases, the spike-ins compete and become reverse transcribed instead of the cellular mRNA. mRNA by itself is a delicate molecule and easily prone to degradation. Therefore, strict quality control measures of fragment range in the Bioanalyzer or Experion is helpful.

Single-Cell Proteomics Proteomics provides information about the biochemical activation state of the translated protein. The activation of a protein is represented by its phosphorylation, acetylation, proteolytic cleavage, ubiquitylation, change in localization, conformation, or abundance within cells [41]. Traditionally, these processes have been analyzed by population assays such as western blotting and immunofluorescence imaging. However, the analysis of these changes in single cells is challenging because of the paucity of protein amplification techniques [18]. FACS has been the most wellestablished technique for determining the relative expression of surface proteins on single cells; it distinguishes individual cells within a mixed population of cells and enriches live cells based on surface markers [42]. FACS can reliably test up to 15 proteins and enables the determination of signaling pathways and networks within individual cells. However, these are markers selected a priori, and it is important to understand which is the best surface marker that can enrich cells before therapeutic use (Fig. 6.2). When the cells are fixed and permeabilized, FACS also helps in determining the intracellular activation state within individual cells. Most commercial flow cytometers require manual sample preparation of cells; however, improvements have been made in microfluidic platforms that allow for handling, sorting, and flow cytometry [42]. Groups of cells can also be fluorescently bar coded with unique signatures of fluorescent dyes, so that they can be mixed together, stained, and analyzed as a single sample. This results in the possibility of high-throughput FACS analysis. Fluorescent bar coding reduces antibody consumption by 10-fold to 100-fold and minimizes pipetting error, staining variation, and the need for normalization. The robustness of data is increased through the combination of control and treatment samples, and the speed of acquisition of data is enhanced [43]. Mass cytometry is a combination of flow cytometry with mass spectrometry and provides measurement of over 40 parameters simultaneously at a single-cell resolution. This process allows for millions of cells to be assayed, enabling sufficient sampling to identify major cell subsets from the heterogeneous cell sample [44]. Unlike conventional FACS, which uses fluorophores as reporters causing spectral overlap, mass cytometry uses unique, stable, heavy metal isotopes of atomic weights different from those employed in mass spectrometry. This allows for greater number of parameters to be analyzed with little signal overlap between parameters. The instrumentation used for mass cytometry is called cytometry by time-of-flight [44]. Disadvantages of mass cytometry are that it cannot determine forward and side scatter, and hence cannot determine cell size and shape; the acquisition rate is slower than FACS; and cells must be fixed before analysis. Thus, mass cytometry can be used to analyze subpopulations of cells but cannot be used to enrich subsets of live cells based on these markers. A long-standing challenge in proteomic analysis is the transformation of high-resolution mass spectrometry as a cell populationeaveraging tool into a single-cell analyzer. Advancements have enabled the identification of 500e800 nonredundant protein groups from single cells using less than 0.2% of the total protein of the cell as starting material

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[45]. However, the main drawback is a low signal-to-noise ratio when considering the low levels of protein within a single cell [42].

Single-Cell Epigenetics Epigenetic modifications such as DNA methylation or histone modifications are functionally relevant changes to the genome that perturb gene expression without altering the DNA sequence. The best-studied epigenetic modification is DNA methylation, which consists of the addition of a methyl group to cytosine residues (5-methylcytosine). Usually, this alteration is inversely correlated with gene expression levels with implications for tumor biology, disease progression, resistance to standard drug treatments, and relapse. Attempts have been made to study epigenetics at the single-cell level. Single-cell genome-wide bisulfite sequencing (scBS-seq) has been used to assess the epigenetic heterogeneity of DNA. In this technique, the DNA of a cell is treated with bisulfite, which results in DNA fragmentation and the conversion of unmethylated cytosines to thymine [46]. Complementary strands of the fragmented DNA are synthesized using adaptor sequences and random oligonucleotides. This step is repeated several times to obtain enough tagged DNA and copy numbers of each fragment. A second adaptor is integrated, PCR amplification is performed, and cDNA libraries are generated, which are subjected to sequencing [46]. The number of CpGs obtained from the analyzed data depends on the depth of sequencing. Furthermore, to understand the complex relationship between DNA methylation and transcription in heterogeneous cell populations, scBS-seq of the genome and RNA seq of the transcriptome have been performed on the same cell. Such analyses have clinical implications in contexts such as in vitro fertilization, in which the number of cells for analysis is limited [47]. Advances in sequencing have cleared the way for many other methods to assess single-cell modifications. These include formaldehyde-assisted isolation of regulatory elements followed by sequencing, chromatin immunoprecipitation sequencing, DNase sequencing, Micrococcal nuclease followed by sequencing, and assay for transposaseaccessible chromatin sequencing. The general principle of these technologies is fragmentation of DNA and sequencing of the regions that have been bound by DNA-binding proteins [48]. This has allowed for improved profiling of DNA-binding proteins, histone modifications, and nucleosomes [49].

ANALYZING SINGLE-CELL DATA As detailed previously, single-cell technologies enable alterations in genes, gene expression profiles, and protein production within single cells to be determined [50]. However, single-cell sequencing technologies generate large amounts of data requiring computational infrastructure and expertise. For example, whole-genome sequencing of 100 individual cells at read lengths of 75 base pairs requires about 15 TB of storage space. Subsequently, the quality of these data needs to be assessed so that variations resulting from noise can be distinguished from true biological variations. The vetted data are then normalized and analyzed to reveal subpopulations of cells or information of clinical relevance. Thus, it is essential from the start to understand whether the biologic question requires a broad net to be cast to capture all information about individual cellular behavior, or whether a targeted approach is beneficial to reveal the most important cellular features affecting human health.

Reducing Noise in Single-Cell Data Noise in data generated from single-cell technologies can be classified into two major types: technical and biological [10,51]. During sequencing, technical variations can occur as a result of insufficient amounts of isolated RNA/ DNA, instability of the minute amounts of isolated RNA, problems in amplification of DNA/RNA, and bias during library preparation [51]. It has been estimated that over 80% of variation in expression patterns in single-cell RNA seq results from technical variations in measurement [51]. Biological noise can be caused by differences in the stages of the cell cycle, the epigenetic status of the cell, and the cellular microenvironment. It can also occur owing to the ineffective isolation of initial cells [26]. Biological confounders account for up to 18% of noise in single-cell RNA seq data and are especially problematic when analyzing rare subpopulations of cells [51]. To build accurate clinical models to guide diagnosis and treatment, it is important to minimize variability and noise in single-cell analysis techniques. This can be achieved by identifying and removing low-quality cells through

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two major steps: (1) quality control measures at each experimental step, and (2) quality control during data analysis [50]. Quality control at the experimental level can be performed at various stages by testing for the viability and size of the isolated cells, visually inspecting entrapment of cells within microfluidic chambers, testing the quality of DNA/ RNA, analyzing the size of the fragments generated, and quality testing the library preparation to ensure the data from each cell reaches a minimum threshold of usability. An interesting solution to detect doublets or multiplets is to mix cell populations from two different species. For example, when studying cardiomyocyte heterogeneity, cardiomyocytes from murine and human cells can be isolated and processed together. During data analysis, seemingly individual cells displaying mixed cDNA patterns can be discarded. The number of single cells analyzed can also be increased, resulting in a higher detection power [52]. However, this approach increases the amount of data generated when the same depth of sequencing is desired and requires greater computation capacity and resources. Another approach is to use nanoliter volumes of sample preparation to yield fewer false positives in gene expression studies compared with microliter formulations [10]. This method requires the use of automated liquid dispensers and robots. Quality control at the data analysis level can be performed by excluding cells with minimal numbers of reads or excluding cells with a high number of poor-quality reads. Furthermore, research groups have attempted to create models to remove biological confounding factors. One group succeeded in determining where cells lay within their paths of differentiation by creating a model that adjusted for cell cycleespecific changes in gene expression [53].

Normalizing Single-Cell Data Once clean data are identified, normalization can be performed to compare data from one cell with data from other cells. Normalization of data can be performed through a variety of techniques and can vary depending on the technology used [54]. This has proven to be challenging for all single-cell analysis methods. For bulk RT-PCR studies, housekeeping genes have been used to normalize gene expression against that of a gene with relatively constant expression across cells. However, at the single-cell level, there is significant variation between these housekeeping genes [55]. This has led to normalization with a combination of housekeeping genes or with genes that are best for the tissue being evaluated [56]. Some researchers even suggest using no normalization owing to these significant variations between cells [57]. As with single-cell RT-PCR analysis, there are also multiple methods for data normalization with RNA sequencing [54]. One method involves median normalization, in which the mean count for all cells of each gene is calculated and a size factor is created by determining the median of the fraction of each sample’s count over the mean across all samples. The size factor is then applied to all samples. Another method to normalize the data that accounts for technical artifacts is the use of spike-ins [54]. This involves introducing synthetic transcripts into each cell’s library at a known concentration. Knowledge of what the expected and observed counts are can then be used as a multiplicative factor to account for technical errors that occur during the process.

Mathematical Identification of Cellular Subpopulations Once the data have been vetted, they can be used in various ways. One specific purpose is to find important subpopulations with unique functions [50]. Single-cell technologies have demonstrated that cells that were once thought to be homogeneous display significant variations in their functional profiles and instead exist as subpopulations. However, perfectly equivalent expression profiles between any two cells are highly unlikely, which pushes the biostatistician to determine how best to group cellular subpopulations. A variety of techniques have been used to cluster cells within a population and reveal subpopulations. Two main categories of clustering include hierarchical and partition clustering. However, many variations exist within these groups [58]. Hierarchical clustering functions by combining or dividing groups and creating a hierarchical structure that demonstrates this order (Fig. 6.4). The two main methods within this technique include agglomerative nesting and divisive analysis [47]. Agglomerative nesting methods involve first arranging the data into a series of sets with one object in each set. For the purposes of single-cell analysis, an object would be considered the data profile from a single cell. A cost function is used to determine which of these sets is “cheapest” to combine. The two objects are then combined, removed from the list, and replaced with a combination of their components. The process is repeated until all items exist within a single group. Divisive analysis differs from agglomerative nesting in that it begins with one set

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FIGURE 6.4 Hierarchical clustering followed by k-means clustering can reveal novel subsets that are altered in disease. Single-cell transcriptional data can be represented by hierarchical clustering with each cell, represented as a column, and each gene, represented as a row in the heat map. There data can further be clustered using k-means clustering to reveal subsets that are altered in disease. Cells in cluster 2 (green box) show an increased frequency in the diseased state, whereas cells in cluster 4 show a decreased frequency in the diseased state (red box). These cells can then be specifically targeted to test for therapeutic outcomes.

containing all objects. The object with the greatest dissimilarity from the rest of the objects is separated from the group. The remaining objects are then evaluated to see whether they should be included in this separated group. The process is repeated until there are as many clusters as there are objects. Each separation is demonstrated within a dendrogram to show the level of separation among the various groups. Once an object is placed into a group, it cannot be reassigned from that group. The specific mathematical approach to clustering can be varied based on the cost function used within the algorithm. In contrast to hierarchical clustering, partition clustering involves a set number of clusters that are defined a priori [47]. This method involves placing objects into a cluster with the closest center. The goal is then to minimize dispersion within each cluster through iterative reallocation of the objects within clusters. Unlike hierarchical clustering, this method enables an object’s cluster assignment to be changed at any step. Two specific examples of partition clustering are k-means clustering and partitioning around medoids (PAM). The difference between these two methods is that the cluster center in k-means is the average of the objects, whereas in PAM the center is an actual object.

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A significant number of modifications have been made to both hierarchical and partition clustering. One example is demonstrated by the addition of fuzzy c-means clustering. The methods of partition clustering listed earlier involve placing an object into a single cluster, which is referred to as “hard assignment” [59]. Fuzzy c-means clustering uses a “soft assignment,” which allows placement of an object into multiple clusters. This becomes useful for objects that lie between two clusters. With partition clustering, a key requirement is to determine how many clusters one should create. Many methods have been presented to attempt to determine the optimal number of clusters, with none obviously superior to the others [59]. These include, but are not limited to, minimum message length criteria with a Gaussian mixture model, minimum description length, Bayes information criterion, Akaike information criterion, and gap statistics. Even with these objective measures, there is no perfect method to determine how many meaningful clusters exist within a given data set.

DETERMINING SUBPOPULATIONS Although cells are defined by their intracellular characteristics (transcription, translation, epigenetics, metabolism, etc), the only way to isolate live cells is to use markers on their surface. Many of these surface markers defined in the literature may have little to no mechanistic relationship to the intracellular function of the cell, but they are selected by educated guesses. For the first time, single-cell technologies allow the correlation of surface markers to the intracellular cellular machinery to ensure that the best surface markers are selected to determine cellular subpopulations. However, to identify the best signature for the cell, a broad net needs to be cast encompassing all surface markers. Single-cell RNA seq offers the widest resolution and can inform decision making. However, the large datasets generated by this technology contain high amounts of technical noise, masking, or amplifying cellular heterogeneity and make it difficult to define rare subpopulations accurately and precisely. This is like to finding a needle in a haystack [60]. From a technical perspective, it is unclear whether single-cell RNA seq will ever approach the accuracy and precision of single-cell qPCR. Such unreliability is not acceptable for prospectively isolating and functionally testing cell subpopulations and generating therapeutic products that will be used in humans. To overcome these shortcomings, a single-cell approach was developed and validated that leverages the comprehensiveness of single-cell RNA seq and the accuracy and precision of single-cell qPCR [60]. Specifically, this approach uses information from single-cell RNA seq, metaanalysis of publicly available microarray databases, and peer-reviewed literature to arrive at 96 genes with which to interrogate cells of interest (fibroblasts, hepatocytes, etc.). These genes include cell cycle, transcriptional, and cell-specific genes that appear to be sporadically expressed by the three screening analyses. The 96 genes are then transcriptionally analyzed in individual cells by qPCR on a reliable Fluidigm Biomark microfluidic chip. The single-cell transcriptional data are subjected to hierarchical and k-means clustering to validate expression of the 96 genes and accurately determine systematic variations within the cells. This is the first part of the analysis, which determines whether putatively homogeneous cells contain undiscovered subpopulations. At this point, if there are subpopulations (i.e., autonomous clusters), it is still not clear whether these subpopulations have any differential function. It is possible that the transcriptional clusters are merely a descriptive curiosity. To determine functional relevance requires isolation of the subpopulations and prospective functional analysis. Only cells with differing functionality are considered “true” autonomous cell subpopulations. To prove functional relevance, the subpopulations need to be isolated with surface markers using existing FACS technology. To do this, single cells are subjected to transcriptional analysis for a second time. Mathematically, the original 96 genes can be winnowed to eight or nine “cluster defining” genes, which leaves 80 or more open channels on each chip. These open channels are used to correlate the cluster defining genes blindly with all known surface markers capable of being used for FACS sorting. This results in a surface marker combination with the highest specificity and sensitivity for each subpopulation (receiver operating characteristic curve closest to 1). The surface marker combination identified is used to isolate the subpopulations of interest accurately and precisely using FACS, and the cells are functionally tested (Fig. 6.5). This technology can be used to speciate any cell type, isolate rare cells, and study cellular alterations in any disease state [33]. This single-cell approach has been used to (1) study cellular heterogeneity in various cell types including fibroblasts, mesenchymal stem cells, hematopoietic stem cells, neural stem cells (NSCs), and glioblastomas [7,33,34,36,61]; (2) determine subpopulation alterations in a wide range of diseases such as diabetes, aging, fibrosis, and cancer [34,35,62,63]; and (3) identify surface markers accurately for the prospective isolation of cells for therapeutic use [33].

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FIGURE 6.5 Development of a rational framework for the identification and isolation of functional cell subpopulations. Seemingly homogeneous cells within a population are subjected to single-cell transcriptional analysis to determine cell subsets based on differential gene expression across 96 genes. The 96 genes are selected from RNA sequencing, publicly available microarray databases, and peer-reviewed literature. The subsets are then probed with all 386 surface markers in parallel with the intracellular functional genes to determine the ideal surface markers that select for these clusters. Because the multiplex chip allows for the probing of only 96 genes, at least five chips are used and single-cell transcriptional data across these chips are displayed. Transcriptional data of the 386 surface markers is then blindly correlated with intracellular genes to determine the surface marker combination that identifies the cell with the highest sensitivity and specificity. Finally, antibodies for the selected surface markers are employed to fluorescence-activated cell sort (FACS) and enrich the functional subsets that can readily be used for investigation and therapy.

Development of Cell-Based Therapies Cell-based therapies have been developed based on ‘legacy’ surface markers derived from the literature and from historical data. Thus, when trials fail, it becomes difficult to determine the best way in which to proceed. Customized therapies require an in-depth knowledge of both disrupted cellular pathways in diseased tissue and cell surface marker information on cells that can bring about the best therapeutic effect. The effectiveness of the Needlestack platform was tested to identify and isolate single-cell subpopulations rationally for therapeutic use. Human subcutaneous adipose-derived stromal cells (ASCs) were selected as the cell therapy source [33]. These cells can easily be isolated from the subcutaneous tissue and were tested widely in preclinical trials based on their ability to produce growth factors and deposit extracellular cells. However, ASCs are a heterogeneous group of cells obtained from excluding hematopoietic cells and endothelial cells and enrich CD34expressing cells from the stromal vascular fraction. If these cells are to be used therapeutically, it is imperative to understand the healing effects of the various cell subsets within this population.

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Heterogeneous Populaon of Progenitor Cells

VASCUAR

BONE FAT

SORTING MECHANISM IN THE OR

CARTILAGE

MUSCLE

Re-implanng in the OR

Isolang Cells in the OR

FIGURE 6.6 Isolating the best cell for any given clinical application. Single-cell technologies can identify which cells in a heterogeneous population will provide the best outcome for a specific clinical application. Thus, heterogeneous cells such as adipose-derived stromal cell (ASCs) that are obtained from lipoaspirates in the clinic can be enriched for these highly potent cellular subsets in the operating room (OR) and delivered immediately to patients who require therapy.

Hence, ASCs were individually isolated and subjected to single-cell qPCR against 96 genes involved in tissue repair. The gene list was informed from the literature and publicly available microarray databases. The resulting data were subjected to k-means clustering to identify subpopulations. The subpopulation most favorable to wound healing was selected. This subpopulation was defined by 18 genes, which left open 78 channels to interrogate cell surface marker expression in a second set of experiments. Upon combining five chips, it allowed for an unbiased and blinded correlation with 386 surface markers. This blinded correlation (mathematically performed by linear discriminant analysis) identified two surface markers that precisely selected the subpopulation of interest [33]. Next, these two surface markers were used to enrich the ASC subpopulation by FACS, and the cells were delivered to diabetic wounds. Diabetic wounds displayed a normalization of the healing response with the enriched ASCs. Interestingly, the application of ASCs depleted of this important subset demonstrated no effect on diabetic healing, whereas application of the unsorted, heterogeneous ASCs improved but did not normalize diabetic healing. This was the first cell-based targeted therapy to normalize diabetic wound healing in a preclinical setting and could be extended to the treatment of any disease (Fig. 6.6). Importantly, this single-cell method for selecting the best cellular subpopulation exists during a time when US Food and Drug Administrationeapproved cell-based therapies such as Apligraf and Carticel have shown efficacy and have resulted in reduced overall health care costs for patients.

CLINICAL IMPLICATIONS OF CELLULAR HETEROGENEITY IN TISSUE REPAIR AND DISEASE Customized therapies need an in-depth analysis of impaired cellular pathways in disease and a granular understanding of cellular subpopulation changes that underlie disease. Although population-averaged assays cannot provide such resolution, the development of novel single-cell technologies provide great promise for targeted basic science and clinical discovery. This section summarizes advancements made using single-cell technologies in understanding the molecular and cellular changes that modulate diabetes, aging, wound healing, cancer, and fibrosis (Fig. 6.7).

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FIGURE 6.7 Understanding the functional relevance of cellular heterogeneity. The development of novel single-cell technologies offers great promise for targeted basic science and clinical discovery. These techniques allow for the comprehensive mapping of cells within various tissues in health and inform us about alterations in cellular subsets during aging or diseases such as diabetes, fibrosis, and cancer.

Cellular Heterogeneity in Diabetes Diabetes brings about cellular and molecular impairments in a wide variety of cell types including stem and progenitor cells leading to tissue dysfunction. In many cases, owing to hyperglycemic memory, these cellular perturbations do not normalize even after a return to normoglycemia, resulting in persistence of tissue dysfunction [64]. Single-cell interrogation of subcutaneous ASCs in type 1 and type 2 diabetes has demonstrated that the global dysfunction in ASCs is brought about by the selective depletion of discrete ASC subpopulations, impairing wound healing in diabetes [33,62]. Similarly, type 1 and type 2 diabetes bring about intrinsic defects within bone marrow progenitor cells through selective depletion in provasculogenic subpopulations. These defects are not correctable by restoring glucose homeostasis [65]. Within the bone microenvironment, single-cell RNA seq has also revealed intrinsic skeletal stem cell impairments caused by hyperglycemic changes within the stem cell niche [66]. Interestingly, single-cell analyses revealed differences within insulin-producing pancreatic b cells, a population long considered to be homogeneous. These studies indicated that adult b-cell subpopulations can differ in size, insulin production, insulin secretion, and precursor cell potential with relevance to an understanding of diabetes and implications for enhancing cell replacement therapies for treating diabetes [67,68].

Cellular Heterogeneity in Wound Healing Wound repair is an example of a highly heterogeneous tissue with several different cell types working in concert at distinct spatiotemporal stages to bring about healing [6]. Immediately after a wound is formed, neutrophils are recruited from the bone marrow as the first line of defense against bacteria. Classically, neutrophils have been considered a homogeneous population of terminally differentiated cells with a conserved function. Their limited proliferation ability, short life span, and low mRNA content (10e20 times lower than leukocytes) have made it difficult to tease apart the various subsets of neutrophils through population techniques [69]. However, single-cell technologies revealed both phenotypic and functional versatility in neutrophils that extend beyond their antimicrobial activity to their impact on disease and their ability to activate other cells such as macrophages [70e72]. With the evolution of single-cell technologies, the definition of a macrophage has evolved as a cell that engulfs and digests pathogens, particles, and dead cells. It is now accepted that tissue macrophages have the unique ability of plasticity, in which the cells modulate their activation state based on external cues such as the presence of infection, growth factors, and cytokines in their microenvironment [73]. The diversity within macrophages is seen at the phenotypic, genetic, and epigenetic levels, leading to subsets of macrophages that are proinflammatory, antiinflammatory, and provascular, or transitioning between these states. Furthermore, there are macrophages in the adult tissue that originate during embryonic development that are not derived from monocytes [74]. Thus, spatiotemporal

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factors within the wounded tissue microenvironment determine the presence of macrophage subpopulations, each potentially with a unique function. Neovascularization follows the inflammatory phase of repair in every tissue. During this phase, blood vessels are at various levels of maturity. Some vessels are intact and are maintaining blood fluidity, some are leaky and aiding the influx of inflammatory cells, and others are actively undergoing angiogenesis. During angiogenesis, endothelial cells are sprouting and proliferating, whereas pericytes within the basal lamina are activated to scaffold and provide structural integrity to the new vessels. Circulating progenitor cells from the bone marrow are also recruited to support new blood vessel formation. Appropriate synchronization of these cells is crucial for neovascularization and healing. However, population assays have been unsuccessful in definitively characterizing pericytes and circulating progenitor cells within the repairing wound. In wound healing, active proliferation and reciprocal interactions of fibroblasts with other cell types in the wound environment, such as keratinocytes, endothelial cells, adipocytes, inflammatory cells, and resident stem cells, are important. Although reduced extracellular matrix deposition by fibroblasts can contribute to nonhealing wounds, excessive extracellular matrix deposition can lead to hypertrophic scarring and fibrosis [75]. Single-cell analyses have led to the identification of various fibroblast subpopulations with distinct functions after injury [7,76]. These technologies have identified unique subsets of fibroblasts that are responsible for the scar response.

Cellular Heterogeneity in Fibrosis Tissue fibrosis is a common complication that underlies impaired tissue regeneration and tissue dysfunction in response to a variety of insults [6,75]. Fibrosis is a poorly understood process, but it is largely attributed to excessive extracellular matrix deposition by fibroblasts. However, fibroblasts are a heterogeneous population of cells [77]. To this end, single-cell technologies have been employed to interrogate fibroblast heterogeneity. It has been demonstrated that CD26þ fibroblasts constitute a distinct subpopulation of dermal fibroblasts, which is the primary cell type for excessive collagen deposition and scarring during wound healing associated with fibrosis [7]. Similarly, heterogeneity in fibroblasts mediating pathology such as pulmonary fibrosis and renal fibrosis have also been described in the literature [78,79]. Moreover, matrix stiffness cues from cross-linked collagen can induce other cells to turn into fibroblast-like cells, further contributing to fibroblast heterogeneity [80]. Macrophages are one of the cell types that deposit collagen in response to matrix stiffness. Thus, cellular heterogeneity in macrophage populations has formed the basis of many fibrosis studies. Traditionally, macrophages have been classified into proinflammatory M1 cells and antiinflammatory M2 cells [81,82]. Time-dependent shifts in relative proportions of M1/M2 macrophages underlie the reparative process as well as dysregulated excessive inflammation in the heart, kidney, and lungs. Comprehensive gene expression analysis of macrophages coupled with surface marker screening revealed that Ceacam1þ/Msr1þ/Ly6C/F4/80/Mac1þ cells, a distinct subpopulation of cells, is the chief contributor to bleomycin-induced fibrosis [83]. Similarly, subpopulations of macrophages that express CD34/CD68 have been found to be more prone to differentiate into myofibroblasts [84]. Single-cell transcriptional analysis has also been employed to study heterogeneity in fibrocytes, which are hematopoietic cells depositing collagen during tissue repair and fibrosis. This revealed the presence of a CD45þ/CD11bþ/ F480þ macrophage subpopulation forming fibrocytes during wound healing [35]. Further research into tissuespecific cellular heterogeneity will help develop therapeutic strategies to control fibrosis. Enhanced understanding of cell heterogeneity in fibrosis could lead to strategies for cellular reprogramming, with implications in wound healing therapeutics, tissue engineering, and regenerative medicine [77].

Cellular Heterogeneity in Aging Aging affects the regenerative capacity of most tissues. At the stem and progenitor cell level, these changes are attributed to both alterations of the intrinsic stem cell state and perturbations in the composition of stem cell subpopulations, which have been difficult to dissect in the past. Single-cell RNA seq has been used to differentiate between these cell-intrinsic and subpopulation differences in hematopoietic stem cells and progenitor cells. From a cell cycle perspective, these studies reveal a reduction in long-term HSCs (LT-HSCs) in the G1 phase with age. From a differentiation perspective, aged short-term-HSCs (ST-HSCs) resemble young LT-HSCs, which demonstrates that aged ST-HSCs fully self-renew and serve as the main source of hematopoietic maintenance in mice [85,86]. Single-cell qPCR has been used to evaluate the effects of aging on subcutaneous ASCs and influence their ability to support neovascularization in a wound healing setting. Although aging does not bring about changes in ASC

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number, viability, or proliferative capacity, the single-cell study demonstrated that aging depletes a vasculogenic subpopulation of ASCs, leading to impairments in wound healing [63]. Single-cell RNA seq has been used to characterize adult NSC and progenitor cell populations, to determine heterogeneity within these cells. This resulted in the finding that NSCs can be clustered into early, mid-, and late-stage subpopulations along the stages of activation and differentiation. This study provides an integrative understanding of the NSC lineage with clinical implications in aging [87].

Tumor Cell Heterogeneity and Drug Resistance Cells within tumors exhibit differential mutations and are derived from multiple lineages resulting in intratumor heterogeneity [52]. Tumor cell heterogeneity is a chief contributor to tumor invasion, metastasis, and resistance to drug therapy [52,88]. Two models have been proposed to drive this heterogeneity. The first, called the clonal evolution model, proposes that most neoplasms originate from a single cell, and the stepwise acquisition of mutations within this clone allows for the formation of more aggressive subclones, leading to tumor progression [89]. The second, theoretically opposing hypothesis, the cancer stem cell model, suggests that only a small subset of cells, called cancer stem cells, have tumorigenic potential, whereas their differentiated progeny have limited proliferation and tumorigenic potential. The elucidation of these two models had depended on xenograft limiting dilution assays and tumor markers from the literature [89]. Powerful single-cell technologies to discern this information had not yet been developed until now. These technologies might be able to suggest that the two tumor models are not mutually exclusive; cells within the tumor may display vast phenotypic plasticity and differentiated tumor cells undergoing dedifferentiation to acquire stemlike properties [90]. Single-cell RNA seq has been used to distinguish transcriptional diversity in genes regulating proliferation, immune response, and oncogenic signaling in cells isolated from human tumors [91]. In human glioblastomas, which is the most common and aggressive form of brain tumor, with an exceptionally low rate of survival, there has been a search for brain tumoreinitiating cells. One study using single-cell qPCR identified a distinct DDR1þ subset from murine and human glioblastomas as the primary driver of aggressive tumorigenicity in vivo [34]. Although glioblastoma is a well-established example, there is limited information from other brain tumors. Single-cell RNA seq has been applied to study other brain tumors, such as from patients with oligodendroglioma, and has revealed stem/ progenitor cell populations and unique differentiation programs within cells of these tumors [92]. In breast cancer, cancer stem cell subsets have been studied, with CD44high/CD24low cells representing a quiescent invasive mesenchymal state and ALDHþ cells representing a more proliferative epithelial state [93]. Advances in single-cell transcriptional profiling have taken these studies a step further and revealed novel targets such as calcium- and zinc-binding protein encoding gene (S100A9) to target breast cancer metastasis [94,95]. Similarly, singlecell sequencing of breast cancer cells has identified subpopulations that are resistant to chemotherapy. IGF1Rþ/ KDM5Aþ/ITGA6þ breast cancer cells, for example, have been found to be resistant to drugs such as Paclitaxel [96]. In addition, in situ single-cell analysis suggested that chemotherapy before human epidermal growth factor receptor 2etargeted therapy can increase treatment resistance as a result of changes in intratumor diversity [97]. In a similar vein, single-cell RNA seq of lung adenocarcinoma cells revealed a unique subset of KRAS G12Dþ/ high RS cells that are resistant to chemotherapy [98]. Single-cell analysis of human colon cancer samples identified cellular subsets with unique signatures such as KRT20negative/(CA1, MS4A12, CD177, SLC26A3)negative that correlate with worse clinical prognosis [99]. Similar approaches have been used to identify POU5F1þ cells as a subpopulation of invasive cells in melanoma [100]. Therefore, new single-cell analysis techniques enable the identification of specific tumor cell subpopulations that are resistant to drugs, mediate tumor metastasis, and are responsible for tumor relapse. These novel targets can be used to develop targeted cell-specific treatments for cancer.

CONCLUSIONS For a long time, our understanding of biology has been defined by measuring population averages of cellular behavior. However, in most cases, population averages do not result in accurate information, because cells within the same population exhibit heterogeneity. Rare but important cells such as stem and progenitor cells are almost unidentifiable in population-averaged studies. The evolution of single-cell technologies such as next-generation sequencing offers for the first time a comprehensive analysis of the entire genome and transcriptome of single cells,

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and the ability to discover rare and previously unidentified cells. These technologies reveal changes within individual cells without previous bias from the literature. However, obtaining single-cell data is a challenging process in terms of time, resources, technical know-how, and the reliability of analyzed data. Moreover, a major concern that accompanies single-cell sequencing technologies is the presence of technical and biological noise that needs to be differentiated from biologic variation and heterogeneity. Whether these challenges can be overcome is doubtful. Thus, it is important to determine in what cases comprehensiveness of biological information is necessary. In clinical situations, particularly those in which patients must be diagnosed or in which therapeutic products must be used in patients, unreliability of data is unacceptable. In such situations, the use of accurate and precise technologies such as single-cell qPCR and FACS remains the norm. Platforms such as the Needlestack combine RNA seq, single-cell qPCR, and FACS to provide reliable single-cell readouts and validate the cellular subpopulations discovered. Development of such accurate and precise single technologies will aid in the fundamental understanding of human biology and guide therapeutic development.

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[80] Dingal PC, et al. Fractal heterogeneity in minimal matrix models of scars modulates stiff-niche stem-cell responses via nuclear exit of a mechanorepressor. Nat Mater 2015;14:951e60. [81] Aurora AB, Olson EN. Immune modulation of stem cells and regeneration. Cell Stem Cell 2014;15:14e25. [82] Cao Q, Wang Y, Harris DC. Macrophage heterogeneity, phenotypes, and roles in renal fibrosis. Kidney Int Suppl (2011) 2014;4:16e9. [83] Satoh T, et al. Identification of an atypical monocyte and committed progenitor involved in fibrosis. Nature 2017;541:96e101. [84] Mesure L, De Visscher G, Vranken I, Lebacq A, Flameng W. Gene expression study of monocytes/macrophages during early foreign body reaction and identification of potential precursors of myofibroblasts. PLoS One 2010;5:e12949. [85] Busch K, et al. Fundamental properties of unperturbed haematopoiesis from stem cells in vivo. Nature 2015;518:542e6. [86] Kowalczyk MS, et al. Single-cell RNA-seq reveals changes in cell cycle and differentiation programs upon aging of hematopoietic stem cells. Genome Res 2015;25:1860e72. [87] Dulken BW, Leeman DS, Boutet SC, Hebestreit K, Brunet A. Single-cell transcriptomic analysis defines heterogeneity and transcriptional dynamics in the adult neural stem cell lineage. Cell Rep 2017;18:777e90. [88] Meacham CE, Morrison SJ. Tumour heterogeneity and cancer cell plasticity. Nature 2013;501:328e37. [89] Cabrera MC, Hollingsworth RE, Hurt EM. Cancer stem cell plasticity and tumor hierarchy. World J Stem Cells 2015;7:27e36. [90] Peterson JA. Single cell heterogeneity in breast cancer. In: Ceriani RL, editor. Immunological approaches to the diagnosis and therapy of breast cancer. US, Boston, MA: Springer; 1987. p. 41e53. [91] Patel AP, et al. Single-cell RNA-seq highlights intratumoral heterogeneity in primary glioblastoma. Science 2014;344:1396e401. [92] Tirosh I, et al. Single-cell RNA-seq supports a developmental hierarchy in human oligodendroglioma. Nature 2016;539:309e13. [93] Brooks MD, Burness ML, Wicha MS. Therapeutic implications of cellular heterogeneity and plasticity in breast cancer. Cell Stem Cell 2015;17: 260e71. [94] Powell AA, et al. Single cell profiling of circulating tumor cells: transcriptional heterogeneity and diversity from breast cancer cell lines. PLoS One 2012;7:e33788. [95] Hayes DF, Paoletti C. Circulating tumour cells: insights into tumour heterogeneity. J Intern Med 2013;274:137e43. [96] Lee MC, et al. Single-cell analyses of transcriptional heterogeneity during drug tolerance transition in cancer cells by RNA sequencing. Proc Natl Acad Sci USA 2014;111:E4726e35. [97] Janiszewska M, et al. In situ single-cell analysis identifies heterogeneity for PIK3CA mutation and HER2 amplification in HER2-positive breast cancer. Nat Genet 2015;47:1212e9. [98] Kim KT, et al. Single-cell mRNA sequencing identifies subclonal heterogeneity in anti-cancer drug responses of lung adenocarcinoma cells. Genome Biol 2015;16:127. [99] Dalerba P, et al. Single-cell dissection of transcriptional heterogeneity in human colon tumors. Nat Biotechnol 2011;29:1120e7. [100] Ennen M, et al. Single-cell gene expression signatures reveal melanoma cell heterogeneity. Oncogene 2015;34:3251e63.

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C H A P T E R

7 Embryonic Stem Cells: Derivation, Properties, and Challenges Irina Klimanskaya Astellas Institute for Regenerative Medicine, Marlboro, MA, United States

INTRODUCTION Embryonic stem cells (ESC) can be viewed as an immortal extension of short-lived pluripotent cells that exist in a preimplantation embryo. These pluripotent cells become all of the tissues of the body during embryo development, and cell lines created in vitro from these pluripotent cells retain important properties: self-renewal and the ability to differentiate into a variety of tissues of all three germ layers. An in vitro research model of these cells established in 1981 [1,2] immediately became indispensable for studying mechanisms of mammalian development, and when ESC were derived from human embryo in 1998 [3], regenerative medicine received a new and promising source of cells for tissue engineering. Cells of any type intended for a therapeutic application have to be functional in vivo, nontumorigenic, and free of pathogens, and it is highly desirable to have a reliable long-lasting source of these cells. When such cells are isolated from donor tissues, their potential for expansion is limited, which restricts the use of this source. As a desirable offthe-shelf product, pluripotent cells seem to be an excellent source of differentiated derivatives: Their ability to selfrenew allows for virtually limitless in vitro expansion, thus enabling large-scale manufacture, and they can be differentiated into a variety of derivatives that in turn can be purified and expanded.

DERIVATION OF EMBRYONIC STEM CELLS Mouse Embryonic Stem Cells In 1981, two independent research efforts resulted in the derivation of the first ESC lines from mouse embryos [1,2]. Both approaches used mitotically inactivated STO cells as feeders and based their assessment of cell morphology on the morphology of mouse embryonic carcinoma (EC) cells maintained in each laboratory. Evans and Kaufman considered critical factors for the success of derivation of pluripotent cell lines to be the window of embryonic development when pluripotent cells that would grow in culture existed in an embryo, the isolation of a sufficiently large number of such cells, and tissue culture conditions supportive of proliferation rather than differentiation of these cells. They first used an artificial delay in implantation induced by ovariectomy, which allowed late-stage embryos to remain free-floating in the uterus and to grow a large number of cells with no further development beyond primary ectoderm. After a 4- to 6-day delay, the blastocysts were recovered and cultured until egg cylinder-like structures formed that were isolated, trypsinized, and subcultured on mitotically inactivated STO cells, and selected for colonies resembling EC cells. Martin based her work on the premise that teratocarcinoma stem cells are derived from pluripotent cells of the periimplantation embryo and produce pluripotency-supporting factors. She used conditioned medium from an established EC cell line as a source of such factors and plated the inner cell masses (ICMs) of mouse blastocysts isolated by immunosurgery [4] on the STO feeder layer in such conditioned medium. Resulting colonies with morphology resembling EC cells were selected and passaged until a high-density culture was established that no Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00007-2

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longer depended on conditioned medium, because ESC were probably making pluripotency-supporting factors themselves. Pluripotent cell lines from both studies had a normal karyotype and differentiated into cells of three germ layers. These works laid the foundation for the huge field of study in cell and developmental biology on pluripotent stem cells and their differentiation. Mouse ESC are typically derived from mouse blastocysts with optional immunosurgery. They can be cultured indefinitely on mitotically inactivated STO cells or mouse embryonic fibroblast feeder layers in the presence of leukemia inhibitory factor (LIF) and fetal bovine serum (FBS), they express several markers of pluripotency, and they have high alkaline phosphatase activity [5]. Common pluripotency markers are transcriptional factor Oct-3/4, originally reported in the ICM of an early blastocyst [6,7], as well as Nanog [8], Sox-2, Rex-1, Dnmt3b, Lin-28 [9], and cell surface antigen SSEA-1, which is expressed on the surface of blastomeres of the eight cellestage embryo and ICM [9a]. ESC have become a common tool for generating transgene mouse models, because they easily aggregate with the ICM of a blastocyst when injected, or with blastomeres, making aggregation chimeras. ESC could also be generated from a single blastomere of a multicell-stage embryo [10] and appear to have the same properties including germ line transmission.

Human Embryonic Stem Cells In 1998, in a groundbreaking report from the group led by James Thomson at the University of Wisconsin [3], the derivation of human ESC was announced. The team used an approach similar to that for deriving mouse ESC: An ICM of a blastocyst, isolated by immunosurgery, was plated onto mouse embryonic fibroblast feeder cells in medium supplemented with LIF, basic fibroblast growth factor (bFGF), and knockout serum replacement (KSR) (Life Technologies) in place of FBS. Several human ESC (hESC) lines were generated and soon became available for use by other researchers worldwide. hESC have the capability of differentiating into any and all cell types of the human body; they immediately became recognized as a promising source of many cell types sought after by regenerative medicine. However, this discovery also evoked great controversy and led to heated discussions in the media because some considered the destruction of a preimplantation human embryo to be the same as killing a human. The use of federal funds for research involving hESC lines is allowed for National Institutes of Healtheregistered cell lines; there were 378 eligible hESC lines as of January, 2017 (https://grants.nih.gov/stem_cells/registry/ current.htm). Other examples of stem cell registries are the University of Massachusetts International Stem Cell Registry http://www.umassmed.edu/iscr/index.aspx and the European Human Embryonic Stem Cell Registry http://www.hescreg.eu/. ESC retain the fundamental property of ICM cells: the ability to give rise or develop into all tissues of the human body. The same markers of pluripotency are found in both ICM and hESC: transcription factors Oct-4, NANOG, and Rex-1; cell surface antigens SSEA-3, SSEA-4, TRA-1-60, and TRA-1-81; and high endogenous alkaline phosphatase activity [11,12]. They retain high telomerase activity and can continue to self-renew indefinitely. Both in vitro and in vivo, when injected into immune-deficient mice, hESC form teratomas, tumors that contain derivatives of all three germ layers, with the most commonly seen ones being bone, cartilage, neural rosettes, and epithelium of the airways and gut.

SOURCES OF HUMAN EMBRYONIC STEM CELLS Blastocyst The first and most commonly used source of hESC is a blastocyst. Many groups noticed that fully expanded and spontaneously hatching blastocysts are not the best candidates for the derivation of ESC: their outgrowth is very prone to differentiation, probably because of some intrinsic commitments already made by the cells of the ICM. In addition, the growth of trophoblast cells can potentially overgrow ICM-originated cells. The latter can be avoided by performing immunosurgery [4], in which a zona pellucidaefree blastocyst is incubated with antibodies that bind to the surface of trophectoderm, and then complement is added. This results in the lysis of trophoblast cells. The trophoblast-free ICM is plated for outgrowth. This procedure increases the efficiency of cell line derivation, but the success rate also depends on the quality and size of the ICM. Common impediments that may be encountered are spontaneous differentiation and apoptosis observed during the first week of ICM or blastocyst outgrowth. As mentioned earlier, Evans and Kaufman used delayed implantation of mouse embryos to achieve a larger ICM;

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Stoikovic and coauthors used a similar approach. They cultured late-stage blastocysts in the presence of Buffalo rat liver (BRL) celleconditioned medium through day 8, which allowed them to obtain larger ICM without its extensive further differentiation, and resulted in the successful derivation of hESC [13].

Morula As an alternative to whole blastocyst, which represents a relatively late stage in preimplantation embryonic development when the specification of the cell fate starts, earlier-stage embryos (morulae) were successfully used [14] to create hESC lines with an efficiency comparable to that for ICM- or whole blastocystederived hESC lines. These hESC appeared to have the same properties as blastocyst-derived hESC, such as a pattern of expression of pluripotency markers, differentiation to all three germ layers, immortality in culture, and a normal karyotype.

Growth-Arrested Embryo While many hESC cell lines have been created using leftover in vitro fertilization (IVF) embryos donated to research by couples undergoing infertility treatment, scientists kept working on alternative ways to make hESC without destroying the embryo. One of the first successfully executed approaches was the derivation of hESC lines from nonviable growth-arrested embryos [15,16]. This demonstrated that such embryos still have viable pluripotent cells that can be used to generate hESC lines.

Somatic Cell Nuclear Transfer A highly publicized approach to hESC derivation is based on somatic cell nuclear transfer (SCNT). During a micromanipulation procedure, an unfertilized egg is enucleated and the nucleus of a donor cell is introduced into the egg via a micropipette. The egg with the donor nucleus then develops into a blastocyst that can be used to isolate ESC. Such ESC would have the same genotype as the donor of the nucleus and can be used to generate an autologous cell type for tissue repair. Several research groups [17,17a,18] succeeded in creating hESC by SCNT using fetal, neonatal, and adult fibroblasts as donor cells for somatic nuclei; the success rate was 25% [18a], which was comparable to that for hESC derivation from a naturally fertilized blastocyst. However, the efficiency of blastocyst formation and hESC derivation seemed to be different for eggs from different donors and can even correlate with the hormonal stimulation protocol [18]. In addition, it was discovered [18a] that a major SCNT reprogramming barrier was associated with the severe methylation of lysine 9 in histone H3 in a human somatic cell genome. The introduction of KDM4A, an H3K9me3 demethylase, during SCNT significantly improved the development and blastocyst formation of SCNT embryos, even when the authors deliberately used eggs from donors whose eggs had previously failed to produce SCNT blastocysts. Although the possibility of creating donor-matched hESC lines by SCNT remains attractive, it takes years for the operator to develop the skills required to perform this procedure with high precision and minimal disturbance to the egg or donor nucleus for successful outcome, so this approach to deriving hESC remains in the realm of only a few groups in the world.

Parthenogenesis Another attractive possibility for generating pluripotent cells without destroying the embryos and overcome the problem of immune compatibility at the same time, is the generation of ESC and their derivatives from activated nonfertilized oocytes, or parthenotes [19], that would carry only maternal human leukocyte antigen (HLA) genes and thus allow to reduce the variability and number of lines required for immune match of the large number of patients. Due to genetic imprinting and the deficiencies of maternal and paternal haploid gene sets, their combined action is required for normal development. Parthenote mammalian embryos which do not have paternal genes are unable to develop to term; however, pluripotent cells produced from such parthenote human embryos seem to have phenotypes, behavior, and differentiation potential similar to those of hESC from blastocysts [20e23]. From an ethical viewpoint, parthenote hESC may be less controversial because no life is destroyed [23a]. A lot of progress has been made in generating human parthenote ESC and their derivatives [24e29], but data on the behavior of such derivatives in vivo are still limited. Several studies showed that a large percentage of parthenogenetic blastomeres was affected by an excessive number of centrioles, a high aneuploidy rate [30], and genetic and epigenetic

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instability [31]. More studies demonstrating the safety and efficacy of parthenote hESC derivatives are needed before this attractive source of HLA-matched cells for regenerative medicine can be fully used.

Single Blastomere To address the ethical controversy of hESC derivation from blastocysts, our group developed an approach based on single blastomere biopsy, a procedure commonly employed in preimplantation genetic diagnostics (PGD) in the course of IVF [32e34]. In this procedure, a hole is made in the zona pellucida of a morula stage embryo using acid Tyrode solution or with the help of a laser, and a single blastomere is extracted. The embryo continues to develop while the blastomere undergoes PGD tests, and a few days later, blastocysts that were “cleared” can be implanted. This procedure has been deemed safe and has resulted in hundreds of healthy babies being born. Using blastomere biopsy, we established several hESC lines from single blastomeres, whereas the parent embryos were allowed to develop to the blastocyst stage and then were cryopreserved [35]. Single blastomeres were first cocultured with the biopsied embryos and then were plated onto feeder cells in microdrops. Outgrowing colonies were treated in the same way as ICM outgrowths: with careful mechanical passaging followed by enzymatic dissociation when an appropriate number of colonies could be achieved. Established hESC lines had properties similar to ICMderived hESC: they stained positively for Oct-4, SSEA-3, SSEA-4, TRA-1-60, TRA-1-81, and alkaline phosphatase; they had a normal karyotype and differentiated into derivatives of all three germ layers both in teratoma assays in nonobese diabetic/severe combined immunodeficiency (NOD-SCID) mice and in vitro. The success rates for hESC derivation from blastocysts and blastomeres can be similar, around 25%e30%. This technique was subsequently used by several other groups to derive hESC lines with high efficiency and/or without destroying the embryo under different conditions including human feeder cells, feeder-free, and even from growth-arrested embryos [36e41]. Single blastomereederived hESC showed a transcriptional profile similar to “conventional” hESC [38], and several differentiated derivatives of single blastomereehESC appeared functional in vitro and in vivo [42,43,43a]. By avoiding the destruction of human embryos, this approach has offered a way to address ethical concerns regarding hESC derivation and provided a way to overcome a major impediment in developing hESC-based therapies.

HUMAN EMBRYONIC STEM CELL MAINTENANCE Similar to early mammalian development, when multiple cellecell and cellematrix interaction produce a variety of both inductive and permissive differentiation signals, ESC culture reproduces such signals with a certain degree of approximation: hESC readily differentiate into all three lineages, but the efficiency of their commitment to each lineage often depends on specific culture conditions, because this artificial system does not provide the same finely tuned orchestration as that which happens during in vivo pattern formation. The strong inclination of hESC to differentiate in culture makes it challenging to maintain hESC in a self-renewing state and is associated with the limitations of a two-dimensional cell culture environment and media components used as substitutes for the short-lived microenvironment of a preimplantation embryo. Meticulous attention to several key factors is required for the successful maintenance of hESC in a pluripotent state for multiple passages: • Microenvironment: defined combination of media and matrix, with or without feeder cells • Maintenance: frequent passaging using gentle and efficient cellematrix disruption • Morphology observation performed daily to assess the culture and avoid subculturing cells with signs of differentiation

Microenvironment hESC are usually cultured on either a monolayer of mitotically inactivated feeder cells or a defined extracellular matrix (for example, Matrigel, vitronectin, laminin-521) in defined media that allow them to be propagated in culture virtually indefinitely, unlike cells of the ICM, which exist only transiently. Like the ICM of a developing blastocyst, which is programmed to differentiate once the implantation starts, pluripotent ESC in culture readily differentiate when the microenvironment changes. Even the most carefully selected and tested combinations of media and matrix will not allow hESC to keep growing in the same dish and remain pluripotent indefinitely, because when colonies

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reach a certain size and become overcrowded, they form a second layer of cells, and then the cells in the colony begin to lose pluripotent marker expression and differentiate. Timely passaging prevents this loss of pluripotency because it releases single cells and small cell clumps, so any matrixecell interactions that were formed during the first days after previous passage and begin to send differentiation signals to hESC are disrupted, and the cells can continue to self-renew. The first hESC lines were derived and cultured under conditions similar to what is used for derivation and maintenance of mouse ESC: mitotically inactivated mouse embryonic fibroblasts (MEF) plated on gelatin, in a medium based on KSR, a proprietary serum substitute formulation. This medium was supplemented with LIF and bFGF [3,11,33,34]. Variations of this system include adding Plasmanate [33] or using a 1:1 mix of Knockout- Dulbecco’s Modified Eagle Medium and F12 medium. However, it was reported that human hESC do not have an active signal transducer and activator of transcription 3 (STAT3) pathway and thus are LIF-independent [44,45], and many researchers stopped using it in the culture medium. Indeed, it appeared that using bFGF alone is sufficient to support the pluripotency of hESC over multiple passages. Other studies have shown the importance of LIF in maintaining hESC in what is called a naive state (discussed subsequently), so it is probably too early to make a conclusion about the need of this factor to maintain pluripotency.

Maintenance It is commonly observed that cultures of hESC contain differentiating cells that, when there are only few of them, usually do not interfere with the successful maintenance of hESC in a pluripotent state or with their differentiation toward a desired derivative. However, if the colonies are allowed to overgrow, soon afterward they become multilayered or begin to touch each other, and spontaneous differentiation usually follows within hours. Although it is possible to rescue even extensively differentiated cultures (for instance, by carefully selecting undifferentiated colony pieces, or “mechanical picking,” it is more practical to prevent the loss of pluripotency by timely passaging based on observing the colony morphology and confluency and by using high-quality reagents and good cell handling practices. For an hESC subculture, there is a wide variety of commercially available dissociating agents. In the past, collagenase used by Thomson and colleagues [3] for hESC derivation and passaging was an enzyme of choice for many laboratories. It allowed the cells to be passaged as clumps rather than as single cells, but it required meticulous attention to colony morphology because when colonies of larger size are harvested, they are more prone to spontaneous differentiation. Trypsin is another popular enzyme, but because of its rapid action, it demands careful techniques during cell harvest to avoid cell damage and death. Mechanical colony dispersion and hand picking allows the selection of colonies of “proper” morphology with minimal stress to the cells and is the most commonly used method to derive new lines (when the outgrowth is small and the removal of differentiating cells is needed to prevent further differentiation of the remaining pluripotent cells). However, this method is operator-biased and demands sufficient experience to avoid selecting for aneuploid cells that may have a growth advantage and thus are the first to form good-sized and “good-looking” colonies. On the other hand, a skilled operator may be able to rescue an aneuploid culture by carefully selecting and dispersing colonies of the right morphology. Other available dissociation methods that allow hESC colonies to gently break into small cell clumps are Accutase, TrypLE (proprietary mix of enzymes, Life Technologies), and “Dissociation Buffer” (containing chelating agents, Life Technologies). hESC can maintain normal karyotype over multiple passages, but they are prone to aneuploidy. There are not enough data to identify reliably which factors cause aneuploidy, although a study showed an association of aneuploidy with high-density culture [46]. The authors demonstrated that 33% of hESC became genetically abnormal after only 5 days of high-density culture. It seemed that lactic acid and acidification of the medium were the main reason for such abnormalities, but interestingly, laminin 521 used as matrix counteracted this effect. Frequent change of medium also prevented such aberrations. Some of most commonly seen chromosome abnormalities, such as trisomy in chromosome 12 or 17 [47], are known to result in hESC survival or growth advantage, so even an initial nonclonal aberration can quickly become prevalent. Other common abnormalities in hESC karyotype are seen with chromosomes 1, 14, and 20 [48]. Dissociating agents can contribute to the quick spread of abnormal cells throughout the population [48a]. This tendency of hESC to undergo clonal aneuploidy reinforces the importance of making frequent karyotyping a part of the routine maintenance of hESC. G-banding with the examination of a minimum of 20 cells complemented by fluorescence in situ hybridization (FISH) with probes for chromosomes 1, 4, 12, 17, and 20 can be considered appropriate karyotyping methods.

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When choosing the most suitable combination of matrix, media, and passaging methods, it is important to remember that in addition to supporting pluripotency, a stable karyotype and reproducibility, the culture system needs to ensure the acceptable efficiency of the differentiation of hESC into the desired derivative under specific protocols. For instance, the formation of embryoid bodies (EB) (plating hESC into low-attachment cell culture plates, where they can aggregate and form cell clumps, differentiating into three germ layers) is frequently used to simulate early differentiation events in mammalian development. However, when hESC are dissociated into single cells, the formation of EB may be impeded by low cell survival, and the yields of EBs and differentiated cells can be much lower [48b].

Morphology The morphology of individual hESC and of colonies was mentioned earlier with regard to the maintenance routine and the choice of culture conditions. hESC have been known to form colonies with sharp “shiny” borders when they are cultured on feeder cells, and cells in such colonies are small with a high ratio of nuclei to cytoplasts and visible nucleoli. When cultured feeder-free, hESC look larger and more spread-out [55]. The borders of the colonies may become sharp after only several days in culture, when the cells become almost “overgrown.” Under these conditions, hESC inside the colonies look relatively large and flat for the first few days and may even resemble early stages of differentiation, which can be confusing to an inexperienced eye; however, after several days, the morphology of such colonies becomes more similar to typical hESC as cells become “packed,” so they become smaller in diameter. However, larger and flatter cells can also indicate the beginning of differentiation, and daily follow-up of the same cell culture dish is important to better understand the nuances of cell morphology, to tell apart which deviations from the “ideal” morphology are associated with cell adaptation after passaging and which are early signs of differentiation.

Evolution of Human Embryonic Stem Cell Derivation and Culture Methods There are several commercially available media with a proprietary blend of growth factors and nutrients that support the maintenance of hESC growth and pluripotency. Some examples are NutriStem (Biological Industries, Israel), TeSR media (Stem Cell Technologies), and E8 (Thermo Fisher Scientific). These media usually come with a detailed protocol prompting the researcher to use a certain matrixemedia combination. Commonly used extracellular matrices are Matrigel, an extracellular matrix produced by Engelbreth-Holm-Swarm mouse sarcoma cells (BD Biosciences), laminin 521 or 511 (BioLamina Sweden), CELLstart, a human placenta-derived extracellular matrix (Life Technologies), and vitronectin. If a special matrixemedia combination is desired, the conditions have to be carefully tested: hESC need to be subcultured for several passages under new conditions before it is known whether a certain combination works for the support of growth and pluripotency and ensures karyotype stability. MEF as feeder cells have been a reference standard for ESC culture for years. Since the beginning of hESC research, there has been great interest in using their derivatives clinically, but coculture of live human cells with live animal cells results in a xenogeneic product and requires more extensive testing for animal viruses. Thus is has been highly desirable to be able to derive and propagate hESC without feeder cells of animal origin or even make them feeder-free. Although xenogeneic cell products are allowed for transplantation in human patients, there are more stringent regulations. For instance, the patient’s blood samples needs to be archived for a prolonged time, which adds to the costs and logistics of clinical applications. The first reports on using feeder cells of human origin demonstrated the feasibility of using cells from different donors and/or different tissues, although some cell types seemed to be more effective in supporting hESC derivation and growth [39,49e53]. Another step in the transition from a xenogeneic to a nonxenogeneic cell product was the feeder-free culture system [54], in which the MEF-conditioned cell culture medium allowed hESC to propagate on Matrigel or laminin and retained their pluripotency and normal karyotype. However, this work was done using cell lines established with live mouse feeder cells, so these feeder-free cultured hESC still classified as a xenogeneic product. We attempted to derive new cell lines from blastocysts that were completely feeder-free. We postulated that an important factor in establishing the outgrowth of pluripotent cells was interaction with an extracellular matrix that was produced and properly organized by feeder cells. MEF were cultured for several days to allow for this matrix to be produced and organized, and then sodium deoxycholate was used to lyse the MEF whereas the preassembled extracellular matrix was left intact. The blastocysts or ICMs were plated onto this matrix in KSR-based derivation medium supplemented with LIF and bFGF, and the outgrowing hESC colonies were dispersed using mechanical passaging or

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trypsineEDTA. hESC cell lines that showed normal karyotype, maintained pluripotency marker expression over multiple passages and differentiated into derivatives of all three germ layers in vivo and in vitro was established as a result of this work [55]. Although derivation of hESC is more challenging than propagation, more feeder-free hESC lines were established using novel defined media such as NutriStem [40].

NAIVE EMBRYONIC STEM CELLS hESC and mouse ESC have many similarities, yet they are different. Mouse ESC have been shown to exist in two states: “naive” and “primed” [56e59]. Not only do naive pluripotent cells express markers of pluripotency such as Oct4, NANOG, Sox2, both X-chromosomes are active. Primed ESC coming from the epiblast (EpiSC) retain the expression of these markers, but only one X chromosome is active [60]. hESC resemble EpiSC rather than naive: the colonies they form are flat, not dome-like as is typical for naive cells, they are polarized [60a]; their metabolism is mostly glycolic whereas primed cells rely on oxidative phosphorylation [61,62]; only one X-chromosome is active; and they are sensitive to single-cell dissociation [57,63]. Naive cells survive single-cell dissociation much better, which allows for greater capacity for expansion (which could be highly desirable for large-scale manufacturing), their doubling time is much shorter, and they have been shown to differentiate more efficiently both in vivo and in vitro [64,65], a highly sought-after property when differentiation is aimed at making derivatives for regenerative medicine. Mouse ESC are usually derived and exist in a naive state, and it takes special effort and culture conditions to isolate EpiSC from an epiblast. On the other hand, most hESC lines were derived and propagated in a primed state, until naive, or “ground state” pluripotent hESC were derived [66,67]. It has been shown that naive hESC can be derived de novo, or existing hESC lines can be converted into the naive state. In both cases, signaling pathways need to be activated or inhibited using a cocktail of bioactive substances and small molecules. Some examples are MEK inhibitor PD0325901, GSK inhibitor CHIR99021, STAT3 inhibitor NSC74859, ROCK inhibitor Y27632, LIF, FGF2, and TGFb1. Interestingly, all such cocktails include LIF, which was once considered to be essential for hESC maintenance in the pluripotent state but then was abandoned after it was shown that hESC do not depend on it for pluripotency. It remains unclear whether the transition from naive to primed hESC happens during the first days, if not hours, of derivation from a naive ICM or whether such primed cells exist in the embryo. RNA sequencing analysis of single cells of human blastocysts showed that at least three types of cells can be identified, and that during derivation of hESC there are changes in gene expression as the cells adapt to culture conditions [68].

HUMAN EMBRYONIC STEM CELL DIFFERENTIATION AND MANUFACTURING FOR CLINICAL APPLICATION Although hESC provide exciting research opportunities for studies related to the mechanisms of development, their biggest promise is in their differentiation potential. There are unmet medical needs for a variety of cell types, including but not limited to pancreatic b cells, hepatocytes, dopamine neurons, oligodendrocytes, retinal cells, cardiomyocytes, vascular cells, cartilage, and bone: derivatives of all three germ layers. Some of these cell types are frequently seen in hESC cultures left to differentiate spontaneously or in EB, and some require specialized derivation protocols for efficient yields. For instance, the neural path seems to be a default choice of ESC in the absence of other differentiating cues [69], and in spontaneously differentiating hESC cultures [69e71] the presence of various neural lineage cells is common. Perhaps this natural tendency of hESC to form derivatives of neural lineage endorsed derivation and studies of such derivatives, and oligodendrocyte progenitors were the first hESC-derived cells to enter clinical trials for spinal cord injury. More clinical trials followed using another hESC-derived progeny of the neural lineage: retinal pigment epithelium (RPE) cells to treat macular degeneration and Stargardt’s disease in the United States and European Union [72,73]. RPE has a unique function: it provides support for the photoreceptor by delivering nutrients and removing shed outer segments. Its morphology is also unique which allow to easily detect these cells as pigmented cobblestone “islets” among various types of differentiating hESC. Unlike many terminally differentiated cell types, RPE retains its proliferation potential, so these pigmented epithelial cells can be relatively easily isolated with high purity and also be efficiently scaled up (Klimanskaya et al., 2004). Such hESC-derived RPE cells can fully differentiate and mature

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after transplantation in animal models, fully integrate into the host’s RPE layer, retain RPE morphology and molecular markers, and provide photoreceptor rescue [43a,73,74]. For this unique type of cells, a relatively small scale hESC culture can result in efficient production of the final product, and considering the small size of the macula and low cell numbers required for injection (in the range of several hundred thousand cells) [72,73], a relatively small-scale cell culture is needed to produce therapeutic doses. However, for other cell types considered for therapeutic applications that cannot undergo multiple population doublings while maintaining potency, or when large numbers of cells are required, various challenges can arise, such as using alternative cell culture systems including suspension culture, microcarrier- or microfluidics-based bioreactors at the hESC stage of the process, or isolation of progenitor type cells that can be expanded further. Whereas there may be a variety of approaches to manufacturing different hESC-originated cell types, the same principles can be applied to producing all hESC derivatives to ensure their safety and efficacy. The US Food and Drug Administration issued a document on tissue donor regulations (Guidance for Industry: Eligibility Determination for Donors of Human Cells, Tissues, and Cellular and Tissue-Based Products (HCT/Ps), https://www.fda.gov/ downloads/biologicsbloodvaccines/guidancecomplianceregulatoryinformation/guidances/tissue/ucm091345. pdf), and these regulations apply to embryo donors. All raw materials used in manufacturing need to meet the safety criteria and be fully characterized. Standard tests performed on manufactured cells include sterility, mycoplasma, common and latent viruses, and endotoxins. Of utmost importance is the safety of the cell product. First, there should be solid proof that the manufacturing process does not result in the presence of residual pluripotent cells in the final product, which can lead to the formation of teratoma, and the assays used to detect such pluripotent cells have to have a high level of sensitivity. Animal studies may be required to confirm that transplantation of the hESC derivative does not lead to tumor formation in immune-suppressed animals. Cell growth in soft agar can be used as an additional in vitro tumorigenicity test, and an assay should be in place to confirm the desirable purity of the derivative. Cells intended for transplantation have to have a normal karyotype confirmed by a rigorous test. The potency of the final product has to be confirmed using physiologically relevant assays: For instance, it could be phagocytosis for RPE, glucose-responsive insulin production for pancreatic b-cells, and bone and cartilage formation for mesenchymal stem cells. For each cell type, the most relevant and technically feasible in vitro assay should be chosen.

CONCLUSIONS ESC help to recapitulate early events in mammalian development and are an excellent system to study selfrenewal and differentiation. Since hESC were derived, numerous studies have demonstrated that they could be a promising source of various cell types for regenerative medicine. There are several alternative sources of pluripotent stem cells: the ICM of a blastocyst, morula, single blastomere, parthenote embryos, embryos generated via SCNT, growth-arrested embryos, and induced pluripotent cells. The same principles established through many years of research using ESC can be applied to the culture and differentiation of all pluripotent cell lines of different origins and sources to ensure their robust propagation in self-renewal state and efficient differentiation. Since the first derivation of hESC the culture methods have been become increasingly more robust, and there is now a variety of approaches to propagating pluripotent stem cells including hESC. Human feeder cells or feeder-free conditions can be used along with several defined mediaeextracellular matrix combinations and a variety of dissociating agents, including xenogeneic-free reagents. At the same time, many challenges are associated with our incomplete knowledge of the fine mechanisms of selfrenewal and differentiation. Further studies should help us to better understand which components of the cell culture system and steps in the maintenance routine are most critical in supporting hESC self-renewal, a stable karyotype, and differentiation efficiency to advance the use of hESC derivatives in regenerative medicine.

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[62] Ware CB, Nelson AM, Mecham B, et al. Derivation of naive human embryonic stem cells. Proc Natl Acad Sci USA 2014;111:4484e9. [63] Lewandowski J, Kurpisz M. Techniques of human embryonic stem cell and induced pluripotent stem cell derivation. Arch Immunol Ther Exp (Warsz) October 2016;64(5):349e70. [64] Dodsworth BT, Flynn R, Cowley SA. The current state of naı¨ve human pluripotency. Stem Cell November 2015;33(11):3181e6. https:// doi.org/10.1002/stem.2085. [65] Duggal G, Warrier S, Ghimire S, et al. Alternative routes to induce naive pluripotency in human embryonic stem cells. Stem Cells 2015. [66] Gafni O, Weinberger L, Mansour AA, et al. Derivation of novel human ground state naive pluripotent stem cells. Nature 2013;504:282e6. [67] Hanna J, Cheng AW, Saha K, Kim J, Lengner CJ, Soldner F, Cassady JP, Muffat J, Carey BW, Jaenisch R. Human embryonic stem cells with biological and epigenetic characteristics similar to those of mouse ESCs. Proc Natl Acad Sci USA May 18, 2010;107(20):9222e7. [68] Yan L, Yang M, Guo H, Yang L, Wu J, Li R, Liu P, Lian Y, Zheng X, Yan J, Huang J, Li M, Wu X, Wen L, Lao K, Li R, Qiao J, Tang F. Single-cell RNA-Seq profiling of human preimplantation embryos and embryonic stem cells. Nat Struct Mol Biol 2013;20:1131e9. [69] Ying QL, Stavridis M, Griffiths D, Li M, Smith A. Conversion of embryonic stem cells into neuroectodermal precursors in adherent monoculture. Nat Biotechnol 2003:183e6. [70] Munoz-Sanjuan I, Brivanlou AH. Neural induction, the default model and embryonic stem cells. Nat Rev Neurosci 2002;3:271e80. [71] Smukler SR, Runciman SB, Xu S, van der Kooy D. Embryonic stem cells assume a primitive neural stem cell fate in the absence of extrinsic influences. J Cell Biol 2006;172:79e90. [72] Schwartz SD, Regillo CD, Lam BL, Eliott D, Rosenfeld PJ, Gregori NZ, Hubschman JP, Davis JL, Heilwell G, Spirn M, Maguire J, Gay R, Bateman J, Ostrick RM, Morris D, Vincent M, Anglade E, Del Priore LV, Lanza R. Human embryonic stem cell-derived retinal pigment epithelium in patients with age-related macular degeneration and Stargardt’s macular dystrophy: follow-up of two open-label phase 1/2 studies. Lancet February 7, 2015;385(9967):509e16. [73] Schwartz SD, Hubschman JP, Heilwell G, Franco-Cardenas V, Pan CK, Ostrick RM, Mickunas E, Gay R, Klimanskaya I, Lanza R. Embryonic stem cell trials for macular degeneration: a preliminary report. Lancet February 25, 2012;379(9817):713e20. https://doi.org/10.1016/S01406736(12)60028-2. [74] Lund RD, Wang S, Klimanskaya I, Holmes T, Ramos-Kelsey R, Lu B, Girman S, Bischoff N, Sauve´ Y, Lanza R. Human embryonic stem cellderived cells rescue visual function in dystrophic RCS rats. Cloning Stem Cells Fall 2006;8(3):189e99.

Further Reading Boiani M, Scho¨ler HR. Regulatory networks in embryo-derived pluripotent stem cells. Nat Rev Mol Cell Biol November 2005;6(11):872e84. Brevini TA, Gandolfi F. Parthenotes as a source of embryonic stem cells. Cell Prolif February 2008;41(Suppl. 1):20e30. Kato R, Matsumoto M, Sasaki H, Joto R, Okada M, Ikeda Y, Kanie K, Suga M, Kinehara M, Yanagihara K, Liu Y, Uchio-Yamada K, Fukuda T, Kii H, Uozumi T, Honda H, Kiyota Y, Furue MK. Parametric analysis of colony morphology of non-labelled live human pluripotent stem cells for cell quality control. Sci Rep September 26, 2016;6:34009. Pan GJ, Chang ZY, Scho¨ler HR, Pei D. Stem cell pluripotency and transcription factor Oct4. Cell Res December 2002;12(5e6):321e9. Robertson EJ. Teratocarcinomas and embryonic stem cellsA practical approach. Oxford: IRL Press; 1987. p. 254. Stojkovic P, Lako M, Stewart R, Przyborski S, Armstrong L, Evans J, Murdoch A, Strachan T, Stojkovic M. An autogeneic feeder cell system that efficiently supports growth of undifferentiated human embryonic stem cells. Stem Cells March 2005;23(3):306e14. Turovets N, Fair J, West R, Ostrowska A, Semechkin R, Janus J, Cui L, Agapov V, Turovets I, Semechkin A, Csete M, Agapova L. Derivation of highpurity definitive endoderm from human parthenogenetic stem cells using an in vitro analog of the primitive streak. Cell Transplant 2012;21(1): 217e34. Wu G, Scho¨ler HR. Role of Oct4 in the early embryo development. Cell Regen (Lond) April 29, 2014;3(1):7. Stojkovic P, Lako M, Przyborski S, Stewart R, Armstrong L, Evans J, Zhang X, Stojkovic M. Human-serum matrix supports undifferentiated growth of human embryonic stem cells. Stem Cells August 2005;23(7):895e902.

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C H A P T E R

8 Alternative Sources of Human Embryonic Stem Cells Svetlana Gavrilov, Virginia E. Papaioannou, Donald W. Landry College of Physicians and Surgeons of Columbia University, New York, NY, United States

INTRODUCTION Human embryonic stem cells (hESC) are conventionally derived from viable preimplantation embryos produced by embryonic stem (IVF) [1]. The derivation of hESC is considered ethically controversial because of the typical destruction of the embryo during this process [2e5]. A human embryo constitutes an object of moral concern [6] owing to its identity as a human at the embryonic stage of development. In biological terms, a human embryo has a distinct, unique, and unambiguous status as a result of this identity. However, the political and moral status of human embryos are in a state of flux. Whereas there is universal opposition to reproductive cloning of humans by any method, there is diversity in opinion regarding the use of human embryos to derive hESC and, subsequently, potential therapies derived from them [7,8]. Ethical and cultural imperatives to respect human dignity from the moment of fertilization conflict with a utilitarian desire to relieve human suffering, even when this comes at the expense of embryonic human life. These conflicting perspectives have fueled an intense debate and have influenced legislative regulation of stem cell research in the United States and internationally [2e5,9,10]. US stem cell research policy was regulated on the federal level by the Dickey Amendment and President Obama’s Executive Order 13,505 and by individual state laws (Box 8.1) [10]. The use of federal funding to derive new hESC that would entail destroying human embryos is forbidden. Also, in many European countries (Austria, Germany, Ireland, Italy, Lithuania, Norway, Poland, and Slovakia), the derivation of hESC from surplus embryos is prohibited [9]. Because stem cell biology is at the forefront of research, legislative acts change rapidly. (For up-to-date legislative regulation of human embryonic stem (ES) cell research, refer to links provided in Box 8.2) [9,10]. Another consideration is the constant demand to derive new human embryonic stem (ES) lines for both basic and clinical applications owing to the loss of genetic and epigenetic stability arising during hESC culture and manipulation [11e14]. Many available hESC lines had been exposed to animal material during derivation or culture [2,15]. It is acceptable to expose hESC lines to products of human origin, but it remains the ultimate goal to pursue hESC derivation under stringent xenogeneic-free conditions for eventual clinical use [2,15]. The debate on embryo-destructive derivation of ES cells often focuses on the moral sensibilities of investigators and their desires for research unfettered by ethical considerations. However, the goal of hESC research is to find therapies that would ease human pain or debilitation caused by illness or injury [2,16,17]. In the latter context, the sensibilities of many millions of the populace (the intended beneficiaries of this work) should be instructive. As a result, a variety of different derivation strategies have been proposed (Fig. 8.1) to avoid using an embryo as a source of human stem cells (detailed information can be found in appropriate chapters of this book or elsewhere) [2,3]. In this chapter, we will discuss two alternative approaches to yielding genetically unmodified hESC that do not interfere with the developmental potential of human embryos: single blastomere biopsy (SBB) and organismically dead embryos (Fig. 8.1) [2].

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BOX 8.1

BRIEF OVERVIEW OF US FEDERAL STEM CELL POLICY The US policy on stem cell research is shaped by the following legislative act and executive order: • The “Dickey amendment,” a rider issued in 1996 that framed all subsequent political discussions regarding human embryonic stem cells (hESC) research. The amendment stated that no federal funding may be employed for (1) the creation of a human embryo or embryos for research purposes or (2) research in which a human embryo or embryos are destroyed, discarded, or knowingly subjected to risk of injury or death (beyond that permitted for fetuses in utero under the Public Health Service Act). • Executive Order (EO) 13,505, which removed barriers to responsible scientific research involving human

stem cells. This EO was issued by President Obama on March 9, 2009 and stated that the Secretary of Health and Human Services, through the director of the National Institutes of Health, may support and conduct responsible, scientifically worthy human stem cell research, including human stem cell research, to the extent permitted by law. In addition, this EO revoked two items issued by President George W. Bush: (1) a presidential statement that permitted work only on hESC lines generated before August 9, 2001, and (2) EO 13,435, which favored all research on stem cells without harming a human embryo.

BOX 8.2

USEFUL LINKS AND RESOURCES FOR INFORMATION ON CURRENT LEGISLATION IN THE UNITED STATES AND INTERNATIONALLY National Institutes of Health (NIH) Stem Cell Information webpage: contains relevant information on current US stem cell policy; NIH Stem Cell Registry with a list of eligible lines for NIH funding. The page also contains public comments on draft NIH human stem cell guidelines that supplement Executive Order 13,505. http://stemcells.nih.gov/index.asp.

International Society for Stem Cell Research webpage: contains comprehensive information on international legislation on human embryonic stem cell research; periodically updated. http://www.isscr.org/.

Single Blastomere Biopsy SBB for the purpose of deriving ES cells was developed by Lanza and colleagues [18e21]. HESC are created from a single blastomere that is removed from the embryo [20,21] by employing a technique that was originally developed for preimplantation genetic diagnosis (PGD) [2,22e24]. This procedure bypasses the ethical issue of embryo destruction, because biopsied embryos continue to develop and reach the blastocyst stage and beyond, as demonstrated by more than a decade of experience with PGD [2,24]. SBB of both murine and human eightcell stage embryos has been used successfully as a source of material to derive ES cell lines (Fig. 8.1) [2,18e21]. The risk associated with embryo biopsy [25] is accepted by patients as part of the PGD procedure, but it would be considered unjustified in a research setting in the absence of a clinical indication [2]. In addition, US regulations forbid research on an embryo that imposes greater than minimal risk, unless the research is for the direct benefit of the fetus (Box 8.1) [26]. To date, none of the hESC lines derived by SBB have been approved for National Institutes of Health (NIH) funding [10].

127

ORGANISMICALLY DEAD EMBRYOS

Reprogramming

ANT

Somatic cell

Transfer of altered somatic cell nucleus

Reprogramming with e.g. OCT4,SOX2 and NANOG

Classical Sperm

SBB

Organismically dead

Oocyte

Zygote Enucleated oocyte

8-cell embryo Biopsy Reprogrammed cell

Blastocyst

Dead embryos

1 bm

Harvesting of live cells

ZP

ICM Implantation in uterus

TE

hESC line

iPS line

hESC line

Isolated ICM Reactivation of CDX2

ANT pluripotent stem cell line

hESC line

Implantation in uterus

FIGURE 8.1 Classical and alternative strategies for the generation of human stem cells by reprogramming with exogenous genes (iPS), transfer of a genetically altered somatic cell nucleus into an oocyte (ANT), the classical derivation of human embryonic stem cells (hESCs) from blastocyst culture, the derivation of hESCs from a biopsied single blastomere (SBB), and the derivation from organismically dead embryos. bm, blastomere; ICM, inner cell mass; iPS, induced pluripotent stem cells; TE, trophectoderm; ZP, zona pellucida. Reproduced with permission from Gavrilov S, Papaioannou VE, Landry DW. Alternative strategies for the derivation of human embryonic stem cell lines and the role of dead embryos. Curr Stem Cell Res Ther 2009a;4:81e6.

ORGANISMICALLY DEAD EMBRYOS Our group proposed the derivation of hESC from irreversibly arrested, nonviable human embryos that died, despite best efforts, during the course of IVF for reproductive purposes [2]. This proposal to harvest live cells from dead embryos is analogous to the harvesting of essential organs from deceased donors. We suggested that the established ethical guidelines for essential organ donation could be employed for the clinical application of this paradigm to generate new hESC lines [2,4,27,28].

Irreversibility as a Criterion for Diagnosing Embryonic Death The modern concept of death is based on an irreversible loss of integrated organismic function [28,29]. Brain death is used as a reliable marker for irreversible loss of integrated function. Diagnosing the death of a patient before the

128

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death of that patient’s tissues is important for the appropriate application of medical resources and for the possibility of organ donation. To apply this concept to a stage of development that precedes the development of the nervous system, we proposed that an irreversible arrest of cell division would mark an irreversible loss of integrated function. Thus, it was necessary to find criteria that would establish irreversible cessation of normal embryonic development before every cell of the embryo has died. Through retrospective analysis of early-stage embryos that had been generated for reproductive purpose but were rejected owing to poor quality and/or developmental arrest, we showed that many of these embryos were in fact organismically dead [28]. Our data showed that the failure of normal cell division for 48 h was irreversible, and despite the possible presence of individual living cells, they indicated an irreversible loss of integrated organismic function: the conceptual definition of death [2,28]. Furthermore, we conducted a prospective study to characterize embryonic death [3,30], in which the progression of arrested embryos, including abnormal blastocysts, was examined in extended culture [27]. Our data demonstrated that developmental arrest observed in some human embryos by embryonic day 6 (ED6) after IVF cannot be reversed by extended culture in conditions suitable for preimplantation embryos, because we saw no morphological changes indicative of developmental progression in most embryos and observed no unequivocal instances of further cell divisions [27]. Moreover, these observations are in line with standard IVF practice, which dictates that such embryos should not be transferred or cryopreserved because they are known not to produce live offspring [27,31e38]. In an attempt to correlate morphology with cell number, we categorized the embryos at ED6 on the basis of gross morphology (Fig. 8.2) [27]. We showed that morphological categorization was of limited value in predicting cell number. Nevertheless, the higher cell number associated with cavitation might predict greater potential for the success of hESC derivation [27]. In addition, we determined the proportion of living and nonliving cells in nonviable ED6 human embryos (Fig. 8.2) and showed that most irreversibly arrested embryos contain a high proportion of vital cells regardless of the stage of arrest, which indicates that harvesting cells and deriving hESC from such nonviable embryos should be feasible [27].

Human Embryonic Stem Cell Lines Derived From Irreversibly Arrested, Nonviable Embryos In fact, the proof of principle for this alternative method has been obtained, because 14 hESC lines were successfully derived from nonviable embryos that were irreversibly arrested by our criteria (Table 8.1) [39,40,41]. The first cell line (hES-NCL9) was derived by Stojkovic and colleagues from 132 arrested embryos [40]. Subsequently, Daley and colleagues derived 11 lines from 413 poor-quality embryos rejected for clinical use [39]. In addition, our group derived two human ES lines: CU1 and CU2 from 159 ED6 irreversibly arrested, nonviable human embryos [41]. Although many arrested embryos might be expected to be aneuploid [42e46], all 14 hESC lines were karyotypically normal; moreover, pluripotency and differentiation potential were demonstrated in vitro and/or in vivo [39,40]; [41].

Morphological Criteria for Predicting the Capacity of Irreversibly Arrested, Nonviable Human Embryos to Develop Into a Human Embryonic Stem Cell Line To define morphological criteria that could be used to predict the capacity of discarded, irreversibly arrested, nonviable embryos to develop into an hESC line, we carried out a retrospective analysis of the morphological progression from ED5 to ED6 in 2480 embryos that were rejected for clinical use [41]. Embryos were given a morphological category commonly used for clinical grading as per standard IVF practice (e.g., single-celled embryo, multicell, morula, blastocyst). If an embryo had reached the blastocyst stage (i.e., showing advanced cavitation), it was given an overall grade of good, fair, or poor, and was also scored for inner cell mass and trophectoderm quality. Our analysis showed that nonviable embryos defined as poor did not improve with extended in vitro culture but retained the capacity to yield hESC lines despite arrested development [41]. We postulated that if derivation efforts were targeted on this subgroup, the derivation success rate could be increased and the production of new hESC lines could be brought closer to clinical application [41].

129

ORGANISMICALLY DEAD EMBRYOS

(A)

(B)

(C)

(D)

(E)

(F)

(G)

(H)

(I)

(J)

(K)

(L)

(M)

(N)

(O)

FIGURE 8.2 Morphology and differential propidium iodide/Hoechst fluorescent nuclear staining of nonviable embryos at ED6. Brightfield images (A, D, G, J, and M) with corresponding fluorescence images (B, E, H, K, and N), and enlarged details (C, F, I, L, and O) as indicated by the green squares (AeC). Category A embryo showing degeneration at embryonic day 6. All nuclei, including nuclear fragments, are pink, indicating that there are no living cells in the embryo. Detail shows pink nucleus from a dead cell (DeF). Category C embryo with living and dead cells is indicated by the blue and pink nuclei, respectively. Detail shows nuclei from one living and one dead cell. Arrow in E indicates a sperm nucleus outside the zona pellucida. (GeI) Category G embryo with living and dead cells as well as fragmented nuclei. Detail shows intact and fragmented nuclei (JeL). Category D embryo with all live cells. Detail shows blue fragmented nucleus (MeO). Category H embryo with many living and a few dead cells. Arrowheads in I and O indicate nuclear fragments. Reproduced with permission from Gavrilov S, Prosser RW, Khalid I, MacDonald J, Sauer MV, Landry DW et al. Non-viable human embryos as a source of viable cells for embryonic stem cell derivation. Reprod Biomed Online 2009b;18:301e8.

130 TABLE 8.1

8. ALTERNATIVE SOURCES OF HUMAN EMBRYONIC STEM CELLS

List of Human Embryonic Stem Cell Lines Derived From Nonviable Organismically Dead Embryos Embryoid Body Assay

Teratoma

Eligible for National Institutes of Health Funding?

References

Cell Line Name

Type of Embryo

Karyotype

Stem Cell Markers

hES-NCL9

Day 6e7 late arrested embryo (16e24 cells)

46,XX

Yes

Yes

Yes

ND

[40]

CHB-1

Day 3 PQE

46,XY

Yes

NR

Yes

Yes

[39]

CHB-2

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CHB-3

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CHB-4

Day 5 PQE

46,XY

Yes

NR

Yes

Yes

[39]

CHB-5

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CHB-6

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CHB-8

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CHB-9

Day 5 PQE

46,XY

Yes

NR

Yes

Yes

[39]

CHB-10

Day 5 PQE

46,XY

Yes

NR

Yes

Yes

[39]

CHB-11

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CHB-12

Day 5 PQE

46,XX

Yes

NR

Yes

Yes

[39]

CU1

Day 6 arrested poor blastocyst

46,XX

Yes

Yes

ND

ND

[41]

CU2

Day 6 arrested early blastocyst

46,XXa

Yes

Yes

ND

ND

[41]

ND, not determined; NR, not reported; PQE, poor-quality embryo. a Putative normal karyotype: possible low level of mosaicism.

CONCLUSION The derivation of hESC from organismically dead embryos is a unique approach because it defines a common ground in the human ES debate. Harvesting live cells from dead human embryos has the likelihood of being accepted by the staunchest opponents of embryo-destructive ES derivation. ES cells generated by this approach appear to be suitable for clinical research. Thus far, 11 human ES lines derived by Daley and colleagues have been included in the NIH stem cell registry and are available for research with NIH funding [10]. Human ES lines generated from organismically dead embryos are of equal quality compared with lines derived by the classical, intracellular massederived approach, but further characterization of these lines is needed [2]. During routine IVF procedures, large proportions of embryos fail to develop properly [45,47,48] and are discarded as being unsuitable for clinical use [2,27]. Despite the low efficiency of isolation of hESC from organismically dead embryos, large-scale derivation is not limited because in the United States alone, nearly half a million such embryos are generated yearly as a by-product of assisted reproductive technologies [2,27]. The prospect of thousands of hESC lines generated by this method and deposited into stem cell banks renders clinical applications based on human leukocyte antigen matching feasible.

References [1] Thomson JA, Itskovitz-Eldor J, Shapiro SS, Waknitz MA, Swiergiel JJ, Marshall VS, et al. Embryonic stem cell lines derived from human blastocysts. Science 1998;282:1145e7. [2] Gavrilov S, Papaioannou VE, Landry DW. Alternative strategies for the derivation of human embryonic stem cell lines and the role of dead embryos. Curr Stem Cell Res Ther 2009a;4:81e6. [3] Green RM. Can we develop ethically universal embryonic stem-cell lines? Nat Rev Genet 2007;8:480e5. [4] Landry DW, Zucker HA. Embryonic death and the creation of human embryonic stem cells. J Clin Invest 2004;114:1184e6. [5] McLaren A. A scientist’s view of the ethics of human embryonic stem cell research. Cell Stem Cell 2007;1:23e6. [6] Guenin LM. The morality of unenabled embryo use e arguments that work and arguments that don’t. Mayo Clin Proc 2004;79:801e8.

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J In Vitro Fert Embryo Transf 1989;6:30e5. [32] Cummins JM, Breen TM, Harrison KL, Shaw JM, Wilson LM, Hennessey JF. A formula for scoring human embryo growth rates in in vitro fertilization: its value in predicting pregnancy and in comparison with visual estimates of embryo quality. J In Vitro Fert Embryo Transf 1986; 3:284e95. [33] Erenus M, Zouves C, Rajamahendran P, Leung S, Fluker M, Gomel V. The effect of embryo quality on subsequent pregnancy rates after in vitro fertilization. Fertil Steril 1991;56:707e10. [34] Giorgetti C, Terriou P, Auquier P, Hans E, Spach JL, Salzmann J, et al. Embryo score to predict implantation after in-vitro fertilization: based on 957 single embryo transfers. Hum Reprod 1995;10:2427e31. [35] Puissant F, van Rysselberge M, Barlow P, Deweze J, Leroy F. Embryo scoring as a prognostic tool in IVF treatment. Hum Reprod 1987;2:705e8. [36] Staessen C, Camus M, Bollen N, Devroey P, van Steirteghem AC. The relationship between embryo quality and the occurrence of multiple pregnancies. Fertil Steril 1992;57:626e30. [37] Steer CV, Mills CL, Tan SL, Campbell S, Edwards RG. The cumulative embryo score: a predictive embryo scoring technique to select the optimal number of embryos to transfer in an in-vitro fertilization and embryo transfer programme. Hum Reprod 1992;7:117e9. [38] Ziebe S, Petersen K, Lindenberg S, Andersen AG, Gabrielsen A, Andersen AN. Embryo morphology or cleavage stage: how to select the best embryos for transfer after in-vitro fertilization. Hum Reprod 1997;12:1545e9. [39] Lerou PH, Yabuuchi A, Huo H, Takeuchi A, Shea J, Cimini T, et al. Human embryonic stem cell derivation from poor-quality embryos. Nat Biotechnol 2008;26(2):212e4. [40] Zhang X, Stojkovic P, Przyborski S, Cooke M, Armstrong L, Lako M, et al. Derivation of human embryonic stem cells from developing and arrested embryos. Stem Cell 2006;24:2669e76. 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C H A P T E R

9 Stem Cells From the Amnion Paolo De Coppi1,2, Anthony Atala2 1

UCL Institute of Child Health and Great Ormond Street Hospital, London, United Kingdom; 2Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

INTRODUCTION In this chapter, we provide an overview of the potential advantages and disadvantages of different stem and progenitor cell populations identified in the amnion and in the amniotic fluid (AF), along with their properties and potential clinical applications. Placenta, fetal membranes (i.e., amnion and chorion), and AF have been extensively investigated as a potential noncontroversial source of stem cells. They are usually discarded after delivery and are accessible during pregnancy through amniocentesis and chorionic villus sampling [1]. Several populations of cells with multilineage differentiation potential and immunomodulatory properties have been isolated from the human placenta and fetal membranes; they have been classified by an international workshop [2] as human amniotic epithelial cells (hAECs) [3e7], human amniotic mesenchymal stromal cells (hAMSCs) [8,9], human chorionic mesenchymal stromal cells (hCMSCs) [10,11], and human chorionic trophoblastic cells (hCTCs). In the AF, two main populations of stem cells have been isolated: amniotic fluid mesenchymal stem cells (AFMSCs) and amniotic fluid stem (AFS) cells. Because of the easier accessibility of the AF compared with other extraembryonic tissues, these cells may hold much promise in regenerative medicine.

PLACENTA: FUNCTION, ORIGIN, AND COMPOSITION During human placental development, a range of cell types is generated, depending on gestation, which can be described as epithelial (because they derive from the amniotic membrane), trophoblastic, and hematopoietic (both derived from the chorionic villi). Placental tissue has contributions from both the fetus (amniotic membrane epithelium, extraembryonic mesoderm, and the two-layered trophoblast) and the mother (decidua basalis). The primitive formation of the placenta occurs from cells of fetal origin, known as trophoblast, that invade the uterine endometrium, form the outer layer of the blastocyst, and produce a network of protrusions, the villi and the lacunae system. On the 7th to 10th day after conception, the chorionic membranes are developed from layers of proliferating placental cells. At day 9 postconception, the inner cell mass induces the formation of the epiblast and hypoblast that subsequently become the amniotic cavity and the yolk sac. The process of gastrulation enables the bilaminar disc to differentiate into the three germ layers (ectoderm, mesoderm, and endoderm), followed by organogenesis [12,13]. During the maturation of the syncytium, the villi establish the maternofetal transport of blood nutrients, oxygen, gases, and waste products, and they differentiate from mesenchymal villi into immature intermediate villi. The placental progenitor stem cells are the cytotrophoblast cells emanating from the trophectodermal layer that result in the villous syncytiotrophoblast, which is a multinucleated aggregate of cytotrophoblast cells, and the extravillous cytotrophoblasts (EVTs) [14]. The trophoblast invasion of the maternal decidualized endometrium is also associated with hormonal secretions, such as human chorionic gonadotrophin, which downregulates maternal cellular immunity and promotes angiogenic activity of the EVTs [15]. By the 12th week of gestation, the placenta has adopted a Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00009-6

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Copyright © 2019 Elsevier Inc. All rights reserved.

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9. STEM CELLS FROM THE AMNION

hemotrophic nutritional function, in particular causing extensive transformation of the maternal spiral arteries into augmented high-flow, low-resistance vessels that provide sufficient nutrients and oxygen for the developing fetus. Human fetal placental cells can be divided into the hAECs, hAMSCs, hCMSCs, and hCTCs. Isolation of these latter cells (hCMSCs and hCTCs) can be realized directly from the placenta or through chorionic villus sampling, a form of invasive ultrasound-guided prenatal diagnosis that entails sampling of the placental tissue from 11 weeks of gestation. Isolation of the heterogeneous population of cells (hAECs, hCMSCs, and hCTCs) can be implemented from different placenta regions through enzymatic digestion with dispase or collagenase in synergy with DNase.

AMNIOTIC FLUID: FUNCTION, ORIGIN, AND COMPOSITION The AF is the clear, watery liquid that surrounds the growing fetus within the amniotic cavity. It allows the fetus to grow freely and move inside the uterus, protects it from outside injuries by cushioning sudden blows or movements and by maintaining consistent pressure and temperature, and acts as a vehicle for the exchange of body chemicals with the mother [16,17]. In humans, the AF starts to appear at the beginning of the second week of gestation as a small film of liquid between the cells of the epiblast. Between days 8 and 10 after fertilization, this fluid gradually expands and separates the epiblast (i.e., the future embryo) from the amnioblasts (i.e., the future amnion), thus forming the amniotic cavity [6]. Thereafter, it progressively increases in volume, completely surrounding the embryo after the fourth week of pregnancy. Over the course of gestation, AF volume markedly changes from 20 mL in the seventh week to 600 mL in the 25th week, 1000 mL in the 34th week, and 800 mL at birth. During the first half of gestation, the AF results from active sodium and chloride transport across the amniotic membrane and the nonkeratinized fetal skin, with concomitant passive movement of water. In the second half of gestation, the AF is composed of fetal urine, gastrointestinal excretions, respiratory secretions, and substances exchanged through the sac membranes [18e21]. The AF is primarily composed of water and electrolytes (98%e99%) but it also contains chemical substances (e.g., glucose, lipids, proteins, hormones, and enzymes), suspended materials (e.g., vernix caseosa, lanugo hair, and meconium), and cells. AF cells are derived both from extraembryonic structures (i.e., placenta and fetal membranes) and from embryonic and fetal tissues [22]. Although AF cells are known to express markers of all three germ layers [23], their exact origin still represents a matter of discussion; the consensus is that they consist mainly of cells shed in the amniotic cavity from the developing skin, respiratory apparatus, and urinary and gastrointestinal tracts [18,24]. AF cells display a broad range of morphologies and behaviors varying with gestational age and fetal development [25]. Under normal conditions, the number of AF cells increases with advancing gestation; if a fetal disease is present, AF cell counts can be either dramatically reduced (e.g., intrauterine death, urogenital atresia) or abnormally elevated (e.g., anencephaly, spina bifida, exomphalos) [26]. Based on their morphological and growth characteristics, viable adherent cells from the AF are classified into three main groups: epithelioid (33.7%), AF (60.8%), and fibroblastic type (5.5%) [25]. In the event of fetal abnormalities, other types of cells can be found in the AF, e.g., neural cells in the presence of neural tube defects and peritoneal cells in case of abdominal wall malformations [26e28]. Most cells present in the AF are terminally differentiated and have limited proliferative capabilities [26,29]. In the 1990s, however, two groups demonstrated the presence in the AF of small subsets of cells harboring a proliferation and differentiation potential. First, Torricelli reported the presence of hematopoietic progenitors in the AF collected before the 12th week of gestation [30]. Then Streubel was able to differentiate AF cells into myocytes, which suggested the presence in the AF of nonhematopoietic precursors [31]. These results initiated new interest in the AF as an alternative source of cells for therapeutic applications.

AMNIOTIC EPITHELIAL CELLS Human amnion consists of amniotic epithelial cells (AECs) on a basement collagenous membrane, an acellular compact layer filled with reticular fibers, a fibroblast layer with Hofbauer cells/histiocytes, and a highly hygroscopic spongy layer with fibrils between the chorion and the amniotic sac [32]. AECs can be obtained with differential enzymatic digestion from the amnion membrane after it is separated from the underlying chorion [33]. The amnion contains epithelial cells expressing surface markers that include both embryonic-specific markers such as the stage-specific antigens (SSEAs) 3 and 5, Tra-1-60, Tra-1-81, and mesenchymal markers CD105, CD90, CD73, CD44, CD29, human leukocyte antigen (HLA)-A, -B, -C, CD13, CD10, CD166, and

AMNIOTIC MESENCHYMAL STEM CELLS

135

CD117. Their immunological properties have not been completely elucidated; however, AECs also appear to be resistant to rejection after allotransplantation, probably owing to their immunosuppression properties (CD59 and HLA-G [34]), which could also lead to their therapeutic role in a disease model. For example, hAECs have been reported to lower the blood glucose levels of streptozotocin-induced diabetic mice several weeks after implantation, potentially by differentiation into b cells [35]. hAECs are considered multipotent cells because of their capacity to differentiate into different lineages. In particular, they have been able to mature under specific culture conditions into neuronal cells that synthesize acetylcholine, norepinephrine, and dopamine [6,36,37]. In vivo, hAECs have been reported to be neuroprotective and neuroregenerative, probably in relation to growth factor secretion. Indeed, studies showed that hAECconditioned media exhibit neurotrophic effects on rat cortical cells [38], and because of the expression of neural markers such as nestin, glial fibrillary acidic protein, and microtubule-associated protein 2, they are inclined toward neuronal lineages. hAECs have also been used to treat peripheral nerve injuries in animal models, in which they have shown to enhance the growth of host neurons and guide regenerative sprouting [68a]. In addition to their neurogenic potential, some relevant work has been done on exploring the hepatic potential of AECs. First, AECs produce albumin and a-fetoprotein, and they show glycogen storage and hepatic differentiation potential in vitro [39,40]. Moreover, in vitro, hAECs had the capacity to metabolize ammonia, testosterone and 17a-hydroxyprogesterone caproate, whereas they expressed hepatocyte markers such as albumin, A1AT, CYP2A4, 3A7, 1A2, 2B6, ASGPR1, and inducible fetal cytochromes. After intrahepatic transplantation into immunodeficient (severe combined immunodeficient [SCID])/beige mice, hAECs demonstrated functional hepatic characteristics [39], and after pretreatment of SCID/beige mice with retrorsine, hAECs expressed mature liver genes, plasma proteins, and hepatic enzymes to a level equal to adult liver tissue. hAECs showed therapeutic efficacy after transplantation in a mouse model of cirrhosis [41]. The transplantation of human amnionederived epithelial cells to the liver appears to have desired therapeutic properties including the secretion of matrix metalloproteinase that instigate fibrinolysis and increase in interleukin-10 concentration. In a liver disease mouse model, amnion epithelial cell transplantation resulted in hepatic engraftment with decreased inflammation, fibrosis, and hepatocyte apoptosis [42]. Zhang et al. infused amnion-derived cells in a carbon tetrachloride e-treated mouse liver, and they showed minimal fibrosis and apoptosis [43]. Ricci et al. used a piece of human amniotic membrane (hAM) to assess fibrosis in rat liver and demonstrated increased antifibrotic properties of the hAM with a reduction in ductular reaction and extracellular matrix (ECM) deposition [44]. Vaghjiani et al. showed that the differentiation of hAECs into hepatic-like cells can remain viable and functional after encapsulation in barium alginate microspheres in vitro, and they can express CYP3A4, which is thought to break down nearly 50% of all therapeutic drugs [45]. Moreover, the cryopreserved amniotic membrane and its by-products have been recognized as significant tools for the treatment of ulceration and epithelial defects (corneal or conjunctival) [46,47]. Nakamura et al. used autologous serum corneal epithelial cells on an amniotic membrane to transplant nine eyes of nine patients with total limbal stem cell deficiency, and they demonstrated improvement in visual acuity and complete corneal epithelialization within 2e5 days [48]. Wang et al. experimented on allogeneic green fluorescent protein (GFP)þ mice intact amniotic epithelium grafts with syngeneic (EGFP-C57BL/6 to C57BL/6 W/t) and allogeneic (EGFP-C57BL/6 to BALB/c W/t) AE cells that were transplanted into the cornea or conjunctiva or inserted into the anterior chambers. The researchers showed that major histocompatibility complex (MHC) class I beta antigens were minimally expressed after implantation [49]. Finally, other potential clinical applications have been explored and AECs have been considered potentially useful for a broad variety of conditions including ophthalmic diseases, lung fibrosis, liver fibrosis, multiple sclerosis, congenital metabolic disorders such as ornithine transcarbamylase deficiency, familial hypercholesterolemia, spinal cord injuries, and Parkinson disease and for allogeneic cell transplantations [50e52].

AMNIOTIC MESENCHYMAL STEM CELLS Mesenchymal stem cells (MSCs) represent a population of multipotent stem cells able to differentiate toward mesoderm-derived lineages (i.e., adipogenic, chondrogenic, myogenic, and osteogenic) [53]. Initially identified in adult bone marrow, where they represent 0.001%e0.01% of total nucleated cells [54], MSCs have since been isolated from several adult (e.g., adipose tissue, skeletal muscle, liver, and brain), fetal (i.e., bone marrow, liver, and blood), and extraembryonic tissues (i.e., placenta and amnion) [55]. The presence of a subpopulation of AF cells with mesenchymal features able to proliferate in vitro more rapidly than comparable fetal and adult cells was described for the first time in 2001 [56]. In 2003, In ’t Anker demonstrated that the AF can be an abundant source of fetal cells that exhibit a phenotype and a multilineage differentiation

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9. STEM CELLS FROM THE AMNION

potential similar to that of bone marrowederived MSCs (BM-MSCs); these cells were named AFMSCs [57]. Soon after that article, other groups independently confirmed similar results.

Isolation and Culture AFMSCs can easily be obtained: in humans, from small volumes (2e5 mL) of second- and third-trimester AF [57a,61], where their percentage is estimated to be 0.9%e1.5% of the total AF cells [58]; and in rodents, from the AF collected during the second or third week of pregnancy [59,60]. Various protocols have been proposed for their isolation; all are based on the expansion of unselected populations of AF cells in serum-rich conditions without feeder layers, allowing cell selection by culture conditions. The success rate of the isolation of AFMSCs was reported by different authors to be 100% [61]. AFMSCs grow in basic medium containing fetal bovine serum (20%) and fibroblast growth factor (5 ng/mL). It was shown that human AFMSCs can be also cultured in the absence of animal serum without losing their properties [62]; this finding is a fundamental prerequisite for the beginning of clinical trials in humans.

Characterization The fetal versus maternal origin of AFMSCs has been investigated by different authors. Molecular HLA typing and amplification of the SRY gene in AF samples collected from male fetuses [57,58] demonstrated the exclusive fetal derivation of these cells. However, whether AFMSCs originate from the fetus or from the fetal portion of extraembryonic tissues remains a matter of debate [62]. AFMSCs display a uniform spindle-shaped, fibroblast-like morphology similar to that of other MSCs populations and expand rapidly in culture [63]. Human cells derived from a single 2-mL AF sample can increase to 180  106 cells within 4 weeks (three passages), and as demonstrated by growth kinetics assays, they possess a greater proliferative potential (average doubling time, 25e38 h) compared with that of BM-MSCs (average doubling time, 30e90 h) [57,58,60,64]. Moreover, AFMSC clonogenic potential has been proved to exceed that of MSCs isolated from bone marrow (86  4.3 versus 70  5.1 colonies) [60]. Despite their high proliferation rate, AFMSCs retain a normal karyotype and do not display tumorigenic potential even after extensive expansion in culture [58,64]. Analysis of AFMSC transcriptome demonstrated that: (1) the AFMSC gene expression profile, as well as that of other MSC populations, remains stable between passages in culture, enduring cryopreservation and thawing well; (2) AFMSCs share with MSCs derived from other sources a core set of genes involved in ECM remodeling, cytoskeletal organization, chemokine regulation, plasmin activation, transforming growth factor-b, and Wnt signaling pathways; and (3) compared with other MSCs, AFMSCs show a unique gene expression signature that consists of the upregulation of genes involved in signal transduction pathways (e.g., HHAT, F2R, and F2RL) and in uterine maturation and contraction (e.g., OXTR and PLA2G10), which suggests a role of AFMSCs in modulating interactions between the fetus and the uterus during pregnancy [63]. Different investigators determined the cell surface antigenic profile of human AFMSCs through flow cytometry (Table 9.1). Cultured human AFMSCs are positive for mesenchymal markers (i.e., CD90, CD73, CD105, and CD166), for several adhesion molecules (i.e., CD29, CD44, CD49e, and CD54), and for antigens belonging to MHC-I. They are negative for hematopoietic and endothelial markers (e.g., CD45, CD34, CD14, CD133, and CD31). AFMSCs exhibit a broad differentiation potential toward mesenchymal lineages. Under specific in vitro inducing conditions, they are able to differentiate toward the adipogenic, osteogenic, and chondrogenic lineage [57,60,63]. Despite not being pluripotent, AFMSCs can be efficiently reprogrammed into pluripotent stem cells (iPS) via retroviral transduction of defined transcription factors (Oct4, Sox2, Klf-4, and c-Myc). Strikingly, AFMSC reprogramming capacity is significantly higher (100-fold) and much quicker (6 days versus 16e30 days) compared with that of somatic cells such as skin fibroblasts. Similarly to iPS derived from other sources, iPS derived from AF cells generate embryoid bodies (EBs) and differentiate toward all three germ layers in vitro; in vivo, they form teratomas when injected into SCID mice [65].

Preclinical Studies After the identification of AFMSCs, various studies investigated their therapeutic potential in different experimental settings. Different groups demonstrated that AFMSCs are able not only to express cardiac and

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AMNIOTIC MESENCHYMAL STEM CELLS

TABLE 9.1 Immunophenotype of Culture-Expanded Second- and Third-Trimester Human Amniotic Fluid Mesenchymal Stromal Cells: Results by Different Groups Markers

Antigen

CD No.

You [57a]

Roubelakis [58]

Tsai [61]

In ’t Anker [57]

Mesenchymal

SH2, SH3, SH4

CD73

þ

þ

þ

þ

Thy1

CD90

þ

þ

þ

þ

Endoglin

CD105

þ

þ

þ

þ

SB10/ALCAM

CD166

nt

þ

nt

þ

Leukocyte common antigen

CD14

nt



nt



gp105-120

CD34

nt







Lipopolysaccharide-R

CD45









Prominin-1

CD133

nt



nt

nt

b1-integrin

CD29

þ

þ

þ

nt

Endothelial and hematopoietic

Integrins

Selectins

Immunoglobulin superfamily

Major histocompatibility complex

b3-integrin

CD61



nt

nt

nt

a4-integrin

CD49d

nt



nt



a5-integrin

CD49e

nt

þ

nt

þ

Lymphocyte function eassociated-1

CD11a

nt

þ

nt



E-Selectin

CD62E

nt

þ

nt



P-selectin

CD62P

nt

þ

nt



Platelet endothelial cell adhesion molecule-1

CD31



þ





Intercellular adhesion molecule (ICAM)-1

CD54

nt

þ

nt

þ

ICAM-3

CD50

nt



nt



Vascular cell adhesion protein molecule-1

CD106

nt

þ

nt



Homing cell adhesion molecule-1

CD44

nt

þ

þ

þ

I (human leukocyte antigen [HLA]-ABC)

none

nt

þ

þ

þ

II (HLA-DR, DP, DQ)

none

nt

nt





nt, not tested.

endothelial-specific markers under specific culture conditions, but also to integrate into normal and ischemic cardiac tissue, where they differentiate into cardiomyocytes and endothelial cells [66e69]. In a rat model of bladder cryoinjury, AFMSCs show the ability to differentiate into smooth muscle and prevent the compensatory hypertrophy of surviving smooth muscle cells [59]. AFMSCs can be a suitable cell source for tissue engineering of congenital malformations. In an ovine model of diaphragmatic hernia, repair of the muscle deficit using grafts engineered with autologous mesenchymal amniocytes leads to better structural and functional results compared with equivalent fetal myoblast-based and acellular implants [62,70]. Engineered cartilaginous grafts have been derived from AFMSCs grown on biodegradable meshes in serum-free chondrogenic conditions for at least 12 weeks; these grafts have been successfully used to repair tracheal defects in fetal lambs when implanted in utero [62]. The surgical implantation of AFMSCs seeded on nanofibrous scaffolds and predifferentiated in vitro toward the osteogenic lineage into a leporine model of sternal defect led to complete bone repair in 2 months [71]. Intriguingly, studies suggested that AFMSCs can harbor trophic and protective effects in the central and peripheral nervous systems. Pan showed that AFMSCs facilitate peripheral nerve regeneration after injury and

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hypothesized that this can be determined by cell secretion of neurotrophic factors [72e74]. After transplantation into the striatum, AFMSCs are capable of surviving and integrating in the rat adult brain and of migrating toward areas of ischemic damage [75]. Moreover, the intraventricular administration of AFMSCs in mice with focal cerebral ischemia-reperfusion (IR) injuries significantly reverses neurological deficits in treated animals [76]. Remarkably, it was also observed that AFMSCs present in vitro had an immunosuppressive effect similar to that of BM-MSCs [77]. After stimulation of peripheral blood mononuclear cells with anti-CD3, anti-CD28, or phytohemagglutinin, irradiated AFMSCs demonstrated a significant inhibition of T-cell proliferation with a dosedependent kinetics [64].

AMNIOTIC FLUID STEM CELLS The first suggestion that the AF may contain undifferentiated cells was based on the observation that AF-derived cells expressed skeletal muscle proteins when cultured in the supernatant of rhabdomyosarcoma cell lines [31]. Subsequently, AF-derived cells were shown to differentiate into osteocytes, adipocytes, and fibroblasts while having a cell marker profile comparable to MSCs [57]. Brivanlou and colleagues were the first to confirm that a subpopulation of AF-derived cells (approximately 0.5%e1% of total live cells) have stem cell potential, by demonstrating the expression of octamer transcription factor-4 (Oct-4) at the transcriptional and protein levels [78]. Remarkably, Karlmark et al. transfected human AF cells with the GFP gene under either the Oct-4 or the Rex-1 promoter and established that some AF cells were able to activate these promoter [79]. Subsequently, we and others used CD117 (c-Kit; type III tyrosine kinase receptor for stem cell factor with essential roles in gametogenesis, melanogenesis, and hematopoiesis) as a means to select the undifferentiated population from the AF [80,81]. These CD117-expressing cells are a heterogeneous population; they were isolated from small [80e83] and large animals [84] as well as humans [80,85] and are known as AFSC.

Isolation and Culture The proportion of c-kitþ cells in the AF varies over the course of gestation, roughly describing a Gaussian curve; they appear at very early time points in gestation (i.e., at 7 weeks of amenorrhea in humans and at embryonic day (E)9.5 in mice) and present a peak at midgestation equal to 90  104 cells/fetus at 20 weeks of pregnancy in humans and 10,000 cells/fetus at E12.5 in mice [81]. Human AFS cells (AFSC) can be derived either from small volumes (5 mL) of second-trimester AF (14e22 weeks of gestation) or from confluent backup amniocentesis cultures. Murine AFSC are obtainable from the AF collected during the second week of gestation (E11.5e14.5) [80,86e88]. AFSC isolation is based on a two-step protocol consisting of the prior immunological selection of ckitepositive cells from the AF (approximately 1% of total AF cells) and in the subsequent expansion of these cells in culture [80,87,89e92]. Isolated AFSC can be expanded in feeder layer-free, serum-rich conditions without evidence of spontaneous differentiation in vitro. Cells are cultured in basic medium containing 15% of fetal bovine serum and Chang supplement [80,92].

Characterization Karyotype analysis of human AFSC deriving from pregnancies in which the fetus was male revealed the fetal origin of these cells [80]. AFSC proliferate well during ex vivo expansion. When cultivated, they display a spectrum of morphologies ranging from a fibroblast-like to an oval-round shape (Fig. 1A). As demonstrated by different authors, AFSC possess great clonogenic potential [80,88]. Clonal AFSC lines expand rapidly in culture (doubling time, 36 h); more interesting, they maintain a constant telomere length (20 kilobase pairs) between early and late passages (Fig. 1B). Almost all clonal AFSC lines express markers of a pluripotent undifferentiated state: Oct4 and NANOG [80,88,89,92,93]. However, they have been proved not to form tumors when injected in SCID mice [80]. Different investigators determined the cell surface antigenic profile of AFSC through flow cytometry (Table 9.2). Cultured human AFSC are positive for embryonic stem (ES) cell (e.g., SSEA-4) and mesenchymal markers (e.g., CD73, CD90, and CD105), for several adhesion molecules (e.g., CD29 and CD44) and for antigens belonging to MHC-I. They are negative for hematopoietic and endothelial markers (e.g., CD14, CD34, CD45, CD133, and CD31), and for antigens belonging to MHC-II.

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AMNIOTIC FLUID STEM CELLS

(A)

(B)

1

2

3

4

(C)

kbp 1

21.2• 7.4• 5.0• 3.6•

6

2

7

3

8

4

9

5

10

11

12

2.0• 13

14

15

16

17

18

19

20

21

22

X

Y

FIGURE 9.1 (A) Human amniotic fluid stem cells (AFSC) mainly display a spindle-shaped morphology during in vitro cultivation under

feeder layer-free, serum-rich conditions. (B, C) Clonal human AFSC lines retain long telomeres and a normal karyotype after more than 250 cell divisions (magnification 20). (B) Conserved telomere length of AFSC between early passage (20 population doublings, lane 3) and late passage (250 population doublings, lane 4). Short-length (lane 1) and high-length (lane 2) telomere standards provided in the assay kit. (C) Giemsa band karyogram showing chromosomes of late-passage (250 population doublings) cells. Picture adapted from De Coppi (2007b).

TABLE 9.2 Surface Markers Expressed by Human c-kitþ Amniotic Fluid Stem Cells: Results by Different Groups Markers

Antigen

CD No.

Ditadi [81]

De Coppi [80]

Kim [3]

Tsai [88]

Embryonic stem cells

Stage-specific antigen (SSEA)-3

None

nt



þ

nt

SSEA-4

None

nt

þ

þ

nt

Tra-1-60

None

nt



þ

nt

Tra-1-81

None

nt



nt

nt

SH2, SH3, SH4

CD73

nt

þ

nt

þ

Thy1

CD90

þ

þ

nt

þ

Endoglin

CD105

nt

þ

nt

þ

Leukocyte common antigen

CD14

nt

nt

nt



gp105-120

CD34





nt



Lipopolysaccharide-R

CD45

þ



nt

nt

Prominin-1

CD133





nt

nt

Integrins

b1-integrin

CD29

nt

þ

nt

þ

Immunoglobulin superfamily

Platelet endothelial cell adhesion molecule-1

CD31

nt

nt

þ

nt

Intercellular adhesion molecule-1

CD54

nt

nt

þ

nt

Vascular cell adhesion protein molecule-1

CD106

nt

nt

þ

nt

Homing cell adhesion molecule-1

CD44

þ

þ

þ

þ

I (human leukocyte antigen [HLA]-ABC)

None

þ

þ

þ

þ

II (HLA-DR, DP, DQ)

None









Mesenchymal

Endothelial and hematopoietic

Major histocompatibility complex

nt, not tested.

Because the stability of cell lines is a fundamental prerequisite for basic and translational research, the capacity of AFSC to maintain their baseline characteristics over passages has been evaluated based on multiple parameters. Despite their high proliferation rate, AFSC and derived clonal lines show a homogeneous, diploid DNA content without evidence of chromosomal rearrangement even after expansion to 250 population doublings [80,89] (Fig. 1C). Moreover, AFSC maintain constant morphology, doubling time, apoptosis rate, cell cycle distribution,

140

9. STEM CELLS FROM THE AMNION

and marker expression (e.g., Oct4, CD117, CD29, and CD44) up to 25 passages [89,92]. During in vitro expansion, however, cell volume tends to increase and significant fluctuations of proteins involved in different networks (i.e., signaling, antioxidant, proteasomal, cytoskeleton, connective tissue, and chaperone proteins) can be observed using a gel-based proteomic approach [89]; the significance of these modifications warrants further investigations but needs to be taken in consideration when interpreting experiments over several passages and comparing results from different groups. AFSC and, more important, derived clonal cell lines are able to differentiate toward tissues representative of all three embryonic germ layers spontaneously, when cultured in suspension to form EBs, and when grown under specific differentiation conditions. EBs consist of three-dimensional aggregates of ES cells, which recapitulate the first steps of early mammalian embryogenesis [93ae93c]. As ES cells, when cultured in suspension and without antidifferentiation factors, AFSC harbor the potential to form EBs with high efficiency: the incidence of EB formation (i.e., the percentage of EB recovered from 15 hanging drops) is estimated to be around 28% for AFSC lines and around 67% for AFSC clonal lines. Similar to ES cells, EB generation by AFSC is regulated by the mammalian target of rapamycin pathway and is accompanied by a decrease of Oct4 and Nodal expression and by an induction of endodermal (GATA4), mesodermal (Brachyury and HBE1) and ectodermal (Nestin and Pax6) markers [92,94]. Under specific mesenchymal differentiation conditions, AFSC express molecular markers of adipose, bone, muscle, and endothelial differentiated cells (e.g., lipoprotein lipase, desmin, osteocalcin, and vascular cell adhesion protein 1). In the adipogenic, chondrogenic, and osteogenic medium, AFSC respectively develop intracellular lipid droplets, secrete glycosaminoglycans, and produce mineralized calcium [86,88]. Under conditions inducing cell differentiation toward the hepatic lineage, AFSC express hepatocyte-specific transcripts (e.g., albumin, a-fetoprotein, and multidrug-resistant membrane transporter 1) and acquire the liver-specific function of urea secretion (Fig. 2A) [80]. Under neuronal conditions, AFSC are capable of entering the neuroectodermal lineage. After induction, they express neuronal markers (e.g., G proteinecoupled inwardly rectifying Kþ potassium channels), exhibit barium-sensitive potassium current, and release glutamate after stimulation (Fig. 2B). Ongoing studies are investigating AFSC capacity to yield mature, functional neurons [95e97]. AFSC can be easily manipulated in vitro. They can be transduced with viral vectors more efficiently than can adult MSCs, and after infection, they maintain their antigenic profile and the ability to differentiate into different lineages [98]. AFSC labeled with superparamagnetic micrometer-sized iron oxide particles retain their potency and can be tracked noninvasively by magnetic resonance imaging (MRI) for at least 4 weeks after injection in vivo [99].

Preclinical Studies Several reports have investigated the potential applications of AFSC in different settings.

(A)

(B)

(C)

FIGURE 9.2 Amniotic fluid stem cell (AFSC) differentiation into lineages representative of the three embryonic germ layers. (A) Hepatogenic differentiation: urea secretion by human AFSC before (rectangles) and after (diamonds) hepatogenic in vitro differentiation. (B) Neurogenic differentiation: secretion of neurotransmitter glutamic acid in response to potassium ions. (C) Osteogenic differentiation: mouse micro-computed tomography scan 18 weeks after implantation of printed constructs of engineered bone from human AFSCs; arrowhead: region of implantation of control scaffold without AFSC; rhombus: scaffolds seeded with AFSC; *, bone deposit after transplantation. Picture adapted from De Coppi (2007b).

AMNIOTIC FLUID STEM CELLS

141

Musculoskeletal System We investigated the osteogenic potential of AFSC by inducing osteogenic differentiation with culture media containing dexamethasone, b-glycerophosphate, and ascorbic acid-2-phosphate. After seeding in a collagen/alginate scaffold, they were implanted subcutaneously in immunodeficient mice. At 18 weeks, microecomputed tomography revealed highly mineralized tissues and blocks of bone-like material [80]. In a study by Sun et al., osteogenic differentiation of human AFSC was achieved using bone morphogenetic protein-7 and seeding on nanofibrous scaffolds, as evident by alkaline phosphatase activity, calcium content, von Kossa staining, and the expression of osteogenic genes. Implantation into the subcutaneous space led to bone formation in 8 weeks with positive von Kossa staining and a radio-opaque profile upon X-ray [100]. A series of experiments by the Goldberg laboratory investigated the osteogenesis of AFSC after seeding on a poly(ε-caprolactone) (PCL) biodegradable polymer. Cells that were differentiated in a three-dimensional PCL scaffold deposited mineralized matrix and were viable after 15 weeks of culture. It was also shown that predifferentiated cells in vitro produced seven times more mineralized matrix when implanted subcutaneously in vivo [101]. When the authors compared AFSC and BM-MSC for osteogenesis after seeding on scaffolds and long-term in vitro culture, they came across some striking results. Although BM-MSC differentiated more rapidly than AFSC, the growth and production of mineralized matrix halted at 5 weeks. In contrast, AFSC continued to produce matrix for up to 15 weeks, which led to a fivefold increase in overall production compared with BM-MSC seeded scaffolds. We demonstrated for the first time the functional and stable long-term integration of AFSC into the skeletal muscle of human a-skeletal actineCre SmnF7/F7 mutant mice, which closely replicates the clinical features of human muscular dystrophy [102]. AFSC were obtained from E11.5e13.5 GFPþ mice through direct CD117 selection immediately after AF collection. Approximately 25,000 freshly isolated AFSC were injected into the tail vein of each animal without previous expansion in culture. Transplanted mice displayed enhanced muscle strength, improved survival rate by 75%, and restored muscle phenotype compared with untreated animals. Not only was dystrophin distribution in GFPþ myofibers similar to that of wild-type animals, GFPþ cells were found to be engrafted into the muscle stem cell niche (as demonstrated by their sublaminal position and by Pax7 and a7-integrin expression). Functional integration of AFSC in the stem cell niche was confirmed by successful secondary transplants of GFPþ satellite cells derived from AFSC-treated mice into untreated SmnF7/F7 mutant mice. To progress toward their application for therapy, the therapeutic potential of cultured AFSC was investigated and 25,000 AFSC, expanded under “embryonic-type” conditions, were intravenously injected into SmnF7/F7 mice. Cultured AFSC regenerated approximately 20% of the recipient muscle fibers compared with 50% when freshly isolated AFSC were used; this highlighted the importance of optimizing cell expansion protocols. Available data suggest that AFSC can differentiate toward myogenic lineages, engraft into the muscle stem cell niche, and participate in muscle regeneration in animal models of muscle injury. Hence, AFSC constitute a promising therapeutic option for skeletal muscle degenerative diseases. Nervous System We previously investigated human AFSC injection in the brain of Twitcher mice (a model of Krabbe globoid leukodystrophy associated with progressive oligodendrocyte and neuronal loss). Human AFSC engrafted into the lateral cerebral ventricles differentiated to cells that were similar to the surrounding environment and survived for up to 2 months. It was also demonstrated that engraftment was variable; 70% of AFSC survived in the brain of Twitcher mice, in contrast to only 30% of AFSC in the brain of normal mice [80]. A study by Prasongchean et al. indicated that treatment with small molecules, which normally leads to neuronal differentiation and grafting of AFSC into environments such as organotypic rat hippocampal cultures and the embryonic chick nervous system, led to no expression of neural cell markers. However, AFSC reduced hemorrhage and increased survival in a chick embryo model of extensive thoracic crush injury. Survival was not improved by mesenchymal or neural cells, or desmopressin. The authors explained that this effect was associated with the secretion of paracrine factors as evidenced by a transwell coculture model [103]. Heart We previously looked at the cardiomyogenic potential of AFSC in vitro and in vivo. AFSC cultured in cardiomyocyte induction media or in coculture with cardiomyocytes demonstrated the expression of proteins specific for cardiomyocytes (atrial natriuretic peptide and a-myosin heavy chain), endothelial (CD31 and CD144), and smooth muscle cells (a-smooth muscle actin). In our first experience with xenogenic transplantation, human AFSC were transplanted in a rat model of myocardial infarction (MI). Cells of the immune system were recruited including

142

9. STEM CELLS FROM THE AMNION

T, B, and NK cells and macrophages, and resulted in cell rejection. We speculated that this may be caused by AFSC expression of B7 costimulatory molecules CD80 and CD86 as well as macrophage marker CD68 [104]. In the next step, we attempted allogenic rat AFSC cardiac therapy by intracardiac transplantation in rats with IR injury. Three weeks after transplantation, a portion of the cells acquired an endothelial or smooth muscle phenotype and a smaller number had cardiomyocyte characteristics. Left ventricular ejection fraction was improved in the animals that had the AFSC injection, as quantified using MRI, and which suggested a paracrine therapeutic effect [82]. We then aimed to investigate this paracrine effect further using a rat model of MI and xenogeneic transplantation of human AFSC administered intravascularly immediately after reperfusion. This was dissimilar to our previous attempt with xenogeneic cellular cardiomyoplasty, which involved intramuscular injection of human AFSC within 20 min of coronary artery occlusion without reperfusion. Intravascular injection of human AFSC and their conditioned medium was cardioprotective and improved cell survival, and it decreased the infarct size from 54% to around 40% in both cases. We also showed that AFSC secrete the actin monomer-binding protein thymosin b-4 (Tb-4), a paracrine factor with cardioprotective properties [105]. Tb-4 has also been implicated as being cardioprotective in MI models that involved BM-MSC injection [106]. In addition to models of myocardial IR injury, we investigated the salutary effects of AFSC in a rat model of right heart failure resulting from pulmonary hypertension. After intravascular injection, AFSC engrafted in the lung, heart, and skeletal muscle reduced brain natriuretic peptide, a surrogate marker for heart failure, and proinflammatory cytokines. Moreover, AFSC differentiated into endothelial and vascular cells forming microvessels, capillaries, and small arteries. A 35% decrease in pulmonary arteriole thickness accompanied the injection [107]. Of relevance, in a seminal article, Rafii demonstrated that it was possible to convert human midgestation AF-derived cells directly into a stable and expandable population of vascular endothelial cells without using pluripotency factors [108]. Notably, it has been shown that ckitþ AFSC prior differentiation express early endothelial transcription factors. Moreover, in vivo, AFSC from both second and third trimesters expanded in hypoxia were able to rescue surface blood flow when locally injected in mice after chronic ischemia damage, and possessed the ability to fix carotid artery electric damage [109]. Hematopoietic System Ditadi and colleagues were the first to demonstrate the hematopoietic potential of murine and human CD117þ/ Lin AFSC [81]. In vitro, the AFSC population in both species displayed multilineage hematopoietic potential, as demonstrated by the generation of erythroid, myeloid, and lymphoid cells. In vivo, cells belonging to all hematopoietic lineages were found after primary and secondary transplantation of murine AFSC into immunocompromised hosts, thus demonstrating the long-term hematopoietic repopulating capacity of these cells. The latter results support the idea that the AF may be a source of stem cells with the potential for therapy of hematological disorders. One of the most exciting applications of AFSC in this setting is in the field of in utero transplantation (IUT). IUT has been proposed as a novel approach for the treatment of inherited hematological disorders (including thalassemia and sickle cell disease) before birth [110]. Clinical translation has been limited by competitive and immunological barriers associated with IUT of adult bone marrowederived HSC [111]. The use of AFSC for IUT could address many of these limitations. AFSC are of fetal origin, and as a result should be able to compete better against host cells compared with adult stem cells (potentially overcoming competitive barriers to engraftment). They are nonimmunogenic to the fetus at any gestational age and are also unlikely to result in maternal immunization because of the tolerogenic properties of the placenta. IUT of AFSC would involve harvesting the cells from the AF, employing in vitro gene therapy to correct the genetic defect, and transplanting back to the donor fetus. Such a combined autologous stem cellegene transfer approach would also address some of the risks associated with administering gene therapy directly to the fetus (in utero gene therapy) [111]. The possibility of performing in vitro gene transfer to harvested ASFC would allow cells to be checked for insertional mutagenesis before transplantation and would obviate the risk of germline transmission of transgenes. In proof of principle studies, Shaw and colleagues showed that IUT of autologous (isolated using ultrasound-guided amniocentesis), expanded, and transduced AFSC resulted in widespread tissue engraftment (including the hematopoietic system) in the ovine fetus [112]. We are investigating the hematopoietic potential of freshly isolated and expanded AFSC after intravenous transplantation in immunocompetent fetal mice, and have obtained stable, multilineage engraftment at neartherapeutic levels using relatively small donor cell numbers. Whether in utero stem cellegene therapy with AFSC would be therapeutic in models of hematological and other congenital disorders remains to be determined. Kidney AFSC have been shown to have plastic regenerative properties in the lung, by differentiating into different lineages according to the type of lung injury taking place in animal models of disease. AFSC injected intravascularly into

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nude mice subjected to hyperoxia-induced pulmonary injury migrated to the lung and expressed human pulmonary epithelial differentiation marker thyroid transcription factor 1 and type-II pneumocyte marker surfactant protein C. After naphthalene injury to Clara cells, AFSC expressed the Clara cellespecific 10-kDa protein [113]. Moreover, in an adult rat model of hyperoxic lung injury, treatment with human AFSC has a reparative potential through paracrine involvement in alveolarization and angiogenesis [114]. In an established nitrofen-induced rat model of lung hypoplasia, lung growth, bronchial motility, and innervation were rescued by AFSC both in vitro and in vivo, which was similarly to results observed before with retinoic acid. The beneficial effect of AFSC was probably related to the paracrine action of growth factor secretion [115]. Those results have been validated in fetal rabbit with a surgically created left diaphragmatic hernia at day 23 (term, day 32). In this model, human AFSC exert an additional effect on tracheal occlusion leading to a decrease in mean terminal bronchiole density, a measure of alveolar number surrounding the terminal bronchioles, without signs of toxicity [116]. Lung The first evidence regarding a nephrogenic potential of AFSC arose from a series of experiments involving the ex vivo growth of murine embryonic kidneys that were injected with labeled AFSC. Whereas AFSC were viable for up to 10 days’ growth, they were also shown to contribute to a number of components of the developing kidney, such as the renal vesicle and S- and C-shape bodies. In addition, the ECM and surrounding cells induced renal differentiation, with the AFSC expressing kidney markers (zona occludens-1, glial-derived neurotrophic factor, and claudin) [91]. In a mouse model of acute tubular necrosis (ATN) involving glycerol injection, luciferase-labeled injected AFSC homed to the injured kidney. This decreased creatinine and blood urea nitrogen (BUN) levels and reduced the number of damaged tubules while increasing the proliferation of tubular epithelial cells. Interestingly, AFSC injected during the acute phase of ATN (between 48 and 72 h) had no effect on creatinine and BUN levels, whereas AFSC injected into the kidney on the same day of glycerol injection resulted in no observed peaks in creatinine or BUN. The authors speculated that this may be the result of AFSC accelerating the proliferation of partially damage epithelial tubular cells while preventing apoptosis [117]. Another laboratory confirmed the protective effect of AFSC in the same mouse model of ATN while comparing their effect with MSC. In addition to results regarding the amelioration of the effect of glycerol, it was demonstrated that MSC were more efficient in inducing proliferation and AFSC were more antiapoptotic. Sedrakyan et al. used Col4a5(/) mice as a mouse model of Alport syndrome to assess the effect of AFSC in renal fibrosis [83]. Early intracardiac administration of AFSC delayed interstitial fibrosis and the progression of glomerular sclerosis and prolonged animal survival. However, AFSC were not demonstrated to differentiate into podocytes, which suggests that the positive effects to the basement membrane were again mediated, as in other model of disease, by a paracrine mechanism [102,118]. It was reported for the first time that AFSCs mixed with organoids made with murine embryonic kidney contributed to the formation of glomerular structures, differentiated into podocytes with slit diaphragms, and internalized exogenously infused bovine serum albumin, attaining unprecedented (for donor stem cells) degrees of specialization and function in vivo [119]. Intestine We looked at the effect of AFSC in a rat model of necrotizing enterocolitis (NEC) [120]. After 24 h of life, NEC rats were randomized to treatments of AFSC, BM-MSC, myoblasts (as a committed negative control), or phosphatebuffered saline (PBS) via intraperitoneal administration. NEC rats treated with AFSC showed significantly higher survival at 7 days compared with all the other groups and had an improved NEC clinical status at 96 h. MRI displayed significantly decreased peritoneal fluid accumulation (a surrogate marker for NEC grade) in the AFSCtreated rats. The improved clinical picture of the pups injected with AFSC was also evident by measurement of intestinal permeability, contraction, and motility. The clinical data were confirmed by histological analysis, demonstrating a decreased amount of villus sloughing, core separation, and venous congestion. The relationship between these therapeutic effects and the presence of AFSC in the intestine was confirmed by tracking AFSC expressing GFP. At 48 h, cells were adherent to the mesentery; at 72 h, they were found in the serosa and muscularis; and at 96 h, they were located in the villi. The low cell numbers in these locations alongside the great clinical differences among treated groups suggested a paracrine effect. Accordingly, when we performed microarray analysis we saw differences in a number of genes involved in inflammation and tissue repair, cell cycle regulation, and enterocyte differentiation. These results were corroborated by immunofluorescence analysis examining cell proliferation and apoptosis. We then sought to investigate the paracrine mechanism by which AFSC mediate their therapeutic effect and established that the number of cryptal cells expressing cyclo-oxygenase-2 (COX-2þ) inversely correlated with the

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degree of intestinal damage. COX-2þ cells were diminished in rats treated with PBS, whereas they were maintained in rats treated with AFSC. Interestingly, although the total number of COX-2þ cells in villi was similar in AFSCtreated and control animals, cryptal COX-2þ cells were significantly increased in the AFSC rats compared with both control animals and pups treated with PBS. This dependence on COX-2 was confirmed when the effect of AFSC was abolished both by selective COX-2 and nonselective COX inhibition, but it remained unaffected by a selective COX-1 inhibitor [121].

CONCLUSIONS Many stem cell populations (e.g., embryonic, adult, and fetal stem cells) as well as methods for generating pluripotent cells (e.g., nuclear reprogramming) have been described. All of them have specific advantages and disadvantages. It has yet to be established which type of stem cell represents the best candidate for cell therapy. However, although it is likely that one cell type may be better than another, depending on the clinical scenario, the discovery of easily accessible cells of fetal derivation, not burdened by ethical concerns, in the AF has the potential of expanding new horizons in regenerative medicine. In fact, amniocentesis is routinely performed for the antenatal diagnosis of genetic diseases and its safety has been established by several studies documenting an extremely low overall fetal loss rate (0.06%e0.83%) related to the procedure [122,123]. Moreover, stem cells can be obtained from AF samples without interfering with diagnostic procedures. Two stem cell populations have been isolated from the AF (i.e., AFMSCs and AFSC) and both can be used as primary (not transformed or immortalized) cells without further technical manipulations. AFMSCs exhibit typical MSC characteristics: fibroblastic-like morphology, clonogenic capacity, multilineage differentiation potential, immunosuppressive properties, expression of a mesenchymal gene expression profile, and a mesenchymal set of surface antigens. However, ahead of other MSC sources, AFMSCs are easier to isolate and have better proliferation capacities. The harvest of bone marrow remains a highly invasive and painful procedure, and the number, proliferation, and differentiation potential of these cells decline with increasing age [124,125]. Similarly, umbilical cord bloodederived MSCs exist at a low percentage and expand slowly in culture [126]. AFSC, on the other hand, represent a class of broadly stem cells with intermediate characteristics between ES and AS cells [29,127]. They express both embryonic and MSC markers, are able to differentiate into lineages representative of all embryonic germ layers, and do not form tumors after implantation in vivo. However, AFSC have been identified only recently and many questions need to be answered concerning their origin, epigenetic state, immunological reactivity, regeneration, and differentiation potential in vivo. AFSC may not differentiate as promptly as do ES cells and their lack of tumorigenesis can be argued against their pluripotency. Although further studies are needed to better understand their biological properties and define their therapeutic potential, stem cells present in the AF appear to be promising candidates for cell therapy and tissue engineering. In particular, they represent an attractive source for the treatment of perinatal disorders such as congenital malformations (e.g., congenital diaphragmatic hernia) and acquired neonatal diseases requiring tissue repair/regeneration (e.g., NEC). In a future clinical scenario, AF cells collected during a routinely performed amniocentesis could be banked, and in case of need, subsequently expanded in culture or engineered in acellular grafts [29,62]. In this way, affected children could benefit from having autologous expanded or engineered cells ready for implantation either before birth or in the neonatal period [128].

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Valproic acid confers functional pluripotency to human amniotic fluid stem cells in a transgene-free approach. Mol Ther 2012;20(10):1953e67. [86] Kim J, Lee Y, Kim H, Hwang KJ, Kwon HC, Kim SK, Cho DJ, Kang SG, You J. Human amniotic fluid-derived stem cells have characteristics of multipotent stem cells. Cell Prolif 2007;40(1):75e90. [87] Siegel N, Valli A, Fuchs C, Rosner M, Hengstschlager M. Induction of mesenchymal/epithelial marker expression in human amniotic fluid stem cells. Reprod Biomed Online 2009;19(6):838e46. [88] Tsai MS, Hwang SM, Tsai YL, Cheng FC, Lee JL, Chang YJ. Clonal amniotic fluid-derived stem cells express characteristics of both mesenchymal and neural stem cells. Biol Reprod 2006;74(3):545e51. [89] Chen WQ, Siegel N, Li L, Pollak A, Hengstschlager M, Lubec G. Variations of protein levels in human amniotic fluid stem cells CD117/2 over passages 5-25. J Proteome Res 2009;8(11):5285e95. [90] Kolambkar YM, Peister A, Soker S, Atala A, Guldberg RE. Chondrogenic differentiation of amniotic fluid-derived stem cells. J Mol Histol 2007;38(5):405e13. [91] Perin L, Giuliani S, Jin D, Sedrakyan S, Carraro G, Habibian R, Warburton D, Atala A, De Filippo RE. Renal differentiation of amniotic fluid stem cells. Cell Prolif 2007;40(6):936e48. [92] Valli A, Rosner M, Fuchs C, Siegel N, Bishop CE, Dolznig H, Madel U, Feichtinger W, Atala A, Hengstschlager M. Embryoid body formation of human amniotic fluid stem cells depends on mTOR. Oncogene 2010;29(7):966e77. [93] Chambers I, Silva J, Colby D, Nichols J, Nijmeijer B, Robertson M, Vrana J, Jones K, Grotewold L, Smith A. Nanog safeguards pluripotency and mediates germline development. Nature 2007;450(7173):1230e4. [93a] Koike M, Sakaki S, Amano Y, Kurosawa HJ. Characterization of embryoid bodies of mouse embryonic stem cells formed under various culture conditions and estimation of differentiation status of such bodies. Biosci Bioeng 2007;104(4):294e9. 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Survival analysis of conjunctival limbal grafts and amniotic membrane transplantation in eyes with total limbal stem cell deficiency. Am J Ophthalmol 2005;140(2):223e30. [97] Toselli M, Cerbai E, Rossi F, Cattaneo E. Do amniotic fluid-derived stem cells differentiate into neurons in vitro? Nat Biotechnol 2008;26(3): 269e70. author reply 270e261. [98] Grisafi D, Piccoli M, Pozzobon M, Ditadi A, Zaramella P, Chiandetti L, Zanon GF, Atala A, Zacchello F, Scarpa M, De Coppi P, Tomanin R. High transduction efficiency of human amniotic fluid stem cells mediated by adenovirus vectors. Stem Cells Dev 2008;17(5):953e62. [99] Delo DM, Olson J, Baptista PM, D’Agostino Jr RB, Atala A, Zhu JM, Soker S. Non-invasive longitudinal tracking of human amniotic fluid stem cells in the mouse heart. Stem Cells Dev 2008;17(6):1185e94. [100] Sun H, Feng K, Hu J, Soker S, Atala A, Ma PX. 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[101] Peister A, Deutsch ER, Kolambkar Y, Hutmacher DW, Guldberg RE. Amniotic fluid stem cells produce robust mineral deposits on biodegradable scaffolds. Tissue Eng Part A 2009;15(10):3129e38. [102] Piccoli M, Franzin C, Bertin E, Urbani L, Blaauw B, Repele A, Taschin E, Cenedese A, Zanon GF, Andre-Schmutz I, Rosato A, Melki J, Cavazzana-Calvo M, Pozzobon M, De Coppi P. Amniotic fluid stem cells restore the muscle cell niche in a HSA-Cre, Smn(F7/F7) mouse model. Stem Cells 2012;30(8):1675e84. [103] Prasongchean W, Bagni M, Calzarossa C, De Coppi P, Ferretti P. Amniotic fluid stem cells increase embryo survival following injury. Stem Cells Dev 2012;21(5):675e88. [104] Chiavegato A, Bollini S, Pozzobon M, Callegari A, Gasparotto L, Taiani J, Piccoli M, Lenzini E, Gerosa G, Vendramin I, Cozzi E, Angelini A, Iop L, Zanon GF, Atala A, De Coppi P, Sartore S. Human amniotic fluid-derived stem cells are rejected after transplantation in the myocardium of normal, ischemic, immuno-suppressed or immuno-deficient rat. J Mol Cell Cardiol 2007;42(4):746e59. [105] Bollini S, Cheung KK, Riegler J, Dong X, Smart N, Ghionzoli M, Loukogeorgakis SP, Maghsoudlou P, Dube KN, Riley PR, Lythgoe MF, De Coppi P. Amniotic fluid stem cells are cardioprotective following acute myocardial infarction. Stem Cells Dev 2011;20(11):1985e94. [106] Bao W, Ballard VL, Needle S, Hoang B, Lenhard SC, Tunstead JR, Jucker BM, Willette RN, Pipes GT. Cardioprotection by systemic dosing of thymosin beta four following ischemic myocardial injury. Front Pharmacol 2013;4:149. [107] Castellani C, Vescovo G, Ravara B, Franzin C, Pozzobon M, Tavano R, Gorza L, Papini E, Vettor R, De Coppi P, Thiene G, Angelini A. The contribution of stem cell therapy to skeletal muscle remodeling in heart failure. Int J Cardiol 2013;168(3):2014e21. [108] Ginsberg M, James D, Ding BS, Nolan D, Geng F, Butler JM, Schachterle W, Pulijaal VR, Mathew S, Chasen ST, Xiang J, Rosenwaks Z, Shido K, Elemento O, Rabbany SY, Rafii S. Efficient direct reprogramming of mature amniotic cells into endothelial cells by ETS factors and TGFbeta suppression. Cell 2012;151(3):559e75. [109] Schiavo AA, Franzin C, Albiero M, Piccoli M, Spiro G, Bertin E, Urbani L, Visentin S, Cosmi E, Fadini GP, De Coppi P, Pozzobon M. Endothelial properties of third-trimester amniotic fluid stem cells cultured in hypoxia. Stem Cell Res Ther 2015;6:209. [110] Ramachandra DL, Shaw SS, Shangaris P, Loukogeorgakis S, Guillot PV, Coppi PD, David AL. In utero therapy for congenital disorders using amniotic fluid stem cells. Front Pharmacol 2014;5:270. [111] Loukogeorgakis SP, Flake AW. In utero stem cell and gene therapy: current status and future perspectives. Eur J Pediatr Surg 2014;24(3): 237e45. [112] Shaw SW, Bollini S, Nader KA, Gastaldello A, Mehta V, Filppi E, Cananzi M, Gaspar HB, Qasim W, De Coppi P, David AL. Autologous transplantation of amniotic fluid-derived mesenchymal stem cells into sheep fetuses. Cell Transplant 2011;20(7):1015e31. [113] Carraro G, Perin L, Sedrakyan S, Giuliani S, Tiozzo C, Lee J, Turcatel G, De Langhe SP, Driscoll B, Bellusci S, Minoo P, Atala A, De Filippo RE, Warburton D. Human amniotic fluid stem cells can integrate and differentiate into epithelial lung lineages. Stem Cells 2008;26(11):2902e11. [114] Grisafi D, Pozzobon M, Dedja A, Vanzo V, Tomanin R, Porzionato A, Macchi V, Salmaso R, Scarpa M, Cozzi E, Fassina A, Navaglia F, Maran C, Onisto M, Caenazzo L, De Coppi P, De Caro R, Chiandetti L, Zaramella P. Human amniotic fluid stem cells protect rat lungs exposed to moderate hyperoxia. Pediatr Pulmonol 2013;48(11):1070e80. [115] Pederiva F, Ghionzoli M, Pierro A, De Coppi P, Tovar JA. Amniotic fluid stem cells rescue both in vitro and in vivo growth, innervation, and motility in nitrofen-exposed hypoplastic rat lungs through paracrine effects. Cell Transplant 2013;22(9):1683e94. [116] DeKoninck P, Toelen J, Roubliova X, Carter S, Pozzobon M, Russo FM, Richter J, Vandersloten PJ, Verbeken E, De Coppi P, Deprest J. The use of human amniotic fluid stem cells as an adjunct to promote pulmonary development in a rabbit model for congenital diaphragmatic hernia. Prenat Diagn 2015;35(9):833e40. [117] Perin L, Sedrakyan S, Giuliani S, Da Sacco S, Carraro G, Shiri L, Lemley KV, Rosol M, Wu S, Atala A, Warburton D, De Filippo RE. Protective effect of human amniotic fluid stem cells in an immunodeficient mouse model of acute tubular necrosis. PLoS One 2010;5(2):e9357. [118] Rota C, Imberti B, Pozzobon M, Piccoli M, De Coppi P, Atala A, Gagliardini E, Xinaris C, Benedetti V, Fabricio AS, Squarcina E, Abbate M, Benigni A, Remuzzi G, Morigi M. Human amniotic fluid stem cell preconditioning improves their regenerative potential. Stem Cells Dev 2012;21(11):1911e23. [119] Xinaris C, Benedetti V, Novelli R, Abbate M, Rizzo P, Conti S, Tomasoni S, Corna D, Pozzobon M, Cavallotti D, Yokoo T, Morigi M, Benigni A, Remuzzi G. Functional human podocytes generated in organoids from amniotic fluid stem cells. J Am Soc Nephrol 2016;27(5):1400e11. [120] Zani A, Cananzi M, Lauriti G, Fascetti-Leon F, Wells J, Siow B, Lythgoe MF, Pierro A, Eaton S, De Coppi P. Amniotic fluid stem cells prevent development of ascites in a neonatal rat model of necrotizing enterocolitis. Eur J Pediatr Surg 2014;24(1):57e60. [121] Zani A, Cananzi M, Fascetti-Leon F, Lauriti G, Smith VV, Bollini S, Ghionzoli M, D’Arrigo A, Pozzobon M, Piccoli M, Hicks A, Wells J, Siow B, Sebire NJ, Bishop C, Leon A, Atala A, Lythgoe MF, Pierro A, Eaton S, De Coppi P. Amniotic fluid stem cells improve survival and enhance repair of damaged intestine in necrotising enterocolitis via a COX-2 dependent mechanism. Gut 2014;63(2):300e9. [122] Caughey AB, Hopkins LM, Norton ME. Chorionic villus sampling compared with amniocentesis and the difference in the rate of pregnancy loss. Obstet Gynecol 2006;108(3 Pt 1):612e6. [123] Eddleman KA, Malone FD, Sullivan L, Dukes K, Berkowitz RL, Kharbutli Y, Porter TF, Luthy DA, Comstock CH, Saade GR, Klugman S, Dugoff L, Craigo SD, Timor-Tritsch IE, Carr SR, Wolfe HM, D’Alton ME. Pregnancy loss rates after midtrimester amniocentesis. Obstet Gynecol 2006;108(5):1067e72. [124] D’Ippolito G, Schiller PC, Ricordi C, Roos BA, Howard GA. Age-related osteogenic potential of mesenchymal stromal stem cells from human vertebral bone marrow. J Bone Miner Res 1999;14(7):1115e22. [125] Kern S, Eichler H, Stoeve J, Kluter H, Bieback K. 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C H A P T E R

10 Cord Blood Stem Cells Kristin M. Page, Jessica M. Sun, Joanne Kurtzberg Duke University, Durham, NC, United States

INTRODUCTION Umbilical cord blood (CB) is firmly established as an unrelated donor source for hematopoietic stem cell transplantation (HSCT) and is a readily available cell source in the evolving fields of regenerative medicine and cellular therapies. Worldwide, there are over 160 public banks with a global inventory of over 700,000 fully characterized, high-quality cord blood units (CBUs) [1], and more than 5 million CBUs have been banked at an estimated 215 family banks. In this chapter, we review the history of CB transplantation (CBT) and banking as well as established and emerging clinical uses of CB in regenerative medicine.

A BRIEF HISTORY CB was first recognized as rich source of hematopoietic stem and progenitor cells (HSCs) several decades ago. In a pivotal series of experiments, Dr. Ted Boyce, working with Dr. Hal Broxmeyer and colleagues, demonstrated that CB HSCs possessed high proliferative potential, could successfully repopulate hematopoiesis in murine models and tolerated cryopreservation and thawing with efficient HSC recovery [2]. This critical work provided the scientific rationale to investigate CB as a potential source of donor HSCs in humans. The first patient to undergo a CBT was a 5-year-old boy with Fanconi anemia. Through prenatal testing, it was determined that his mother was pregnant with an unaffected human leukocyte antigen (HLA) matched sibling. Upon the sibling’s birth, the CB was collected into a sterile bottle containing preservative-free heparin and transported to Dr. Broxmeyer’s laboratory, where it was diluted with tissue culture media and dimethyl sulfoxide (DMSO). The CB was then cryopreserved and stored under liquid nitrogen until it was transported to Paris, France, where the transplant would occur. The clinicians caring for the family elected to wait until the healthy sibling donor was 6 months of age so that she could serve as a backup bone marrow donor if the cord blood transplant failed. In 1988, Dr. Eliane Gluckman performed the first CBT in the world using the sibling’s CB as the donor [3]. The child successfully engrafted with his sister’s cells and has remained healthy with full donor chimerism ever since. Building on this initial success, additional related donor CBTs were performed in selected centers over the next 5 years [4e7]; the first unrelated donor CB bank was established by Dr. Pablo Rubinstein at the New York Blood Center in 1992. In the following year, Dr. Joanne Kurtzberg performed the first unrelated donor CBT at Duke University in a 4-year-old child with relapsed T-cell leukemia. Early experience with unrelated CBT demonstrated that partially HLA-mismatched, banked unrelated donor CB could successfully restore hematopoiesis with an incidence of graft versus host disease (GvHD) lower than expected and that engraftment was associated with the total nucleated cell (TNC) dose available relative to the recipient’s body size [8,9]. With the extension into the unrelated donor setting, the fields of CBT and banking expanded rapidly. In 1995, EUROCORD was established by Dr. Eliane Gluckman; it continues to operate as an international CBT registry on behalf of the European Group for Blood and Marrow Transplantation. In 1996, the Foundation for the Accreditation of Cellular Therapy (FACT) was established by its parent organizations, the International Society for Cellular Therapy and the American Society of Blood and Marrow Transplantation (ASBMT). The following year, the International NetCord Foundation was established to serve as a registry for international public banks. The first Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00010-2

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international public CB banking standards were created by NetCord and the first accreditation program resulted from a collaboration between FACT and NetCord in 1999. Although participation is these programs is considered voluntary, many public CB banks are required to received accreditation from FACT/NetCord or the American Association of Blood Banks to participate in registries and receive reimbursement for units distributed for transplantation. In the United States, the National Marrow Donor Program (NMDP) established the Center for Cord Blood in 1999 and expanded the unrelated donor registry to include publically banked CBUs. In 2006, after passage of the Stem Cell Act of 2005, Congress established the CW Bill Young Cell Transplantation Program. This program awarded contracts to the NMDP to create a single point of access donor registry and coordinating centers for CB and adult donors, to the Center for International Blood and Marrow Transplant Research to develop a stem cell outcomes database, and to the Health Resources and Services Administration of the Department of Health and Human Services to administer these programs and establish the national cord blood inventory (NCBI), a US network of public CB banks. This program was reauthorized in 2010 and 2015. A major goal of the program is to create a large repository of high-quality CBUs from donors of diverse ancestry to enable access to HSCT for as many patients as possible. In the United States, the Food and Drug Administration (FDA) regulates unrelated donor CB as biological product and issued final guidance for public banks to obtain a Biological License Agreement (BLA) in 2011. Seven public banks in the United States have obtained a BLA.

CORD BLOOD BANKING Historically, after the birth of a baby, the CB and placenta were discarded as medical waste. With the advent of CBT, methods to reliably collect, process, test, cryopreserve, and store CB were developed with the goal of creating banks that could provide ready access to safe CBU products suitable for transplantation. Since the establishment of the first unrelated donor CB bank, the banking industry exploded, with an estimated worldwide inventory of more than 6 million CBUs stored in a combination of public and family CB banks [10]. Although the “shelf life” of banked cord blood is not known, successful transplants have been performed with units stored as long as 25 years. As the field has grown and advanced, accrediting agencies have emerged and developed standards for quality banking practices. In the United States, the FDA regulates unrelated CB as a biological product.

PUBLIC VERSUS FAMILY (OR PRIVATE) BANKS There are two main types of CB banks. Public banks collect, process, and store donated CBUs intended for unrelated allogeneic transplantation at no expense to the donor or their family. Mothers consent to donate their baby’s CB for public use and thereby relinquish all future rights to the unit. In the United States and most other countries, public banks are subject to strict regulatory oversight and use stringent volume, cell count, sterility, donor eligibility, and medical history specifications to determine which collected CBUs are banked. Because cell count is a critical determinant of CBU use, only larger units containing sufficient cells for a single or double CBT of an adultsized individual are banked. As a result, many donated units do not qualify to be banked and are discarded or used for research, depending on the consent given by the donor family. Family (or private) banks are generally “for-profit” businesses that charge the parents an initial processing fee and an annual storage fee to store CB exclusively for use by the child or the family. In actuality, the likelihood of using a privately banked CB for transplantation is low [11]. Therefore, the American Academy of Pediatrics, American Congress of Obstetricians and Gynecologists (ACOG), and ASBMT and other similar organizations worldwide do not recommend banking CB for personal use unless there is a family history of a disease (e.g., malignancy or hemoglobinopathy) that is amenable to HSCT [11e13]. To facilitate banking in these instances, many public and family banks offer “directed donor” programs, some of which waive charges associated with the banking process. Family banks are not subjected to the same regulatory oversight as public banks, although this varies among countries. Family banks generally use less stringent criteria for banking, which leads to wide variations in volume and cell content of the private inventory. According to a study of the safety of autologous CB infusions to treat children with acquired brain injury, CBUs from family banks were inferior to those stored in public banks with respect to collection volume, total nucleated cell count (TNCC), and CD34þ count (Fig. 10.1) [14]. Because the use of CB is likely to expand, particularly in the field of regenerative medicine, the

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FIGURE 10.1 Distributions of quality variables. In (AeC), the distribution of autologous CBUs is compared with the entire Carolinas Cord Blood inventory with respect to collection volume (A), TNC (B), and CD34 content (C). In (D), TNC of autologous CBUs (represented as red circles) and National Cord Blood Inventoryeeligible Carolinas Cord Blood Bank CBUs (represented as blue squares) are compared. CBUs, cord blood units; TNC, total nucleated cell. Used with permission from Sun J, Allison J, McLaughlin C, et al. Differences in quality between privately and publicly banked umbilical cord blood units: a pilot study of autologous cord blood infusion in children with acquired neurologic disorders. Transfusion 2010;50(9):1980e7.

indications and criteria that a CBU must meet for use may change. In response to these changes, the role of regulatory oversight in family CB banking may need to be optimized.

PUBLIC CORD BLOOD BANKING PROCEDURES Donor Recruitment and Consent Mothers who are potentially eligible for donation are identified based on clinical characteristics defined by the individual bank. At the Carolinas Cord Blood Bank (CCBB), donations are accepted from healthy mothers (aged 18 years) who are carrying healthy singleton pregnancies of at least 34 weeks’ gestation and who provide written informed consent for donation before collection. Consent gives permission to collect the CBU. For potentially eligible units, the mother is approached again to give written informed consent for banking the unit for use in transplantation. These mothers provide a medical and family history, a maternal blood specimen to screen for certain communicable diseases, and a review of medical records of the infant and maternal donors. The mother also gives permission for the CB to be used for research if it does not meet specifications for banking.

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Collection Techniques CB can be collected from either vaginal or cesarean births, either before delivery of the placenta (in utero) by obstetrical staff or after delivery of the placenta (ex utero), allowing for dedicated, trained bank staff to perform collections. Although the delivery method is dictated by the clinical status of the mother and infant, the method of collection is determined by staffing and collection site practices. Generally, reports have observed higher collection volumes after cesarean compared with vaginal deliveries [15e17] and when CB is collected in utero compared with ex utero [18], although not all reports have agreed [19]. Both collection methods continue to be used routinely, but in utero collections are more common, likely owing to the additional personnel expenses associated with ex utero collections. Regardless of the collection method, CB is typically collected by cannulating the umbilical vein, allowing the placental blood to drain by gravity into collection containers with anticoagulant, most commonly citrate phosphate dextrose. Closed system collection bags are standard, because they have been shown to reduce bacterial contamination [20].

Volume and Cell Count Considerations The goal of public CB banks is to provide quality CBUs for allogeneic transplantation for all patients in need. Because successful CBT requires a minimum TNC per kilogram dose, thresholds for banking have been established based on the TNCC. It is well-established that collection volume and TNCC are closely correlated. Therefore, many banks weigh the CBU after collection to estimate the volume and then ship to the processing laboratory only units that surpass a specified volume threshold. Low-volume units are discarded at the collection site because they are unlikely to have sufficient TNCC. Alternatively, some banks measure TNCC at the collection site or use other criteria to determine which units proceed to processing. Efforts to increase collection volume have focused on two general approaches: identifying donations likely to have higher collection volume and developing techniques to obtain the maximal volume from an individual donation. Increased donor birth weight and older gestational age have been closely associated with higher collection volume and TNCC [21e23], although data suggest that collections from younger infants (aged 34e37 weeks’ gestation) are more likely to have higher progenitor cell content, as measured by CD34þ and colony-forming unit (CFU) content (Fig. 10.2) [24]. Also, despite comparable collection volumes among donors of different races or ethnicities, the TNCC, CD34þ, and CFU content adjusted for collection volume (counts/mL) are lower in individuals of certain racial backgrounds, particularly African Americans, compared with Caucasian donors, even after adjusting for other

42.0 40.0

Overall p CDHA > HA where MCPM is monocalcium phosphate monohydrate, TTCP is tetracalcium phosphate, a-TCP is a-tricalcium phosphate, DCPD is dicalcium phosphate dehydrate, DCPA is dicalcium phosphate anhydrous, OCP is octacalcium phosphate, and CDHA is calcium-deficient HA. By exploiting these solubility patterns, researchers are able to tailor and modify the biodegradation rate of these CaP bioceramics to better suit the rate of new bone formation. For example, biphasic CaP compounds composed of HA and b-TCP were designed by incorporating b-TCP into HA ceramics. As a consequence, these biphasic bioceramics exhibited a greater rate of resorption than pure HA ceramics [18]. More specifications regarding the biodegradability and solubility trends of CaP compounds can be reviewed elsewhere [12,19]. In the 1970s and 1980s, research into CaP bioceramics focused on their development for dental and orthopedic applications in the form of dense or porous blocks and granules [15,19,20]. Some of the most common dental applications include alveolar ridge augmentation, fillers for periodontal bony defects, tooth root replacement surgery, and coatings for metal screws for dental implants. Typical orthopedic applications included repair of bony defects as well as coatings for orthopedic metal implants usually used for hip and knee replacement surgery [19]. Although these more traditional CaP bioceramics have shown some clinical success, they still lack some important properties that severely restrict their clinical application, in particular with respect to their handling properties. CaP blocks or granules usually require invasive surgery to reach the desired implant site and achieve complete filling of the bone defect. Furthermore, because granules have poor cohesion, they run the risk of migrating out of the implant site, which can lead to unpredictable bone growth or other complications such as blood clotting [21]. From a surgeon’s point of view, trying to shape granules or blocks to fit into an irregularly shaped bony defect site can prove challenging and lead to implant migration [22]. Therefore, these forms of CaP compounds have proven to be of limited value when the defect site is not easily accessible or when microinvasive percutaneous surgery would be a preferred method for implantation. Moreover, CaP in bulk form is considerably brittle in nature and exhibits poor mechanical strength and fracture toughness, which limits its uses to nonloading bearing sites [11]. These complications, coupled with the fact that most surgical trends lean toward the development of minimally invasive surgical techniques, have warranted the development of a bone substitute material that exhibits superior handling properties and can be injectable or moldable [23]. These demands have resulted in the development of calcium phosphate cements (CPCs).

Calcium Phosphate Cements A current challenge for surgeons is to be able to place a bone substitute material directly into the defect site using the least invasive method possible. With the development of more advanced minimally invasive surgical procedures such as percutaneous surgery, there has been a demand for researchers to develop bone substitute materials that can be injectable, such as CPCs [11,24]. CPCs are a synthetic, self-setting bone substitute material that was first discovered by Brown and Chow in the early 1980s [19,25]. This “first-generation” cement consisted of an equimolar amount of TTCP and DCPA mixed with water in a liquid-to-powder ratio (LPR) of 1:4 to form a paste. This paste differed from the more traditional granule and block forms of CaPs in the sense that upon setting and hardening in vivo, the paste would sustain a rapid phase transformation resulting in the formation of HA or brushite with no hazardous acidic or basic by-products [25,26]. Details regarding this setting mechanism and the final products of the set cements are discussed further on in this chapter. Basic Properties The discovery of CPCs has opened up a new era for bone substitute materials in which handling properties are of upmost importance. This improvement in handling properties allowed for CPCs to be employed in minimally invasive surgery in which it also exhibited the unique ability to be molded and shaped to fill abnormal and complex defect sites where they were then able to set in situ [27]. In addition, because CPCs set in an aqueous environment at 37 C, their CDHA or brushite end product is more similar to that of natural apatites rather than CaP granules and blocks, which are sintered at high temperatures. This low-temperature formation also means that organic molecules such as polymers and living cells could be safely incorporated into the cement, which could further improve their biological response [28].

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Similar to all CaP compounds, CPCs possess excellent biological properties owing to their chemical resemblance to the mineral component of bone [29]. As a consequence, several studies have shown that CPCs are highly biocompatible and osteoconductive materials with the ability to stimulate new bone growth by developing osteoconductive pathways [30]. Interestingly, it was shown that CPCs create chemical bonds to the surrounding host bone, which makes them suitable for fixating metal devices in mechanically weak bone. In addition, studies have shown that certain CPC formulations can even exhibit antimicrobial behavior as well as support osteoblast cell adhesion and gene expression in vitro [29]. CPCs also require several important properties to be employed to the best of their ability. For example, the cement requires suitable rheological properties with respect to viscosity and cohesion to ensure good bonding of the CaP particles to each other while providing adequate cell permeability. Conversely, CaP blocks and granules tend to be denser with minimal inherent porosity, which restricts their biodegradation and new bone formation. Good cohesion of CPCs is also needed to confine the CaP particles to the implant site and avoid migration throughout the body, which could lead to potential complications. CPCs also need to exhibit a suitable setting time with minimal heat output during the setting reaction to be relevant for clinical applications. Finally, the rate of resorption for the cements needs to be tailored to the rate of new bone formation so that the cement can provide adequate mechanical strength until it is replaced by new bone tissue [11,31]. CPCs are a rising class of bone substitute materials that have garnered much attention in the bone tissue engineering field. They are widely considered to be an ideal material for many bone repair and augmentation applications owing to their unique combination of biocompatibility, osteotransductivity, injectability, moldability and manipulation into bony defects, and self-setting ability in situ without producing toxic by-products. In this regard, CPCs are a more attractive material than prefabricated CaP blocks or granules for bone replacement applications [29].

CLASSES OF CALCIUM PHOSPHATE CEMENTS In general, CPCs can be categorized into two main types based on the formed end products, i.e., apatite or brushite. The end product of the set cement is mainly determined by the solubility of the CaP precursor compounds and the pH of the setting reaction. Generally, a poorly crystalline HA or CDHA (apatite) forms when the pH value is greater than 4.2 whereas brushite, otherwise known as DCPD, forms when the pH value is less than 4.2 [32]. With respect to solubility characteristics, brushite-forming cements have a higher solubility rate than apatiteforming cements. Therefore, brushite cements tend to resorb faster than apatite cements both in vitro and in vivo [14,26].

Apatite Cements All apatite-forming CPCs undergo a phase transformation that results in the formation of a poorly crystalline precipitated HA and/or CDHA as the final end product [32]. This end product has a strong similarity to the mineral phase in bone and teeth, which is primarily attributed to the fact that this type of cement is formed in an aqueous environment, resulting in a poorly crystalline structure. In some instances, a full phase transformation does not always occur and therefore small traces of unreacted CaP precursor compounds may remain in the matrix [29]. The most common CaP precursor compounds for apatite-forming cements include tricalcium phosphate (TCP) and/or TTCP [26]. For TCP-based cements, the particle size, degree of crystallinity, and crystal phase all strongly influence its degree of reactivity. For example, higher degrees of reactivity have been reported when the crystal phase was thermodynamically less stable. To this end, amorphous calcium phosphate (ACP) is considered to be the most reactive because it has the least stable crystal phase, followed by a-TCP and finally b-TCP [33]. Furthermore, the smaller the particle size, the more reactive the compound is, owing to the higher surface area available for reaction with the aqueous environment [32,34]. Interestingly, for apatite-forming cements, the forces holding the newly formed, highly interconnected CDHA crystals together are relatively weak, which means that detachment of the crystals in the network from each other can easily occur, especially once the passive dissolution process has been initiated. This is considered to be advantageous from biological and biodegradation points of view, because osteoclasts are better suited to ingest the smaller, fragmented apatite crystals [35]. The solubility of these cements also has an important role in regulating the biodegradation of the cements in vivo. For example, because apatite-forming CPCs and bone share a similar biological composition, they also exhibit similar solubility characteristics. Thus, like normal bone mineral, these CPCs are

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relatively insoluble in aqueous solutions at a neutral pH but become more soluble as the pH becomes more acidic. This is a unique feature that allows osteoclasts to facilitate controlled dissolution during the bone remodeling process [36]. Although apatite cement degrades faster than stoichiometric HA, the rate of degradation is still relatively slow compared with brushite cements, resulting in degradation times that may take several years to decades [37].

Brushite Cements As previously noted, all brushite-forming CPCs precipitate as DCPD as its final cementitious end product, in which the pH value of the setting reaction in the beginning is below 4.2. A wide variety of formulations have been proposed within the literature, but formulations of b-TCP þ MCPM [38], and b-TCP þ H3PO4 [29,39] tend to be most commonly used. Contrary to apatite cements, all brushite cements can form under only one type of setting reaction: the acidebase interaction, as described in more detail subsequently in this chapter. This characteristic is attributed to the chemical composition of brushite cement formulations and also the fact that DCPD can be precipitated only in a solution with a pH less than 6. Therefore, it is also understood that all formulations of this type of cement are acidic by nature during the setting and hardening phase [12,39]. For example, the pH of cements formed by a reaction between MCPM and b-TCP is initially low (w2.5) before the pH eventually begins to rise to around 6 as the hardening phase progresses [39]. By modifying the starting formulations, the pH of the reaction, and thus several important properties of the cement, can be controlled. For example, replacing MCPM with H3PO4 in b-TCPebased formulations maintains a highly acidic cement paste for the first approximately 30 s of the setting reaction. Subsequently, the pH begins to rise in a similar fashion, as observed in MCPM þ b-TCP formulations. Several major advantages can be obtained when MCPM is replaced by H3PO4 in b-TCP-based formulations, most notably: (1) more control over the reaction and chemical composition; (2) more homogeneity in the cement matrix, resulting in higher tensile strengths; and (3) prolonged setting times, which is desired specifically for brushite cements for which setting times are generally too fast [39]. The setting time of brushite cements is determined primarily by the solubility of the alkaline phase present within the CaP formulation. The more soluble the formulation, the faster the setting time is. Furthermore, a reduction in the solubility can be accomplished by increasing the alkalinity of the formulation, effectively prolonging the setting time. It is well-known that the alkalinity of HA is greater than b-TCP, which in turn is greater than a-TCP. In other words, the solubility of these three CaP compounds is: HA < b-TCP < a-TCP [14,40]. Therefore, by adding different types CaP compounds with varying pH and solubility characteristics to the formulation, the setting time can be modified. MCPM-based formulations, for instance, have been mixed with HA, b-TCP, or a-TCP. It was shown that the formulation containing HA took several minutes to set, whereas the formulation containing b-TCP took only 30e60 s to set, and finally the setting time for the a-TCPecontaining formulation was further reduced to a matter of only several seconds [38]. Alternatively, incorporating additives into the cements that act by inhibiting the precipitation and growth of DCPD crystals has proven to be a successful strategy for increasing the setting time [41]. This effect has been observed for cements when setting retardants such as glycolic acid [42], sodium citrate, and citric acid were incorporated into the cement formulations [29]. Brushite cements set within a relatively short period of time from a liquid state, while these cements are highly bioresorbable and biocompatible. This latter feature is attributed to the fact that DCPD has a higher solubility and better stability under physiological conditions compared with apatite [12]. Therefore, it has been shown that brushite cements have a faster in vivo biodegradation rate than apatite cements. This rapid resorption rate also means that these cements have a rapid and significant decrease in mechanical strength after in vivo implantation [32]. Moreover, the setting times of brushite cements are generally so rapid that they require a large volume of liquid phase to keep the cement paste injectable and workable over a reasonable time. This means that high LPRs are often employed for these cements and this higher liquid content results in a more porous, inherently weaker cement as the final end product [11]. Poor mechanical strength coupled with short setting times and constricted handling properties have limited the clinical applicability of brushite cements, in particular with respect to load-bearing applications [29,43].

PHYSIOCHEMICAL PROPERTIES Setting/Hardening Mechanism The setting of CPCs is a continuous process that always begins with the dissolution of CaP precursor compounds into the liquid phase [17,26]. Generally, all CPCs are composed of two main phases: a solid/powder phase and a

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liquid phase. Generally, the solid phase consists of one or more of CaP compounds listed in Table 34.1, whereas the liquid phase is typically either distilled water or an aqueous solution containing, e.g., phosphate-buffered saline, sodium phosphate, ammonium phosphate, citric acid, or sodium silicate [29]. Once the solid and liquid phases are combined and thoroughly mixed, dissolution of the CaP precursor compounds into the aqueous environment occurs. The rate of this dissolution highly depends on the pH of the liquid phase as well as the chemical composition of the CaP precursor compounds. Once the Ca2þ and PO4 3 ions dissolve in solution, a supersaturation level is reached. Subsequently, these reactants begin to precipitate back onto the surface of the remaining powder constituents, usually in the form of CDHA (apatite) or DCPD (brushite) [32,44]. What is left is a highly viscous and moldable self-setting paste that can be placed directly into a bone defect site to aid in the healing process [11]. Generally speaking, once the cement is placed within the body, most of the phase transformation occurs within the first 6 h of setting; approximately 80% of the CaP precursor compounds convert to the final end products in this time frame [29]. An important factor behind the setting reaction of CPCs is the relative stability and solubility of the CaP compounds. This solubility is why a chemical transformation is induced once the CaP compound(s) are mixed with an aqueous solution. During the precipitation phase of the setting reaction, the newly formed CDHA or DCPD crystals grow epitaxially in the form of needle or rod-like microstructures that become highly interconnected and entangled. This highly entangled network of crystals provides structural rigidity to the hardened CPCs. Moreover, this highly interconnected web of crystals means that the microstructure and nanostructure of the cements possess an inherently intrinsic porosity usually between 40% and 60%, with pore sizes ranging between about 0.1 and 10 mm [11,29]. This formation of a web of needle or rod-like crystals during the setting reaction, as described previously, also leads to the adherence of crystalline grains, which results in hardening of the cement. After a sufficient time has passed, usually at least 24 h, most of the crystals have completely formed; a cement matrix is left composed of some areas of highly dense, compacted crystals whereas other areas are highly porous owing to large separations of the crystals [29]. For some apatite-forming CPCs, water is generally not employed as a reactant but rather acts as a medium. Because of this, the amount of water needed for setting to occur is minimal. Conversely, for brushite-forming CPCs, water always actively participates in the setting reaction because it is necessary for DCPD to form. Therefore, these cements are commonly referred to as hydraulic cements [30,44]. Chemical Reaction The chemical reaction that takes place during the setting of CPCs depends on several key variables such as the pH of the reaction, the LPR, the particle size, and, most importantly, the chemical composition of the CaP precursor compounds. Two main types of setting reactions can be discerned: the acidebase interaction and hydrolysis [14,26,32]. AcideBase Interaction The first type of setting reaction, the acidebase interaction, employs the use of the general principles of chemistry in which, once an acidic compound and a basic compound are combined, they interact and produce a neutral compound [29]. In this instance, a relatively acidic CaP compound reacts with a relatively basic one to produce a neutral end product. A typical example of this reaction can be observed in what is considered the first CPC formulation developed by Brown and Chow in the 1980s [25]. This cement formulation consisted of acidic DCPA and basic TTCP compounds that, once combined in an aqueous liquid phase, reacted to form a relatively neutral, poorly crystalline precipitated HA, as shown in Eq. (34.1) [25,29]: 2Ca4(PO4)2O þ 2CaHPO4 / Ca10(PO4)6(OH)2 (34.1) It was initially thought that the reaction of DCPA with TTCP would result in the formation of stoichiometric HA (Ca/P ionic ratio ¼ 1.67), but upon further investigation it was determined that only the first nuclei were stoichiometric HA. As the reaction continued, the nuclei would continue to grow and enlarge in the form of CDHA (Ca/P molar ratio of 1.5e1.67) crystals [25,29]. Hydrolysis Interaction The second type of setting reaction involves the hydrolysis of CaP compounds in a liquid phase. This hydration process is mildly exothermic and is composed of five stages: (1) initiation, (2) induction, (3) the acceleration, (4) deceleration, and (5) termination [29]. Unlike the previous acidebase interaction reaction, this reaction employs the use of only one CaP precursor compound, and therefore the Ca/P molar ratio remains the same from the beginning to the end of the reaction [45]. Typical CaP precursor compounds include ACP, DCPA, CDHA, OCP, TTCP,

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b-TCP, and a-TCP; a-TCP is the most commonly used in the literature. Once one of these single-phase CaP compounds is mixed with an aqueous solution, most reprecipitate into CDHA crystals during the setting of the cement. An example of the chemical reaction using a-TCP as the starting compound and water as the aqueous solution can be seen in Eq. (34.2) [29]: 3aCa3(PO4)2 þ H2O / Ca9(HPO4) (PO4)5(OH) (34.2) For the hydrolysis of a-TCP, the literature indicates that the reaction is controlled through the dissolution of Ca2þ and PO4 3 ions at the surface of the a-TCP particles. Therefore, the surface area, or particle size, has a crucial role in the setting process as well as the time it takes for the cement to set fully [46,47]. Results from another study indicated that the incorporation of 2 wt% CDHA into the a-TCP powder phase to act as a nucleation sites accelerated the kinetics of the setting reaction [46]. A comparison of these two setting mechanisms, with examples from the most common CaP formulations for both apatite and brushite cements, is shown in Fig. 34.2. Setting Times Setting time properties are greatly important for all CPC formulations. This is especially true from the perspective of the surgeon, as the setting time dictates the amount of time the surgeon can mold and/or inject the cement as well

FIGURE 34.2 Classification of calcium phosphate cement setting mechanisms with examples from the most common formulations. Cements are classified based on their final phase transformation end product (apatite or brushite), the number of components in the powder phase (single or multiple), the type of setting reaction (hydrolysis interaction or acidebase interaction), and the microstructure progression during the setting reaction. CDHA, calcium-deficient hydroxyapatite; DCP, dicalcium phosphate; DCPA/DCPD, dicalcium phosphate anhydrous/dicalcium phosphate dehydrate; MCP, monocalcium phosphate; MCPM/MCPA, monocalcium phosphate monohydrate/monocalcium phosphate anhydrous; SEM, scanning electron microscopy; TTCP, tetracalcium phosphate; a-TCP, a-tricalcium phosphate; b-TCP, b-tricalcium phosphate. Reprinted with permission from Ginebra M-P, Canal C, Espanol M, Pastorino D, Montufar EB. Calcium phosphate cements as drug delivery materials. Adv Drug Deliv Rev 2012;64(12):1090e1110.

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as close the wound safely in the operating theater. Setting times must be slow enough to give the surgeon time to implant and mold the cement into the defect site, and fast enough not to delay the operation. There are two main methods for how researchers can measure the setting time of the cements: the Gillmore needles method (ASTM C266-89) and the Vicat needle method (ASTM C191-92) [29]. Both involve placing a weighted needle onto the surface of the cement at different time points to visualize whether an indentation was created. Once a full indentation has not been created on the surface, the cement is considered to be set. More specifically, Gillmore needles have been used to measure the initial and final setting times of cements [48]. For instance, a lighter, thicker needle is used to determine the initial setting time whereas a heavier, thinner needle is used to determine the final setting time [49]. Both the Gilmore needle and Vicat needle methods are highly user-dependent because their results are based on the concept of “visible indentation,” which can make reproducibility of data between research groups challenging. Moreover, these methods provide no information about the progression of the setting reaction that controls the setting process and time. Therefore, other methods such as impedance spectroscopy have garnered much attention as a way to monitor the setting reactions of CPCs in situ. More information regarding this technique can be found elsewhere [50]. With respect to its clinical use, the initial setting time translates to the amount of time the surgeon has to implant and work with the cement, whereas the final setting time translates to the time when the wound can be closed, as depicted in Fig. 34.3. The cement should not move or deform between the initial and final setting time because this could induce cracks. Therefore, setting time criteria have been defined as follows: 3 min  I < 8 min I e CT  1 min F  15 min where CT ¼ cohesion time, I ¼ initial setting time, and F ¼ final setting time. In this instance, CT is defined as the time when the cement paste no longer disintegrates in an aqueous solution [49]. The initial setting time window is relatively large and depends on the application procedure of the specific cement. For example, for dental applications, shorter initial setting times are usually preferred, whereas longer initial setting times are needed for orthopedic procedures. Regardless, the maximum final setting times for all applications generally should not exceed 15 min [49]. Strategies to Improve Setting Times A wide array of strategies has been developed over the years to modify the setting rate of the cements so that they fall within the desired time window. Generally, brushite cements react and set much faster than apatite cements. This means that the setting reaction for brushite cements needs to be delayed whereas the setting of apatite cements needs to be accelerated to meet clinical requirements [14,40]. One of the most straightforward and common approaches toward this end is to change the particle size of the CaP precursor compounds [51]. Setting times can be significantly reduced by decreasing the mean particle size, because the size reduction increases the surface area of the particles that can then interact with the liquid phase, thus accelerating the setting reaction and hydration kinetics of the

FIGURE 34.3 A diagram outlining the setting time stages of calcium phosphate cements from a clinical perspective where CT ¼ cohesion time, I ¼ initial setting time, and F ¼ final setting time [29].

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cement [26,34,46,47]. When this is the case, higher amounts of dissolved Ca2þ and PO4 3 ions can be reached in a shorter time, increasing the degree of supersaturation that favors crystal nucleation and growth [52]. Another simple approach to modifying the setting time is to change the amount of liquid phase used in the cement. As is the case for most apatite cements, the amount of liquid phase used is brought to a minimum in an attempt to reduce the setting time, whereas the opposite approach is employed for brushite cements. This is one reason why apatite cements have a tendency to be more viscous and difficult to inject compared with less viscous brushite cements [26,29]. The setting time can also be reduced by incorporating soluble orthophosphates such as Na2HPO4 or NaH2PO4 into the liquid phase to reduce the pH. Creating a more acidic environment increases the solubility of the CaP compounds and accelerates the setting reaction, which shortens the overall setting time [33,53]. Furthermore, the presence of these sodium orthophosphates prevents unreacted CaP particles from becoming isolated, which could prolong the setting time [29]. Organic acids such as lactic, glycolic, tartaric, citric and malic have also been extensively used to influence the setting time of cements and have been well-described elsewhere [54]. It was also observed that for some cases, an initial thermal treatment of the CaP precursor compounds extended the setting reaction. For example, when a-TCP particles were first heated to 500 C, it increased the final setting time from the order of several minutes to upward of a few hours [55]. Finally, it is common practice to add a nucleating agent such as nanosized HA to the powder component of the cements to reduce the setting time [26,51].

Injectability CPCs have the potential to be injectable, which sets them apart from other bone substitute materials, especially for applications involving minimally invasive surgery, in which injectability is considered a prerequisite [29]. Although there has been much debate in the literature in terms of defining injectability [15], the overall fundamental consensus is that it can be defined as the ability of the CPC paste to be extruded through a syringe, usually with a 12-gauge (i.e., 2 mm in diameter) needle 10 cm long attached to the end, although other needle dimensions have been reported [56e58]. Generally, it is measured by calculating the weight percentage of the CPC paste that can be extruded from the syringe either manually by hand or with the aid of a mechanical bench in which a force of no more than 100 N is applied [15,59]. To understand the injectability of these cements better, one must take into account the flow dynamics of the paste within the syringe. Because CPCs are composed of a powder and liquid phase, they are considered to be a biphasic material. When this biphasic material is subjected to a pressure gradient, as is the case during extrusion through a syringe, the liquid phase has a tendency to flow faster than the powder phase, which can result in local changes of the final cement composition. More specifically, because of the pressure gradient present in the syringe, the paste located closest to the plunger of the syringe is subjected to a higher pressure than the paste farther from the plunger. This higher pressure forces the liquid to flow faster than the CaP particles to the point where the paste in this region becomes depleted of liquid and all that is left is a wet powder instead of a paste [15,56,58]. The opposite effect occurs for the paste in the region farthest from the plunger, where the pressure is lower and the paste is saturated by liquid. Because these effects are dynamic, the size of the region that is depleted of liquid grows and enlarges during injection to the point where this wet powder zone eventually reaches the tip of the syringe and plugs it, rendering the remaining material noninjectable. This phenomenon, in which a phase separation occurs once pressure is applied to the paste, is commonly referred to as filter pressing [15]. As a result of the phase separation, the final composition of the extruded paste becomes compromised and not fully controllable. This is caused by the change in the LPR that takes place, which can influence the setting behavior, mechanical properties, and biological performance of the final cement product. Hence, good cohesion of the paste is of utmost importance to avoid problems associated with filter pressing [60]. Strategies to Improve Injectability As mentioned previously, an important advantage of CPCs is their ability to be injected for use in minimally invasive surgical techniques, although this ability can become compromised when filter pressing occurs [54,56,57]. In fact, in some cases complete injectability cannot be attained even when relatively large-diameter syringe tips without a cannula are used. This implies that filter pressing can occur even under very small loading forces [11]. Thus, there is a need to develop strategies to improve the injectability of the cements. Strategies involving the addition and/or modification of certain variables have been shown to be successful for improving the injectable properties of CPCs. These variables include modifying the particle size and shape [56], the viscosity and rheology of the paste [54,57,61e64], and the LPR [15,57,65].

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Particle Size and Shape With respect to the particle size, it has been shown that the smaller the particle size, the greater the injectability [56,65]. For the particle shape, powders with a more spherical shape have been known to exhibit better injectability because the particles roll more easily and produce less shear resistance when they are extruded from the syringe. Interestingly, more spherical powders also require less liquid phase to turn into a slurry or paste, because no liquid becomes captured between the particles and therefore the liquid can be used more efficiently to wet the particles fully [11,15]. Viscosity It is well-known that the ability of the cements to be injected is highly constrained owing to the rheological properties of the paste during extrusion. Reducing filter pressing or eliminating it from occurring by modifying the viscosity of the paste is one possible solution to improving injectability properties. One approach would be to reduce the ability of the liquid phase to pass through the powder phase. This could be achieved by increasing the viscosity of the liquid phase or by reducing the permeability of the powder phase [54,56,57,61e63]. The viscosity of the liquid phase can be modified by using additives. These additives must be nontoxic and not inhibit the phase transformation that takes place during the cement setting reaction [11]. In principle, the use of additives in the liquid phase can increase the viscosity by several orders of magnitude, resulting in the formation of a more putty-like paste. Furthermore, they can improve wettability and increase the surface charge of the CaP particles owing to the adsorption of ions on the surface [64]. Studies have shown that the addition of organic additives, such as citric acid, increased injectability by delaying the hydration time [11,66]. Furthermore, citric acid is a favorable additive for use in CPCs because it is present as citrate ions in bone and is a triprotic acid, which means that it can release or donate three hydrogen ions. This is advantageous in the sense that these citrate ions can be adsorbed onto the surface of CaP particles and give them a negative charge, and by doing so it can create a repulsion between the negatively charged particles. This repulsion enhances the mobility of the particles, increasing the injectability of the paste, because the formation of any particle agglomerations is eliminated [61]. The behavior of CPCs also depends on the type of acid additive used. For example, it was observed that the use of acetic acid, a monoprotic acid, resulted in an increase in the rate of apatite formation, which made the cement less injectable. Conversely, when citric acid was employed, the CPC became more macroporous and easier to inject, although the rate of apatite formation was reduced [59]. Another additive that was shown to improve the injectability as well as cohesion of CPCs is hydroxypropyl methyl cellulose (HPMC). HPMC, a polysaccharide gelling agent, has the ability to hydrogenate in water, forming a viscous solution in the process. This increase in viscosity improves the cohesion of the paste and makes it easier to extrude out of a syringe under minimal force with no occurrence of filter pressing [66,67]. In another study, researchers added xanthan, a type of polysaccharide, to the cement formulation to exert a lubricating effect on the interface of the CaP particles, which improved injectability [56]. Yet another technique decreases the viscosity to improve injectability. One way to achieve this is to increase the amount of liquid phase used. The idea behind this is that as more liquid phase is used, the viscosity becomes lower and the friction between the CaP particles as well as the cement paste between the syringe walls is reduced. This would make the cement easier to inject, although issues such as filter pressing are likely to occur [64,68]. Finally, the viscosity of CPCs is a constantly changing property because it is related to the setting mechanism. Traditionally, the viscosity decreases immediately after mixing the liquid and powder phases together, followed by a strong increase during the setting of the cement until hardening occurs. The viscosity needs to be at a high enough value to where extravasation of the liquid from the powder can be prevented. This can be accomplished if the viscosity falls between 100 and 1000 Pa s1 and an appropriate window for injection is defined [11]. Liquid-to-Powder Ratio Any adjustment to the LPR can have a significant influence on not only the injectability but also the setting time, porosity, cohesion, resorbability, and strength [68]. For example, high LPRs are known to increase the injectability of cements by reducing the friction between CaP particles and the syringe wall during extrusion, but this additional liquid can delay the setting time, result in poor cohesion [69], and reduce the mechanical integrity of the cement because the higher liquid content leaves behind a greater microporous structure upon dilution [11,56]. Studies have shown how the injectability of CPC formulations with LPRs between 3.85 and 4.50 g/mL have been relatively unaffected, but once this range increases to 4.50e5.00 g/mL the degree of injectability drops by nearly 100% [64]. Consequently, it seems that only major shifts in the general properties of CPCs occur once a certain LPR threshold is met.

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Cohesion Besides injectability, CPCs need to possess adequate rheological properties such as viscosity and cohesion to be employed successfully for clinical applications [15,70]. In this instance, cohesion is defined as the ability of a cement paste to keep its geometrical integrity during setting in an aqueous environment [15,26]. It can be measured by evaluating the amount of solid particles that are released from the paste under aqueous conditions, typically water or Ringer’s solution before setting fully [71]. The level of cohesion is directly correlated to the degree of affinity among CaP particles within the cement, where stronger cohesive properties can result from an increase in attractive van der Waal forces and/or a decrease in repulsive electrostatic forces among the CaP particles [15]. The ideal cohesion of CPCs occurs when no disintegration of CaP particles can be observed while immersed in an aqueous environment for a period of time exceeding the final setting time of the cement [49,60]. This level of cohesion can be achieved by maintaining a high level of viscosity for the cement paste, usually through the use of gel-forming polymers or other cohesion promoters such as sodium alginate and carboxymethyl cellulose [71,72]. In the past, poor cohesive properties were associated with poor biocompatibility and negative in vivo reactions, such as inflammation triggered by the leaching and release of CaP microparticles in and around the implant site [37]. It was argued that CPCs could be considered a prime candidate for the use in vertebroplasty spinal surgery, provided that excellent cohesion of the CPC could be achieved and maintained [32]. Poor cohesive properties of CPCs used for vertebroplasty can lead to an increased risk for blood clotting. It was suggested that blood clotting is triggered by interfacial interactions between blood and CaP microparticles that fragmented off the cement and leached into the bloodstream. Coupled with the fact that vertebrae are highly perfused with blood and positioned in close proximity to the heart and lungs, this compounds the problem and exposes the patient to potentially life-threatening complications. Therefore, it is of utmost importance to maintain a high level of cohesion for CPCs to prevent the release of CaP particles and avoid in vivo complications [26]. Strategies to Improve Cohesion The level of cohesion is directly associated with the interaction among CaP particles within the cement. Bearing that in mind, common approaches to improving cohesion are focused on reducing the mean particle size and liquid content, thus improving the van der Waal attractive forces and decreasing electrostatic repulsive forces among the CaP particles [60]. CPC pastes are considered to be non-Newtonian fluids, which means that their viscosity is a function of shear forces and time [65,70]. This poses a challenge to researchers to be able to maintain and control the viscosity, and therefore the cohesion, of the cements. However, some success has been achieved through the strategy of increasing the viscosity of the liquid phase by incorporating water-soluble polymeric hydrogels into the formulation [60,67]. Most notably, polysaccharides [32,53,67,71,72] polyacrylic acid [29], and gelatin [69] are among the most common hydrogels used owing to their biocompatibility and favorable rheological properties. Moreover, only small amounts of these polymers, on the order of just a few weight percent, need to be added to CPCs to improve their viscosity and cohesion significantly, which makes them less susceptible to washout effects [29]. In some instances, it has been shown that adding certain types of these water-soluble polymers can prolong the setting time of CPCs. To address this issue, sodium phosphate aqueous solutions have been used as the liquid phase to act as a catalyst in the setting time [73]. Other water-soluble polymers, such as sodium alginate, exhibit the unique ability to pectize (i.e., form into a jelly or gel-like substance) when placed in contact with calcium ions. This enables putty-like pastes to be fabricated, although only a few polymers displaying this unique property have been accepted for parenteral use [14,40]. Regardless, the addition of water-soluble polymers has been shown to be a promising technique to improve the viscosity and cohesion of CPCs so that the spectrum of their clinical applications can widen.

STRATEGIES TO IMPROVE THE MECHANICAL PROPERTIES The development of CPCs has progressed significantly with respect to the optimization of their physiochemical, handling, and biological properties. Nevertheless, CPCs exhibit poor mechanical properties, because their tensile and shear strength is lower than that of teeth and bone [43]. Coupled with their inherent brittle behavior, this has limited their clinical applicability to noneload bearing or pure compression-loading sites [74,75]. The compressive strength of CPCs is the most commonly reported mechanical property in the literature, although this property offers minimal insight into the mechanical integrity and reliability of the CPC [15]. Generally, brittle materials are more likely to fail under tension or shearing rather than compression. This phenomenon is attributed

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to the fact that fracturing of brittle ceramics such as CPCs initiates in preexisting flaws such as microcracks or macropores located throughout the cement matrix. It is easier to propagate a crack in tension rather than in compression [76]. Furthermore, loads within the body are generally cyclic, which indicates that fatigue properties should be reported as well [15,77]. Moreover, it is difficult to compare compressive strengths of CPCs with those of trabecular bone because bone is typically much less brittle than bioceramic cements. Therefore, it is highly recommended that the mechanical properties of CPCs be reported as more appropriate parameters such as the Weibull modulus to make a more comparative analysis between bone tissue and cement [77]. Another factor that should be taken into account regarding the mechanical properties of CPCs is that their mechanical performance in vivo may vary considerably from in vitro measurements under laboratory conditions. For example, brushite cements are highly soluble under physiological conditions compared with apatite cements, so their mechanical properties will decrease rapidly upon implantation. This sudden and spontaneous dissolution is also why brushite cements are typically combined with less soluble materials, e.g., HA or b-TCP, thus creating a biphasic CPC [15]. Based on these shortcomings, it is clear why CPCs are restricted to noneload bearing applications. An improvement in the mechanical performance of CPCs would broaden their applicability in the operating theater to more load-bearing defect sites, e.g., spinal surgical procedures such as vertebroplasty and kyphoplasty [75,78]. To this end, the brittleness of the cements would need to be reduced along with an improvement in their fracture toughness. This toughening can be achieved by reducing the porosity of the cement matrix or combining CPCs with a polymer matrix to form a composite material [79]. Generally, these polymers are incorporated into the cement by modifying the liquid phase with polymeric additives [78] in the form of fibers [75], or by using a dual setting system in which dissolved monomers simultaneously cross-link during the cement setting reaction [80,81], as depicted in Fig. 34.4. More common formulations of various CPCepolymer composite materials have been comprehensively reviewed elsewhere [11,74].

Porosity The mechanical strength of CPCs is directly related to their microstructural characteristics. The physical entanglement of crystals that precipitate during the setting reaction, as described earlier, results in the formation of a bulk cement matrix that determines its mechanical integrity. Several factors can affect the outcome of the cement matrix, such as the final setting products (i.e., apatite or brushite), the degree of phase transformation, the crystal size, and the porosity; with the latter being one of the most important factors. The porosity of CPCs mainly originates from the presence of an excess amount of liquid phase that has not fully reacted with the CaP particles. When this is the case, this excess liquid leaves behind open voids or pores within the entangled cement crystal matrix [78]. Generally, these pores can range from a few nanometers to upward of 8e12 mm in diameter and consume between 22% and 55% of the total cement volume [82,83]. The presence of these pores makes it much easier for microcracks to propagate throughout the cement matrix, increasing the cement porosity and decreasing the mechanical properties [29], as shown in the inverse exponential relationship [84]: CS ¼ CS0exp(-KP)

(34.3)

where CS is the compressive strength of the cement at a set porosity, CS0 is the maximum theoretical strength, K is a constant, and P is the porosity [84]. Therefore, the mechanical properties of cements can be improved by decreasing Reinforcement Strategies for CPCs

Intrinsic (i.e.Porosity)

Compression

FIGURE 34.4

Bimodal Particle Size Distribution

Extrinsic Dual Setting System

Fiber Reinforcement

Mechanical reinforcement strategies for calcium phosphate cements used in load-bearing applications [78].

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the matrix porosity [26]. As a rule of thumb, the tensile strength of CPCs seems to increase two-fold with every 10 vol% decrease in porosity. For example, tensile strength values of 5, 10, and 20 MPa would correlate to porosity values of 60%, 50%, and 40%, respectively [14,40]. Consequently, one method to achieve this reduction in porosity is to apply pressure to compact the CaP particles [73,84,85]. In some cases, this method has been shown to reduce the porosity from 50% to 31%, which resulted in an increase in compressive strength by almost one order of magnitude, but this method is typically unsuitable for most clinical applications [84]. Because porosity is directly attributed to the amount of liquid phase used, a more common approach to reducing it is to increase the LPR by reducing the amount of liquid phase added to the cement [14,40]. When this happens, the space between the CaP particles is reduced, enabling a more compact crystal structure to form, as illustrated in Fig. 34.5 [52]. The influence of the LPR on porosity explains why the porosity of apatite cements is usually different from that of brushite cements. The general trend is that for brushite cements the average pore size is larger than for apatite cements, but the total porosity is usually smaller than for apatite cements. This is because brushite cements have an increased water consumption during the setting reaction that also forms larger crystal sizes, which makes the average pore size greater but the total porosity smaller than apatite cements [78]. Although a reduction in the liquid content leads to a decrease in porosity, this approach is limited because all CPC formulations require a minimum amount of liquid to wet all of the CaP particles fully and create a paste [56]. Moreover, a large reduction in the liquid phase can have a negative effect on the rheological properties of the paste, such as increasing the viscosity to the point where the paste becomes noninjectable [29]. Therefore, another effective strategy for reducing the porosity of the cements is to minimize spacing between the CaP particles. This can be achieved by using a two-pronged approach: (1) create a bimodal CaP particle size distribution to fill the space among particles; and (2) create a high surface charge, or z-potential, on the particles. By using a bimodal particle size distribution, the smaller particles are able to occupy the space normally occupied by any

FIGURE 34.5 Schematic depiction of how the (A) particle size of the powder phase and (B) the liquid-to-powder ratio (LPR) influences the microstructure and porosity of calcium phosphate cement (CPCs). A reduction in calcium phosphate (CaP) particle size increases the particle surface area, which means a higher degree of supersaturation can be achieved that favors crystal nucleation. This leads to more frequent and smaller needle-like crystals to form instead of larger plate-like crystals that typically form when larger CaP particles are used, effectively changing the microstructure and porosity of the cement. Modifying the LPR can also fine-tune the microstructure and porosity of CPCs. Lowering the LPR reduces the spacing among CaP particles, leading to a more compact and dense crystal microstructure. The opposite effect can be seen by increasing the LPR. Reprinted with permission from Ginebra M-P, Canal C, Espanol M, Pastorino D, Montufar EB. Calcium phosphate cements as drug delivery materials. Adv Drug Deliv Rev 2012;64(12):1090e1110.

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excess liquid, thus reducing the porosity for both apatite [86] and brushite cements [82,83]. Adding a high surface charge by increasing the z-potential on the CaP particles is thought to aid in dispersing and breaking up any particle agglomerates by reducing their attractive forces. This can be achieved by adding charged ions such as tartrates or citrates to the liquid phase of CPC formulations. These charged ions are then able to adsorb to the particle surface, effectively increasing the z-potential and removing particle agglomerations, leaving behind a more evenly dispersed particle network that exhibits smaller particle spacing [87]. Applying these two principles has been shown to increase the plastic limit (i.e., reduce the need to use more liquid) and reduce the porosity for both apatite and brushite cements. For apatite cements, a reduction in the porosity from 37% to 25% with an increase in the compressive strength from 50 to 79 MPa has been reported [86], and for brushite cements, a reduction in porosity from 30% to 23% resulted in an improvement in compressive strength from 23 to 42 MPa [82]. Finally, although a reduction in porosity leads to an improvement in the mechanical properties of CPCs, this change in the microstructure can have negative implications on the bioresorbability and osteotransductive properties of the cement. For example, the lack of an interconnected porous structure inhibits the growth of newly formed bone into the cement. This means that new bone formation would be dictated by the passive dissolution rate of the cements, resulting in a reduction in the rate of new bone formation. Therefore, it is important to understand and take into account how porosity may affect other properties of the cement, especially with respect to the biodegradability [14,40].

Dual Setting System Another approach to increasing the mechanical strength of CPCs involves using a dual setting system. An example of this strategy involves adding polymeric components to the cement that can be cross-linked by binding Ca2þ ions with carboxylic acid or organic phosphate moieties within the polymer chain. The addition of the liquid phase then initiates the dual setting reaction composed of (1) the traditional dissolution-precipitation reaction of the bioceramic cement component, and (2) deprotonation of the acid groups of the polymer to induce the formation of intrachained or interchained Ca2þ-acid cross-links. The result of this process is a reduction in brittleness and an increase in the strength of the cement [88]. Another example of the dual setting system involves using reactive monomer units that are added to the cement formulation by first dissolving them in the liquid phase. An initiator is then added to the solid phase of the cement so that once the two phases are mixed, a simultaneous reaction is initiated that involves the traditional dissolutionprecipitation setting reaction of the CaP particles and a gelation-polymerization reaction of the monomer units [78]. Within the first several minutes, a hydrogel matrix with embedded CaP particles is formed, followed by a phase transformation of the CaP particles that takes place during the setting reaction of the cement. Finally, a highly interconnected, reinforcing hydrogel matrix is formed that is located within the porous cement microstructure, as depicted in Fig. 34.6. This approach enables the possibility of adding a high amount of polymer to the cement, which translates to a potentially significant increase in its strength and toughness. Furthermore, the rheological properties of the cement paste can remain relatively stable. Both of these characteristics are attributed to the fact that the Aqueous Monomer Solution Monomer Polymerization Reaction

Dual Setting Cement System

Mixing

CPC Setting Reaction CaP Powder Phase

FIGURE 34.6 Schematic diagram depicting the setting mechanism of dual-setting calcium phosphate cements (CPCs) with the formation of interconnected matrices of hydrogel and precipitated calcium phosphate (CaP) crystals [78].

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monomer units are small, water-miscible liquids that exhibit low viscosity such that even high concentrations of the monomers do not alter the initial viscosity of the cement [78]. Further studies regarding this concept can be found elsewhere [80,81,89]. A dual setting system can also be applied when using pure inorganic materials such as silica. Studies describing different approaches to adding silica to the cement, along with its influence on CPCs setting, handling, and mechanical properties can be found elsewhere [78,90,91].

Fiber Reinforcement Toughening of brittle cements by using fibers has proven to be one of the most successful approaches, particularly when using long continuous fibers as a reinforcing matrix [32,75,79,92]. The mechanical performance of these fiberreinforced CPCs depends on the complex interaction among all of the components of the composite cement [79]. The mechanical properties are also time-dependent because both the fibers and the cement matrix have the potential to degrade once they are implanted in the body to allow for tissue regeneration. These components are illustrated in Fig. 34.7, in which contributions to the mechanical behavior of the cement are associated with the strength and stiffness of the fibers and cement matrix, the toughness of the matrix, interfacial interactions between the matrix and fibers, and supplementary effects of the polymeric additives or agglomerations [79]. Excellent reviews detailing how various fiber parameters (i.e., type, volume fraction, orientation, aspect ratio, tensile modulus, and fiberematrix interface properties) can influence the mechanical properties of CPCs can be read elsewhere [75,76,78,79]. Mechanics of Fiber-Reinforced Calcium Phosphate Cements The concept of improving the mechanical performance of CPCs by producing fiberecement composites was initially adopted from the field of materials engineering, where, for instance, the development of fiber-reinforced hydraulic cements and concretes for civil and industrial engineering purposes has been studied extensively. For these applications, the incorporation of fibers into the cement has been a highly effective strategy to improve fracture toughness as well as flexural and tensile strength [76]. To understand better how fibers can improve the mechanical properties of CPCs, the failure pattern of fiber-free CPCs must first be understood. Like most ceramics, CPCs are brittle materials, which means that their failure is characterized by a sudden, catastrophic fracture with minimal plastic deformation occurring before the failure point. Therefore, fracture deformation and the work needed to induce failure are relatively low. From an engineering

Fiber Aspect Ratio Fiber Volume Content

Fiber Orientation

Fiber Mechanical Properties

Mechanical Behavior of Fiber-Reinforced CPCs

Degradation Kinetics

Fiber-Matrix Intreface Properties

Matrix Mechanical Properties Fiber Architecture

FIGURE 34.7 Illustration of the components for fiber-reinforced calcium phosphate cements (CPCs) that influence their mechanical perfor-

mance [79].

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FIGURE 34.8 Schematic drawing of the three mechanisms of how fibers can mechanically reinforce cements [76].

standpoint, this brittle behavior is detrimental for mechanically loaded materials and therefore must be compensated for, in this case by incorporating fibers. The reinforcement effect of fibers in CPCs creates a tougher, more ductile cement that is more suitable for load-bearing applications because they will exhibit a higher tolerance to impact loading and sample flaws [79]. There are three main mechanisms according to how a fiber can mechanically reinforce a cement: fiber bridging, crack deflection, and frictional sliding (Fig. 34.8). In fiber bridging, fibers that bridge crack together within the matrix, thus effectively dissipating the fracture energy and delaying any further opening and propagation [79]. For crack deflection, fibers act as barriers that extend the path distance at which the crack needs to travel through the matrix. It has also been reported that these two mechanisms contribute to the high fracture toughness of human bone [93]. Frictional sliding consists of frictional sliding occurring at the fiberematrix interface when fibers begin to pull out of the cement. This mechanism leads to a stress transfer that ultimately results in increased energy dissipation and fracture toughness of fiber-reinforced CPCs [76]. The load-bearing capacity of fibers increases with an increase in the elastic modulus as well as tensile strength. The diameter of the fibers dictates the total interface area for a given fiber volume fraction, which affects both the homogeneity and workability of the composite. Furthermore, the fiber length to diameter ratio, known as the aspect ratio, is greatly important because loading of the composite needs to be transferred from the matrix to the fibers via the interface. A reinforcing effect is observed only when the fiber length exceeds a critical value, lc [94]. Extensive research from civil engineering revealed that the optimum fiber content to reinforce cementitious materials is typically less than 5 vol%, whereas the fiber content for fiber-reinforced CPCs is usually one order of magnitude higher. This higher fiber content is partially attributed to the fact that only a moderate load transfer has been achieved for fiber-reinforced CPCs because the interface strength has not yet been optimized for bioceramic cements [92]. Finally, because mechanical testing of CPCs is not strictly regulated, caution must be exercised when comparing mechanical values from different studies, because many variables during sample preparation and testing can strongly influence the final results [78]. For example, the strength of dried samples is typically superior to that of hydrated samples, because excess water acts as a lubricant within the entangled crystal network structure of the cement. Moreover, the macroporosity and microporosity of the cement can also be heavily influenced during sample preparation, in which any precompacting of the paste would result in a denser, less porous sample that would exhibit higher strength values compared with uncompacted samples [71,85].

CLINICAL APPLICATIONS Oral, Maxillofacial, and Craniofacial Applications CPC was first approved for human use by the US Food and Drug Administration in 1996. Since then, the number of CPC products and clinical indications has increased considerably, most notably in the oral, maxillofacial, and craniofacial fields [17]. CPCs are increasingly used in these fields because the CPC is stressed only moderately in these applications. Furthermore, the excellent handling properties of the CPC provide the ability to mold the material, which is highly advantageous from a surgeon’s perspective. These cements have been used to repair neurosurgical burr holes, contiguous craniotomy cuts, and other cranial defects. They have also been indicated for use in sinus augmentation procedures and orbital reconstruction surgery [27,95]. Further information outlining CPCs use for oral, maxillofacial, and craniofacial applications can be found elsewhere [29].

CONCLUSION

607

Dental Applications Clinical use of CPCs for dental and intraoral applications has been relatively lacking compared with their use in other fields such as craniomaxillofacial and orthopedic indications. However, research has indicated that CPCs can be highly beneficial in a number of dental and intraoral procedures. To understand better the potential of CPCs to stimulate new bone formation for dental applications, a group of researchers conducted a study in which they implanted prefabricated CPC blocks into the alveolar bone of dogs. They started by extracting all of the mandibular premolar teeth; then they waited 1 month for the alveolar bone to reduce in size to make room to implant 8-mm CPC preset blocks. Over the course of 1 month the researchers observed that the CPC block was slowly being replaced by bone. Histopathologic images further proved that the implant site exhibited features similar to those of natural bone. In addition, it was observed that the coronal half of the implanted CPC was firmly attached to the natural bone [96]. In another study, a CPC was injected into artificially created periodontal defect sites, where it was observed that the cement acted as a scaffold for new bone growth and promoted healing of the periodontal tissue [97]. CPCs have also been indicated for use in direct pulp capping, where investigators compared a self-setting CPC with that of calcium hydroxide. It was concluded that both materials were successful in producing secondary dentin after 24 weeks [98]. Finally, other studies have shown promising results for the use of CPC in restoring enamel carious cavities, for fillings in root canal procedures, for alveolar ridge augmentation, as sinus lifts, for repair of the cleft palate, and as supportive agents for dental implants [17,29]. It has been proven over the years that CPCs can exhibit superior results in these dental applications owing to their excellent osteoconductivity, unique self-hardening properties, excellent affinity to bone defect surfaces, and gradual resorption and replacement by new bone [17].

Orthopedic Applications CPCs are predominantly used for orthopedic applications because of their distinctive handling and biological properties. The cements have been applied to a wide range of clinical procedures including: hip fractures, tibial plateau fractures, fixation of bone screws and titanium implants, vertebral body fillings and augmentation of osteoporotic-induced vertebral bodies, and distal radius fractures [36,99]. Vertebroplasty and Kyphoplasty Two orthopedic procedures, vertebroplasty and kyphoplasty, have become increasingly popular areas in which CPCs have been employed. These procedures aim to treat osteoporosis-induced vertebral compression fractures by augmenting, stabilizing, and restoring weakened vertebra to their normal functional state and height as best as possible [100]. Vertebroplasty involves the direct injection of CPC into the fractured vertebral body, whereas an inflatable balloon tamp is used for kyphoplasty to create a cavity in the vertebral body into which CPC can then be injected to fill the cavity. Results from both procedures have shown promising results with respect to faster healing times of the vertebral body [32]. The type of CPC, and thus its properties, have a crucial role in the overall outcome and success of the procedures. For example, the CPCs must exhibit some key desirable properties such as easy injectability, excellent cohesion to prevent possible leaching of CaP particles into the bloodstream where blood clotting could be triggered, high radiopacity, mechanical properties similar to healthy vertebral bone, and a resorption rate that is comparable to normal bone remodeling times [31]. Owing to these demands, only a few CPC formulations have been allowed for clinical use in this field, and even these formulations are not necessarily considered ideal. For example, several major topics of interest still need to be investigated in more detail to optimize and tailor CPC formulations for future use. In particular, understanding the quasi-static compressive strength and fatigue performance of CPC formulations is highly interesting, especially when considering the in vivo life expectancy of the cement [32]. Overall, it has been demonstrated that the use of CPCs for vertebral bone compression fractures has substantial clinical benefits to the point where this treatment method is on the rise and is growing in popularity [11].

CONCLUSION The rapid development of CPCs confirms that it exhibits unique characteristics that render these types of bioceramics highly suitable as a bone substitute material. CPCs are formed in a simple procedure by combining a CaP powder phase with a liquid phase to form a cementitious paste at body temperature. This paste undergoes a phase transformation at which it eventually hardens in vivo into a material that shares a composition similar to

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the inorganic component found in natural bone. This self-setting capability allows for CPCs to be injectable so that they can be used in minimally invasive surgical techniques; they are also moldable so that they can adapt and fit into irregularly shaped bone defect sites. Their chemical similarity to natural bone gives them excellent osteotransductive properties. The setting times as well as cohesiveness and injectability of CPCs have improved considerably, indicating that CPCs have a promising future in the field of bone regeneration. However, the perfect bone grafting material does not yet exist, and CPCs have much room for improvement and development. In particular, CPCs are still regarded as an inadequate material for use in load-bearing sites, which severely restricts their clinical applicability. Nevertheless, many CPC products have become commercially available that are suitable for several applications, particularly in dental, maxillofacial, and craniofacial surgeries. It is anticipated that a better understanding of the fracture mechanics of CPCs will lead to an improvement in their mechanical properties, which would significantly expand their clinical relevance.

List of Acronyms and Abbreviations ACP Amorphous calcium phosphate CaP Calcium phosphate CDHA Calcium-deficient hydroxyapatite CPC Calcium phosphate cement CT Cohesion time DCPA Dicalcium phosphate anhydrous DCPD Dicalcium phosphate dehydrate F Final setting time HA Hydroxyapatite I Initial setting time LPR Liquid-to-powder ratio MCPM Monocalcium phosphate monohydrate OCP Octacalcium phosphate TTCP Tetracalcium phosphate a-TCP a-Tricalcium phosphate b-TCP b-Tricalcium phosphate

Glossary Cohesion The ability of a cement paste to keep its geometrical integrity during setting in an aqueous environment Injectability The ability of the CPC paste to be extruded through a syringe, usually with a 12-gauge (i.e., 2 mm in diameter) needle 10 cm long attached to the end Osteoconduction The ability to have bone tissue grow on a material’s surface Osteotransduction The ability of a material to be resorbed and replaced by new bone tissue Stoichiometric The quantitative relationship between reactants and products within a chemical reaction

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Tissue responses of calcium phosphate cement: a study in dogs. Biomaterials 2000;21(12): 1283e90. Constantz BR, Ison IC, Fulmer MT, Poser RD, Smith ST, VanWagoner M, et al. Skeletal repair by in situ formation of the mineral phase of bone. Science 1995;267(5205):1796e9. Miyamoto Y, Ishikawa K, Takechi M, Toh T, Yuasa T, Nagayama M, et al. Histological and compositional evaluations of three types of calcium phosphate cements when implanted in subcutaneous tissue immediately after mixing. J Biomed Mater Res 1999;48(1):36e42. Mirtchi AA, Lemaitre J, Terao N. Calcium phosphate cements: study of the b-tricalcium phosphatedmonocalcium phosphate system. Biomaterials 1989;10(7):475e80. Bohner M, Van Landuyt P, Merkle H, Lemaitre J. Composition effects on the pH of a hydraulic calcium phosphate cement. J Mater Sci Mater Med 1997;8(11):675e81. Bohner M. Physical and chemical aspects of calcium phosphates used in spinal surgery. Eur Spine J 2001;10(2):S114e21. Bohner M, Merkle H, Van Landuyt P, Trophardy G, Lemaitre J. Effect of several additives and their admixtures on the physico-chemical properties of a calcium phosphate cement. J Mater Sci Mater Med 2000;11(2):111e6. Marin˜o FT, Torres J, Hamdan M, Rodrı´guez CR, Cabarcos EL. Advantages of using glycolic acid as a retardant in a brushite forming cement. J Biomed Mater Res B Appl Biomater 2007;83(2):571e9. Dorozhkin SV. Calcium orthophosphate cements for biomedical application. J Mater Sci 2008;43(9):3028e57. Lacout J, Mejdoubi E, Hamad M. Crystallization mechanisms of calcium phosphate cement for biological uses. J Mater Sci Mater Med 1996; 7(6):371e4. Zoulgami M, Lucas A, Briard P, Gaude´ J. A self-setting single-component calcium phosphate cement. Biomaterials 2001;22(13):1933e7. Ginebra M, Driessens F, Planell J. Effect of the particle size on the micro and nanostructural features of a calcium phosphate cement: a kinetic analysis. Biomaterials 2004;25(17):3453e62. Liu C, Shao H, Chen F, Zheng H. Effects of the granularity of raw materials on the hydration and hardening process of calcium phosphate cement. Biomaterials 2003;24(23):4103e13. Driessens F, Boltong M, Bermudez O, Planell J. Formulation and setting times of some calcium orthophosphate cements: a pilot study. J Mater Sci Mater Med 1993;4(5):503e8. Khairoun I, Boltong M, Driessens F, Planell J. Limited compliance of some apatitic calcium phosphate bone cements with clinical requirements. J Mater Sci Mater Med 1998;9(11):667e71. Despas C, Schnitzler V, Janvier P, Fayon F, Massiot D, Bouler J-M, et al. High-frequency impedance measurement as a relevant tool for monitoring the apatitic cement setting reaction. Acta Biomater 2014;10(2):940e50. Bohner M. Reactivity of calcium phosphate cements. J Mater Chem 2007;17(38):3980e6. Ginebra M-P, Canal C, Espanol M, Pastorino D, Montufar EB. Calcium phosphate cements as drug delivery materials. Adv Drug Deliv Rev 2012;64(12):1090e110. Burguera EF, Xu HH, Weir MD. Injectable and rapid-setting calcium phosphate bone cement with dicalcium phosphate dihydrate. J Biomed Mater Res B Appl Biomater 2006;77(1):126e34. Leroux L, Hatim Z, Freche M, Lacout J. Effects of various adjuvants (lactic acid, glycerol, and chitosan) on the injectability of a calcium phosphate cement. Bone 1999;25(2):31Se4S.

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[55] Bohner M, Luginbu¨hl R, Reber C, Doebelin N, Baroud G, Conforto E. A physical approach to modify the hydraulic reactivity of a-tricalcium phosphate powder. Acta Biomater 2009;5(9):3524e35. [56] Bohner M, Baroud G. Injectability of calcium phosphate pastes. Biomaterials 2005;26(13):1553e63. [57] Khairoun I, Boltong M, Driessens FM, Planell J. Some factors controlling the injectability of calcium phosphate bone cements. J Mater Sci Mater Med 1998;9(8):425e8. [58] Burguera EF, Xu HH, Sun L. Injectable calcium phosphate cement: effects of powder-to-liquid ratio and needle size. J Biomed Mater Res B Appl Biomater 2008;84(2):493e502. [59] Montufar E, Maazouz Y, Ginebra M. Relevance of the setting reaction to the injectability of tricalcium phosphate pastes. Acta Biomater 2013; 9(4):6188e98. [60] Bohner M, Doebelin N, Baroud G. Theoretical and experimental approach to test the cohesion of calcium phosphate pastes. Eur Cell Mater 2006;12(1473e2262):26e35. [61] Wang X, Ye J, Wang H. Effects of additives on the rheological properties and injectability of a calcium phosphate bone substitute material. J Biomed Mater Res B Appl Biomater 2006;78(2):259e64. [62] Ratier A, Freche M, Lacout J, Rodriguez F. Behaviour of an injectable calcium phosphate cement with added tetracycline. Int J Pharm 2004; 274(1):261e8. [63] Habib M, Baroud G, Gitzhofer F, Bohner M. Mechanisms underlying the limited injectability of hydraulic calcium phosphate paste. Acta Biomater 2008;4(5):1465e71. [64] Gbureck U, Barralet JE, Spatz K, Grover LM, Thull R. Ionic modification of calcium phosphate cement viscosity. Part I: hypodermic injection and strength improvement of apatite cement. Biomaterials 2004;25(11):2187e95. [65] Baroud G, Cayer E, Bohner M. Rheological characterization of concentrated aqueous b-tricalcium phosphate suspensions: the effect of liquid-to-powder ratio, milling time, and additives. Acta Biomater 2005;1(3):357e63. [66] Sarda S, Fernandez E, Nilsson M, Balcells M, Planell J. Kinetic study of citric acid influence on calcium phosphate bone cements as waterreducing agent. J Biomed Mater Res 2002;61(4):653e9. [67] Cherng A, Takagi S, Chow L. Effects of hydroxypropyl methylcellulose and other gelling agents on the handling properties of calcium phosphate cement. J Biomed Mater Res 1997;35(3):273e7. [68] Hesaraki S, Moztarzadeh F, Sharifi D. Formation of interconnected macropores in apatitic calcium phosphate bone cement with the use of an effervescent additive. J Biomed Mater Res 2007;83(1):80e7. [69] Bigi A, Bracci B, Panzavolta S. Effect of added gelatin on the properties of calcium phosphate cement. Biomaterials 2004;25(14):2893e9. [70] Liu C, Shao H, Chen F, Zheng H. Rheological properties of concentrated aqueous injectable calcium phosphate cement slurry. Biomaterials 2006;27(29):5003e13. [71] Ishikawa K, Miyamoto Y, Kon M, Nagayama M, Asaoka K. Non-decay type fast-setting calcium phosphate cement: composite with sodium alginate. Biomaterials 1995;16(7):527e32. [72] An J, Liao H, Kucko NW, Herber RP, Wolke JG, van den Beucken JJ, et al. Long-term evaluation of the degradation behavior of three apatiteforming calcium phosphate cements. J Biomed Mater Res 2016;104(5):1072e81. [73] Chow LC, Eanes ED. Octacalcium phosphate. Karger Medical and Scientific Publishers; 2001. [74] Dorozhkin SV. Calcium orthophosphate-based biocomposites and hybrid biomaterials. J Mater Sci 2009;44(9):2343e87. [75] Canal C, Ginebra M. Fibre-reinforced calcium phosphate cements: a review. J Mech Behav Biomed Mater 2011;4(8):1658e71. [76] Zhang J, Liu W, Schnitzler V, Tancret F, Bouler J-M. Calcium phosphate cements for bone substitution: chemistry, handling and mechanical properties. Acta Biomater 2014;10(3):1035e49. [77] Morgan JP, Dauskardt RH. Notch strength insensitivity of self-setting hydroxyapatite bone cements. J Mater Sci Mater Med 2003;14(7): 647e53. [78] Geffers M, Groll J, Gbureck U. Reinforcement strategies for load-bearing calcium phosphate biocements. Materials 2015;8(5):2700e17. [79] Kru¨ger R, Groll J. Fiber reinforced calcium phosphate cementseOn the way to degradable load bearing bone substitutes? Biomaterials 2012; 33(25):5887e900. [80] Wang J, Liu C, Liu Y, Zhang S. Double-network interpenetrating bone cement via in situ hybridization protocol. Adv Funct Mater 2010; 20(22):3997e4011. [81] Dos Santos LA, Carrodeguas RG, Boschi AO, De Arruda AC. Dual-setting calcium phosphate cement modified with ammonium polyacrylate. Artif Organs 2003;27(5):412e8. [82] Engstrand J, Persson C, Engqvist H. The effect of composition on mechanical properties of brushite cements. J Mech Behav Biomed Mater 2014;29:81e90. [83] Hofmann M, Mohammed A, Perrie Y, Gbureck U, Barralet J. High-strength resorbable brushite bone cement with controlled drug-releasing capabilities. Acta Biomater 2009;5(1):43e9. [84] Barralet J, Gaunt T, Wright A, Gibson I, Knowles J. Effect of porosity reduction by compaction on compressive strength and microstructure of calcium phosphate cement. J Biomed Mater Res 2002;63(1):1e9. [85] Chow L, Hirayama S, Takagi S, Parry E. Diametral tensile strength and compressive strength of a calcium phosphate cement: effect of applied pressure. J Biomed Mater Res 2000;53(5):511e7. [86] Gbureck U, Spatz K, Thull R, Barralet J. Rheological enhancement of mechanically activated a-tricalcium phosphate cements. J Biomed Mater Res B Appl Biomater 2005;73(1):1e6. [87] Barralet JE, Tremayne M, Lilley KJ, Gbureck U. Modification of calcium phosphate cement with a-hydroxy acids and their salts. Chem Mater 2005;17(6):1313e9. [88] Greish Y, Brown P, Bender J, Allcock H, Lakshmi S, Laurencin C. Hydroxyapatiteepolyphosphazane composites prepared at low temperatures. J Am Ceram Soc 2007;90(9):2728e34. [89] Christel T, Kuhlmann M, Vorndran E, Groll J, Gbureck U. Dual setting a-tricalcium phosphate cements. J Mater Sci Mater Med 2013;24(3): 573e81. [90] Geffers M, Barralet JE, Groll J, Gbureck U. Dual-setting brushiteesilica gel cements. Acta Biomater 2015;11:467e76.

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[91] Alkhraisat MH, Rueda C, Jerez LB, Marin˜o FT, Torres J, Gbureck U, et al. Effect of silica gel on the cohesion, properties and biological performance of brushite cement. Acta Biomater 2010;6(1):257e65. [92] Xu HH, Eichmiller FC, Giuseppetti AA. Reinforcement of a self-setting calcium phosphate cement with different fibers. J Biomed Mater Res 2000;52(1):107e14. [93] Ritchie RO. The conflicts between strength and toughness. Nat Mater 2011;10(11):817e22. [94] Brandt AM. Cement-based composites: materials, mechanical properties and performance. CRC Press; 2009. [95] Aral A, Yalc¸in S, Karabuda ZC, Anil A, Jansen JA, Mutlu Z. Injectable calcium phosphate cement as a graft material for maxillary sinus augmentation: an experimental pilot study. Clin Oral Implants Res 2008;19(6):612e7. [96] Sugawara A, Fujikawa K, Kusama K, Nishiyama M, Murai S, Takagi S, et al. Histopathologic reaction of a calcium phosphate cement for alveolar ridge augmentation. J Biomed Mater Res 2002;61(1):47e52. [97] Shirakata Y, Oda S, Kinoshita A, Kikuchi S, Tsuchioka H, Ishikawa I. Histocompatible healing of periodontal defects after application of an injectable calcium phosphate bone cement. A preliminary study in dogs. J Periodontol 2002;73(9):1043e53. [98] Lee S-K, Lee S-K, Lee S-I, Park J-H, Jang J-H, Kim H-W, et al. Effect of calcium phosphate cements on growth and odontoblastic differentiation in human dental pulp cells. J Endod 2010;36(9):1537e42. [99] Ryf C, Goldhahn S, Radziejowski M, Blauth M, Hanson B. A new injectable brushite cement: first results in distal radius and proximal tibia fractures. Eur J Trauma Emerg Surg 2009;35(4):389e96. [100] Ferguson SJ, Steffen T. Biomechanics of the aging spine. Eur Spine J 2003;12(2):S97e103.

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C H A P T E R

35 Biologic Scaffolds Composed of Extracellular Matrix for Regenerative Medicine Michelle Scarritt, Mark Murdock, Stephen F. Badylak University of Pittsburgh, Pittsburgh, PA, United States

INTRODUCTION The goal of regenerative medicine is to rebuild or replace damaged tissues and enable structural and functional restoration of the tissue. Many regenerative medicine strategies involve the use of scaffolds composed of synthetic polymers or biologic materials. The purposes of these scaffolds are to provide at least temporary mechanical support and to act as a temporal and spatial guide to promote cell attachment, migration, proliferation, and/or differentiation during the development of site-specific neotissues. Synthetic scaffold materials have a known chemistry that permits controlled chemical and structural modification such that they can be manufactured to almost any physical and mechanical specifications. However, synthetic materials invariably elicit a proinflammatory host response and occasionally can elicit a foreign body response. Bioscaffolds composed of extracellular matrices (ECM) as well as those composed of purified individual components of ECM (such as collagen) have also been used as surgical mesh materials. Numerous biologic scaffolds have been used in human clinical applications to treat a variety of tissue defects (Table 35.1). Biologic scaffolds are not as fine-tunable as synthetic scaffolds; however, biologic materials are typically more biocompatible and have inductive properties that can modulate cell behavior [1]. This chapter discusses the production and application of biologic scaffold materials for regenerative medicine applications with a focus on intact acellular ECM scaffolds.

EXTRACELLULAR MATRIX: FUNCTION AND COMPONENTS ECM consists of a complex mixture of structural macromolecules and bioactive factors which not only act as scaffolding for cellular growth and motility, which gives shape to tissues, but as a bioactive, instructional environment that can regulate multiple facets of cell behavior including adhesion, migration, differentiation, proliferation, and even survival. All multicellular organisms produce ECM although the composition can vary greatly between species. Plant ECM, for example, includes components of the cell wall, such as the polysaccharide cellulose. Chitosan, a biopolymer that has abundant commercial and biomedical applications, is derived from chitin, a nitrogenous polysaccharide that is a component of the exoskeleton of insects and crustaceans, the radulae of mollusks, the scales of fish and amphibians, and the beaks or shells of cephalopods [2]. Although these polysaccharides can be considered naturally occurring scaffold materials, they will not be discussed in this chapter. Here, the focus is on mammalian ECM, which is highly conserved among species and includes the basement membrane and the interstitial matrix between cells. The basement membrane is a thin, fibrous, sheet-like matrix that underlies an epithelium, mesothelium, or endothelium. The interstitial matrix, on the other hand, is a gel consisting of proteins and polysaccharides that surround cells and act as a compression buffer. Resident cells produce and assemble the components of the ECM intracellularly and then secrete them into the extracellular space by exocytosis. Once secreted, these components assemble to Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00035-7

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TABLE 35.1

Examples of Extracellular MatrixeBased Products

Product TM

AlloDerm

Company

Material

Processing

Applications

LifeCell Corporation

Human dermis

Natural, dry sheet

Soft tissue repair

TM

AlloMax

Becton, Dickinson and Company

Human dermis

Natural, dry sheet

Soft tissue repair

AlloPatchÒ

Musculoskeletal Transplant Foundation

Human dermis

Natural, dry sheet

Tendon and soft tissue repair

ArthroFlexÒ

Arthrex

Human dermis

Preservon proprietary preservation process

Soft tissue repair

AxisTM

Coloplast

Human dermis

Natural, dry sheet

Pelvic organ prolapse and stress urinary incontinence

AxoGuardÒ

AxoGen

Porcine SIS

Multilaminar sheets

Nerve repair

BellaDerm

MTF Biologics

Human dermis

Natural, dry sheet

Repair integumental tissue or soft tissue

BiodesignÒ

Cook Biotech

Porcine SIS

Natural, dry sheet

Dura mater repair

CorMatrixÒ

Cook Biotech

Porcine SIS

Natural, dry sheet

Cardiac tissue repair

DermACELLÒ

Stryker

Human dermis

Preservon proprietary preservation process

Soft tissue repair

DermaMatrixTM

DePuy Synthes

Human dermis

Natural, dry sheet

Repair integumental tissue or soft tissue

DuraGuard

Baxter

Bovine pericardium

Cross-linked, hydrated sheet

Dura mater and soft tissue repair

DuraMatrixÒ

Stryker

Bovine dermis

Natural, dry sheet

Dura mater repair

FlexHDÒ

Ethicon

Human dermis

Natural, hydrated sheet

Soft tissue repair

GlyadermÒ

Euro Skin Bank

Human dermis

Preserved in 85% glycerol

Dermal repair and plastic surgery

GraftJacketTM

Wright Medical Group

Human dermis

Natural, dry sheet

Tendon and ligament repair

Integra HuMendTM

Integra LifeSciences

Human dermis

Natural, dry sheet

Soft tissue reconstruction

IntegraÒ Reinforcement Matrix

Integra LifeSciences

Porcine dermis

Natural, dry sheet

Soft tissue reconstruction and tendon repair

MIRODERMÒ

MiroMatrix Medical Inc

Porcine liver

Sheet cut from decellularized organ

Wound repair

MIROMESHÒ

MiroMatrix Medical Inc

Porcine liver

Sheet cut from decellularized organ

Soft tissue repair

OASISÒ Wound Matrix

Smith & Nephew

Porcine SIS

Natural, dry sheet

Wound repair

Peri-GuardÒ

Baxter

Bovine pericardium

Cross-linked, dry sheet

Pericardial and soft tissue repair

PerioDermTM

MTF Biologics

Human dermis

Natural, dry sheet

Dental, integumental, and soft tissue repair

PermacolTM

Medtronic

Porcine dermis

Cross-linked sheet

Soft tissue repair

PriMatrixÒ

Integra LifeSciences

Fetal bovine dermis

Natural, dry sheet

Wound repair

SurgiMendÒ

Integra LifeSciences

Fetal bovine dermis

Natural, dry sheet

Soft tissue repair

SuspendÒ

Coloplast

Human fascia lata

Natural, dry sheet

Prolapse and stress urinary incontinence

615

EXTRACELLULAR MATRIX: FUNCTION AND COMPONENTS

TABLE 35.1

Examples of Extracellular MatrixeBased Productsdcont’d

Product

Company

Material

Processing

Applications

TissueMendÒ

Stryker

Fetal bovine dermis

Natural, dry sheet

Tendon repair

Vascu-GuardÒ

Synovis Surgical

Bovine pericardium

Cross-linked, hydrated sheet

Vascular reconstruction

VeritasTM

Baxter

Bovine pericardium

Cross-linked, hydrated sheet

Soft tissue repair

Boston Scientific

Fetal bovine dermis

Natural, dry sheet

Soft tissue repair

Becton, Dickinson and Company

Porcine dermis

Natural, dry sheet

Soft tissue repair

Zimmer Biomet

Porcine dermis

Cross-linked, hydrated sheet

Tendon repair

XenformTM XenMatrix

TM

Zimmer Collagen Repair PatchÒ SIS, small intestinal submucosa.

form a fibrous mesh. The composition of ECM is tissue-specific, highly dynamic, and crucially important in organ and tissue development, homeostasis, and response to injury. Each component of the ECM participates in the continuous cross-talk between cells and their environment. This general mechanism is known as dynamic reciprocity and is capable of influencing gene expression [3] (Fig. 35.1). The ECM consists of proteins, glycosaminoglycans (GAGs), glycoproteins, and small molecules that function individually and in conjunction to support and influence cell behavior. The main components are discussed in more detail subsequently; there are excellent textbooks and reviews that describe the ECM in depth [4].

Collagen Collagen is the most abundant protein in the human body. Fittingly, it is also the most common biologic scaffold material. The main function of collagen is to give structural support to resident cells. The domain structure of collagen is unique: a triple helix consisting of three distinct a-chains called “tropocollagen” [5]. Proper threedimensional (3D) folding requires a glycine residue in every third position of the individual a-chains with prolines and hydroxyprolines often flanking the glycines. The helical structure is further stabilized via hydrogen bonds between the chains such that the resulting structure is highly stable. The compliance of a tissue containing collagen depends on the form and arrangement of collagens as well as the degree of mineralization [6]. Accordingly, the 28 types of collagen have been categorized into several structural classes including fibrillar (types I, II, III, V, and XI), fibril-associated collagens with interrupted triple helices (types IX, XII, XIV, XVI, and XIX), short-chain (types VIII and X), basement membrane (type IV), multiple triple helix domains with interruptions (types XV and XVIII), membrane-associated collagens with interrupted triple helices (types XIII

FIGURE 35.1 ECMecell interactions are a two-way street. Although the ECM originates from the resident cells, the relationship between the ECM and cells is a feedback loop. Dynamic reciprocity, a mechanism whereby ECM could influence gene expression which in turn would modify and influence cell-secreted products, was proposed in 1982 by Mina Bissell [3a]. Multiple groups have since demonstrated that the ECM is able to affect cell signaling directly not only by acting as a reservoir for growth factors and other cell-secreted molecules, but also directly through mechanotransduction. Cells are constantly degrading, secreting, and remodeling their environment and thus dictate the composition, mechanics, and geometry of the ECM. In turn, the ECM is able to affect the morphology, polarity, and phenotype of cells, which ultimately affects cell and tissue function. ECM, extracellular matrix.

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35. BIOLOGIC SCAFFOLDS COMPOSED OF EXTRACELLULAR MATRIX FOR REGENERATIVE MEDICINE

and XVII), and other (types VI and VII) [7]. Type I collagen accounts for over 90% of the collagen in the human body and is found in fibrous tissues such as skin and tendons. Fibroblasts are the most common cell type that produces collagen. In addition to its structural properties, collagen has inherent functional properties such as the stimulation or inhibition of angiogenesis and the promotion of cellular proliferation, differentiation, and attachment [8]. For these reasons, collagen is often used as a coating for cell culture vessels not only to promote cell attachment but also to influence cell phenotype. As the main component of connective tissue, collagen is a highly conserved protein that is ubiquitous among mammalian species and accounts for approximately 25%e35% of all body proteins [9]. Inherent common amino acid sequences and epitope structures exist within collagen molecules across species [10,11]. These common antigens appear to account for the lack of an adverse immune response when xenogeneic collagen is used as an implantable scaffold material. Upon implantation, if left in its native ultrastructure, collagen implants are subjected to the fundamental biological processes of degradation and integration into adjacent host tissues. However, structural modifications such as chemical cross-linking may retard or prevent degradation and integration. Collagen can be extracted from tissues such as tendons and ligaments, solubilized, and then reconstituted into fine strands that then can be fashioned to mimic body structures such as heart valves, blood vessels, and skin. Although collagen provides considerable mechanical strength in its natural state, the necessary mechanical and physical properties of tissue engineered products for cardiovascular and orthopedic use often require chemical manipulation of collagen-based materials. Extracted and reconstituted collagen is usually stabilized by chemical cross-linking methods and then terminally sterilized by irradiation or ethylene oxide treatment before clinical use. Chemical processing methods include glutaraldehyde treatment, carbodiimide treatment, dye-mediated photooxidation, exposure to polyepoxy compounds, and glycerol treatment. Most methods of chemical cross-linking increase the strength of collagen at the cost of decreasing the rate of in vivo degradation and negatively affecting cell attachment, proliferation, differentiation, and tissue remodeling. Exposure to chemical cross-linking agents can also change the biocompatibility of a collagen-based material, leading to a foreign body response. Bovine and porcine type I collagen provide readily available sources of ECM scaffold material for numerous clinical applications. Isolated collagen and collagen scaffolds have been widely used in cosmetic surgery, bone grafts, artificial skin for burns, wound care, and multiple other reconstructive applications. Examples of collagen scaffolds include Contigen (Becton, Dickinson and Company, Franklin Lakes, NJ), CosmoDerm and CosmoPlast (Allergan, Inc., Santa Barbara, CA), CollaGUARD/Collieva and CollaRx (Innocoll, Inc., Ashburn, VA), Chondro-Gide (Geistlich Pharma AG, Wolhusen, Switzerland), and Menaflex (formerly Collagen Meniscal Implant [CMI]; ReGen Biologics, Inc., Franklin Lakes, NJ). For these reasons, collagen has become a favorite substrate for many tissue engineering and regenerative medicine applications. The tissue and species source of collagen and its treatment before use are important considerations in the design of tissue engineered devices.

Fibronectin There are two forms of fibronectin in animals: soluble plasma fibronectin and insoluble cellular fibronectin. The soluble form is produced by the hepatocytes of the liver and is a major component of blood. Deposition of plasma fibronectin at a wound site leads to the formation of a blood clot, which is critical to proper wound healing [12]. The insoluble form is a glycoprotein within the ECM. This form of fibronectin is secreted by cells, mainly fibroblasts, as a soluble protein dimer that is subsequently clustered with other fibronectins and arranged into an insoluble fibril matrix. Fibronectin is a protein dimer consisting of two nearly identical approximately 250-kDa polypeptide chains linked by a pair of C-terminal disulfide bonds [13]. Each monomer consists of repeating units termed types I, II, and III [14]. Type I repeating units are approximately 40 amino acids long; they contain two disulfide bonds and appear 12 times in a single fibronectin. Type II units are approximately 60 amino acids long; they contain two disulfide bonds and appear twice in a fibronectin molecule. Type III units are 90 amino acids long; they contain no disulfide bonds and appear 15e17 times. The lack of disulfide bonds in type III units permits partial unfolding under applied force [15]. All three types form a b sandwich composed of two antiparallel b-sheets. Although a single gene encodes fibronectin, a number of splice variants exist, resulting in at least 20 variants in humans [16]. Splicing can occur within type III repeats, leading to “extra” units termed EIIIA (between repeats III11 and III12) and EIIIB (between III7 and III8). Between III14 and III15 is a nonhomologous stretch of amino acids of variable length called the V-region, which represents another site for alternative splicing. This V-region can be partially or completely excluded, which subsequently leads to the exclusion of an integrin a4b1 binding domain. These repeating units

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and splice variants are arranged into several functional domains that permit fibronectin to bind to collagen, heparin, chondroitin sulfate, heparan sulfate, fibrin, other fibronectins, and integrins a5b1, avb1, avb3, a5b6, a3b1, a8b1, aIIbb3, and a4b1. Integrins are transmembrane receptors that facilitate cellecell and cellematrix interactions and can relay information to the cell about the composition and mechanics of the ECM. The sequence arginine-glycine-aspartate (known as the RGD sequence) is located within III10 and is the site for cell attachment via a5b1 and aVb3 integrins. Thus, fibronectin is involved in cell adhesion, migration, differentiation, and growth. Because of its cell-binding RGD sequence, recombinant fibronectin is often used to coat culture vessels to promote cell attachment and proliferation [17]. For these reasons, fibronectin has been conjugated to porous 3D poly(carbonate) urethane scaffolds for vascular tissue engineering and poly(D,L-lactide-co-glycolide) scaffolds for periodontal tissue engineering, among others [18,19]. One interesting preclinical study used modified fibronectin fragments to bind soluble growth factors and integrins simultaneously; injection of these growth factoreloaded fragments was able to improve wound healing and bone regeneration [20].

Laminin Laminins are highemolecular weight, heterotrimeric proteins that are found in the basement membrane, specifically the basal laminae. Laminins contain an a-chain, a b-chain, and a g-chain. There are five genetic variants of the a-chain, four of the b-chain, and three of the g-chain. To date, 15 heterotrimers resulting from unique combinations of these chain variants have been identified in vivo, 12 of which were identified in mammals [21]. Unlike collagen, laminins do not form fibers. Instead, they self-assemble and self-bind to form independent mesh-like structures. Laminins also associate with nidogens and collagens via entactin, fibronectin, and perlecan. Through these intermolecular bonds, a network is formed that allows the basal lamina to resist tensile forces. Heparan sulfates, b1 and b4 integrins, dystroglycan, and other surface receptors are able to bind directly to laminin moieties. The binding of laminin to these receptors and receptor-like proteins facilitates cell adhesion, migration, and differentiation [21]. With regard to the use of laminin as a biologic scaffold, laminins are a critical component of matrix materials generated from the basement membrane of tissues via decellularization, a topic that will be discussed later in the chapter. On their own, laminins have been used as a substrate for culturing cells in vitro. Two groups demonstrated that mouse embryonic stem cells can be grown on recombinant laminin-511 for long-term culture [22,23]. Human embryonic stem cells and induced pluripotent stem cells have also been grown with recombinant laminin 511 without the need for a feeder layer [24].

Elastin As its name implies, elastin lends elasticity to tissues so that it can return to its original shape after stretching or contracting. Elastin is particularly prevalent in the ECM of the lungs, skin, ligaments, and blood vessels, where physiological forces result in intermittent stretching. Similar to fibronectin, elastin is synthesized from only one gene, but it can undergo extensive alternative splicing, resulting in tissue-specific elastin variants. Fibroblasts and smooth muscle cells secrete a soluble, nonglycosylated, hydrophobic form of elastin known as tropoelastin. These elastin precursors undergo posttranslational modification and are packaged inside chaperone molecules and secreted. In the extracellular space, tropoelastins covalently bind by patterning alternating hydrophobic and hydrophilic sequences to form a highly insoluble, cross-linked array, a process known as coacervation [25]. The tropoelastin coacervates are associated with microfibrils, particularly fibrillin-1, by the molecules fibulin-5 and fibulin-4 [26]. Lysyl oxidases cross-link the tropoelastins; after the formation of a multitude of cross-links between the tropoelastin monomers, the resulting elastin polymer becomes an insoluble fiber [27]. Electrospun recombinant human tropoelastin has been evaluated for use in wound healing [28]. Commercially, it is marketed as a product called Dermalastyl or Elastatropin for use in skin care to prevent wrinkles.

Glycosaminoglycans/Proteoglycans GAG are long unbranched polysaccharides consisting of a repeating disaccharide unit. GAG are classified into four groups based on their core disaccharide: heparin/heparin sulfate, chondroitin sulfate/dermatan sulfate, keratan sulfate, and hyaluronic acid (HA). All GAG are synthesized in the Golgi apparatus except HA, which is synthesized at the cell membrane. Unlike protein or nucleic acid synthesis, GAG production is not template driven. GAG

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are highly heterogeneous because they are subjected to dynamic modification by processing enzymes [29]. These modifications include sulfation and epimerization, which make GAG highly polar and allow them to attract water molecules and bind cations such as calcium, potassium, and sodium. Thus, GAG are able to act as a physiological lubricant or shock absorber. One or more GAG covalently bonded to a protein are considered a proteoglycan. Proteoglycans are heavily glycosylated proteins. The protein core of the proteoglycan is synthesized in the rough endoplasmic reticulum and then posttranslationally modified with glycosylations in the Golgi apparatus. A tetrasaccharide bridge connects a GAG to a serine residue of a protein core. Chondroitin sulfate is the most prevalent GAG. It is usually found as part of a proteoglycan, and its function largely depends on the protein to which it is attached. As part of a large aggregating proteoglycan, such as aggrecan, neurocan, versican, or brevican, chondroitin sulfate is critical to maintaining the structural integrity of a tissue. Aggrecan is a major component of cartilage in which chondroitin sulfate lends its highly charged sulfate groups to attract water and provide electrostatic repulsion to generate resistance to compression. Chondroitin can be extracted from the cartilage of cow and pig tissues for use as a biologic. Thus far, it has been investigated as a dietary supplement, drug, or injection in clinical trials to alleviate osteoarthritis; however, the symptomatic benefit is reported to be minimal [30]. Dermatan sulfate is distinguished from chondroitin sulfate by the presence of iduronic acid instead of glucuronic acid. Like chondroitin sulfate, dermatan sulfate is found in proteoglycans such as decorin, biglycan, and versican. Dermatan sulfate is mostly found in skin, but it is also present in tendons, blood vessels, heart valves, intestinal mucosa, and lungs, where it may have a role in coagulation, wound repair, response to infection, differentiation, morphogenesis, cell migration, carcinogenesis, and fibrosis [31]. Although the clinical application of dermatan sulfate is limited, companies such as Bioibe´rica (Barcelona, Spain) manufacture and sell it for potential use in enhancing wound repair or preventing coagulation. Keratan sulfate is found in cornea, bone, nervous tissue, and cartilage, where it is part of proteoglycans such as lumican, mimecan, aggrecan, osteoadherin, keratocan, and fibromodulin. Like other GAG, it is highly hydrated and is able to absorb shock experienced by joints. Keratan sulfate has been demonstrated to be absent or aberrant in macular corneal dystrophy [32]; however, keratan sulfate has no widespread clinical uses. Heparan sulfate is found in all animal tissues in various proteoglycan forms. Transmembrane heparan sulfate proteoglycans such as syndecans and glypicans are critical to cell signaling processes. Other proteoglycans are integral to the basement membrane, such as perlecan and agrin. Heparan sulfate can also be bound to a variety of extracellular proteins including collagen XVIII, growth factors, chemokines, cytokines, enzymes, and coagulation factors. Thus, heparan sulfate has a critical role in regulating developmental and biologic processes. Largely based on its ability to bind growth factors and cytokines, heparan sulfate and its mimics have been reported to promote wound repair and tissue regeneration [33,34]. Unlike other GAG, HA (also known as hyaluronan) is a polysaccharide that is not sulfated and is not found as a proteoglycan. HA is synthesized by integral membrane hyaluronan synthases that directly extrude it into the extracellular space [35]. HA is able to absorb significant amounts of water to act as a swelling force that permits tissues to resist compression. Thus, it is most commonly found in load-bearing joints and in the ECM of tissues such as skin, cartilage, connective tissue, epithelial tissue, and neural tissue [36,37]. HA is also found in nonanimal sources such as roots and tubers (potatoes and sweet potatoes) [38]. In addition to mechanical support, HA influences cell proliferation and migration. Moreover, HA facilitates these cellular functions during the late stages of wound healing and acts as a promoter of inflammation in the early stages [39]. HA has been extensively investigated as a natural scaffold material for tissue reconstruction. It has been used clinically in an injectable form for treatment of osteoarthritis of the knee, for cosmetic surgery as a dermal filler, and for ophthalmic surgery, among other applications. Examples of HA-based materials include Belotero (Merz Aesthetics, Raleigh, NC), JUVE´DERM (Allergan, Inc.), Hyalomatrix, Hyalofill, Hyalogran, Hyalosafe, and Elevess (Anika Therapeutics, Inc., Bedford, MA), and Restylane and Restylane Perlane (Q-Med Corporation, Fort Lauderdale, FL).

Matrix-Bound Nanovesicles Matrix-bound vesicles (MBV) are a component of the ECM. They were first identified by Huleihel et al. as nanovesicles embedded within ECM scaffolds derived from porcine urinary bladder matrix (UBM), small intestinal submucosa (SIS), and dermis [40]. Like exosomes, MBV are nanosized, lipid membraneebound vesicles with a size distribution from 10 to 200 nm. MBV were isolated from acellular porcine tissues by enzymatic digestion followed

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by ultracentrifugation at increasing g values. MBV contain distinct profiles of microRNAs and proteins with some mutual cargo among the three ECM scaffolds evaluated. In vitro, MBV can recapitulate some of the regenerative effects observed when cells are treated with ECM, which highlights the potential of MBV as a key mechanism mediating the regenerative abilities of ECM scaffolds. For example, Huleihel et al. observed neurite outgrowth in neuroblastoma cells treated with MBV and activation of macrophages toward a phenotype associated with a more regenerative, anti-inflammatory, constructive remodeling response. MBV may be useful for direct injection to a damaged site to encourage local anti-inflammatory and pro-remodeling pathways, or as an additive to enhance implanted scaffold performance. Identification of the variability and potential modification of MBV cargo will be critical to developing clinically relevant strategies with MBV.

INTACT AND SOLUBILIZED EXTRACELLULAR MATRIX AS A SCAFFOLD MATERIAL Tissue Procurement Intact ECM can be isolated from virtually any tissue. Decellularization protocols have been established for heart valves, blood vessels, skin, nerves, skeletal muscle, tendons, ligaments, ovaries, testes, small intestine, urinary bladder, liver, and other tissues. ECM is routinely harvested for research purposes and for clinical application from different species including humans, pigs, cows, and horses [1]. ECM may alternatively be collected from cells grown in vitro [41]. ECM scaffold materials harvested from different species, tissues, or cells have unique structural, functional, and biochemical characteristics. For instance, ECM scaffolds derived from porcine SIS (SIS-ECM) consist of about 90% collagen, most of which is collagen type I, and minor amounts of other collagen types III, IV, V, and VI [42]. On the other hand, although they contain the same collagen types as SISeECM, ECM scaffolds composed of porcine UBM (UBM-ECM) contain greater amounts of Col III as well as Col VII, which are important components of the epithelial basement membrane [43]. ECM from various sources also differ in the amount and distribution of GAGs, including heparin, heparin sulfate, chondroitin sulfates, and HA [36,37]; adhesion molecules such as fibronectin and laminin [43,44]; the proteoglycan decorin and the glycoproteins biglycan and entactin [1]; as well as various growth factors including transforming growth factor-b [45], basic fibroblast growth factor (b-FGF) [46], and vascular endothelial growth factor (VEGF) [47]. In addition, ECM scaffolds are distinctive in their protein composition from location to location within various tissues: for example, endocrine versus exocrine loci within the pancreas, or, the valvular versus mural loci within the heart. It is assumed that preservation of the intact ECM composition as well as its intrinsic ultrastructure and 3D architecture, especially its collagen fiber architecture, are fundamentally important in processes such as cell recruitment, migration, proliferation, and differentiation during neotissue formation in vivo [43,48].

Decellularization Conceptually, intact ECM consists of all of the structural and functional components secreted by the resident cells in their native 3D microarchitecture. In reality, the content and microarchitecture of intact ECM are invariably altered during the decellularization process. ECM will necessarily lose some of its functional components as they are washed out by detergents or other liquid washes. The reagents used in these washes are occasionally retained to a small degree in the final ECM product. Moreover, various decellularization processes and reagents are known to alter the orientation, porosity, or fiber size of the main collagen network differentially, potentially affecting biocompatibility and bioactivity. Decellularization processes are designed to balance the removal of cellular material with the preservation of the composition, mechanical integrity, and biological activity of the remaining ECM. Decellularization processes attempt to achieve the metrics of decellularization established in 2011 by Crapo et al.: less than 50 ng of double-stranded DNA per milligram of dry weight ECM, no visible nuclei in hematoxylin eosinestained histological sections, and any residual DNA being fewer than 200 base pairs in length [49]. The process from native tissue to ECM varies dramatically between tissue types, considering factors such as tissue density, vasculature, and lipid content. Decellularization processes necessarily include multiple steps, i.e., liberation of desired tissues from surrounding tissues, decellularization (several methods are described subsequently), disinfection, lyophilization, and terminal sterilization. Most commercially available intact ECM scaffolds are processed into a sheet form before decellularization by methods that include trimming and spreading of the original tissue to facilitate the removal of cellular components and debris. In addition, the decellularization of whole organs has been achieved through

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perfusing detergents or other solutions through the tissue’s native vascular network [50]. Common approaches to decellularization include a combination of physical and chemical treatments, e.g., sonication, agitation, freeze-thawing, and washes with various proteolytic detergents and solvents (reviewed in Gilbert et al. [51]). The decellularization process effectively removes xenogeneic and allogeneic cellular antigens that may be recognized as foreign by the host and results in an adverse inflammatory response or overt immune-mediated rejection [52,53]. The main ECM components are highly conserved between species and are well-tolerated by xenogeneic recipients [10,11]. Over 10 million patients have received ECM-based materials in clinical settings, most of which are of xenogeneic origin, and no immune rejection complications have been reported. Residual amounts of DNA and certain immunogenic species-specific antigens, such as galactosyl-a-1,3-galactose (a-Gal epitope), have been shown to be present in ECM scaffolds but fail to activate complement or bind immunoglobulin M antibody, possibly owing to the small amount and widely scattered distribution of the antigen [54e57].

Postprocessing After decellularization, it is often desirable for the ECM to be processed further. Such processing may include lyophilization (freeze-drying) or vacuum pressing before terminal sterilization to avoid leaching of soluble factors (for example, VEGF and b-FGF), and extend the product’s shelf life. The production of multilaminate forms of the ECM also improves the device’s handling and allows various geometric constructs to be built, including tubes [58], cones [59], and multilaminate sheets [60,61]. Lyophilized scaffolds can be processed further to yield powdered ECM, solubilized ECM, or ECM in hydrogel form [62] for use as minimally invasive injectable scaffolds [63,64] or in combination with sheets of ECM to produce sheet-powder hybrid scaffolds. Whereas each of the processing steps will change the overall composition and structure of the prepared ECM compared with those found in vivo, intact ECM preparations retain a multitude of structurally and functionally active proteins [61,65]. ECM scaffolds that are not chemically cross-linked are rapidly degraded in vivo. Typically, 50% of a noncrosslinked SIS-ECM scaffold is degraded within one month postimplantation and the scaffold is usually completely degraded within a three month time frame, as demonstrated in the repair of a urinary bladder defect or the Achilles tendon [66,67]. ECM degradation leads to an initial decrease in overall strength during the early phase of in vivo remodeling, followed by an increase in strength caused by the deposition of site-specific ECM and the formation of functional site-appropriate neotissue by infiltrating cells in response to their experienced mechanical stresses [58,66,68,69]. Soluble factors within ECM scaffold materials (that is, growth factors, and the release of biologically active cryptic peptides resulting from degradation of the ECM material [8,70]) are thought to be directly involved in the processes of neotissue formation including angiogenesis, mononuclear cell infiltration, cell proliferation, cell migration, and cell differentiation [45,71,72]. The release of soluble factors along with the rapid degradation of the ECM appear to be essential processes for constructive remodeling to occur. This fact is highlighted by an altered remodeling profile in clinical applications using scaffolds that have been chemically cross-linked using glutaraldehyde, carbodiimide, or hexamethylene-diisocyanate, or nonchemical methods. Whereas chemical crosslinking can increase the mechanical strength and reduce the rate of scaffold degradation of an ECM scaffold, this modification results in the formation of a chronic, proinflammatory, foreign body type of tissue response and a reduced level of constructive remodeling [71,73]. Similar to the host response to synthetic polymer scaffolds, an adverse host response is heralded by a predominance of M1-like activated macrophages, a high level of proinflammatory cytokines, and the formation of a type 1 T helper (Th-1) (cell-mediated rejection) response [74]. In contrast, intact, nonchemically modified ECM scaffold materials show a host immune response characteristic of accommodation and integration, as demonstrated by the increased presence of M2-like activated macrophages resulting in low levels of proinflammatory cytokines and the establishment of a Th-2 type of response [75e77]. The initial host immune response to ECM scaffolds appears to be critically important in determining subsequent processes including scaffold degradation, release of matricryptic peptides, host cell recruitment, and angiogenesis, among others [8,72].

Hydrogels A form of ECM that is receiving increased scientific and clinical attention is the ECM hydrogel. Unlike powder or sheet forms, ECM hydrogels are injectable, which makes them minimally invasive; and by virtue of their liquid state at room temperature, they can fill irregularly shaped volumes before they self-assemble into a gel at body temperature. The mechanical stiffness as well as the polymerization time of the resulting gel is fine-tunable by altering the concentration of solubilized ECM in the solution, which makes it a physically versatile biomaterial. Commonly used terminal sterilization methods including g- or electron beam irradiation and ethylene oxide exposure have been shown to inhibit ECM hydrogel formation. However, supercritical CO2 sterilization may permit sterile ECM

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hydrogel formation [78]. ECM hydrogels have the benefit of retaining the inherent bioactivity of the native matrix despite the loss of its native 3D microarchitecture during the solubilization process [79]. Hydrogels are typically created using powdered ECM as a starting point and then enzymatically digesting the structural components until complete solubilization is achieved. Pepsin is popularly used to solubilize ECM, and it has been observed that proteolysis with other enzymes (papain, for example) can impede the polymerization behaviors of the resulting solution. Similar to solid ECM scaffolds, hydrogels will degrade under physiological conditions, typically within one month of implantation, depending on the source ECM and concentration, and it is supposed that this degradation is essential in achieving their favorable host response [79,80]. Many studies suggest that hydrogels from various source ECMs have vast potential to regenerate bone, adipose tissue, skeletal muscle, and even central nervous system tissue [81e83]. Ventrix, Inc. developed a hydrogel for cardiac repair and regeneration called VentriGel. To date, Ventrix is in Phase I clinical development for VentriGel.

Host Response A unique feature of intact ECM compared with other types of bioscaffolds is the favorable host response that it elicits. As stated previously, over 10 million patients have received xenogeneic ECM-based materials in clinical settings and no immune rejection complications have been reported. The lack of an adverse immune response to the use of xenogeneic collagen in implantable scaffold materials has been attributed to the common nature of amino acid sequences and surface epitopes among species [10,11] although work has shown the immunomodulation effects of ECM degradation products [70]. Allogeneic and xenogeneic collagen is subjected to the fundamental biological processes of degradation and integration into adjacent host tissues when left in its native ultrastructure. However, structural modifications designed to alter the rate of degradation and remodeling (e.g., cross-linking) may impair the desired healing and regenerative response. Beyond simply being tolerated by the host, intact ECM has been implicated in directing functional remodeling in many tissue types including skin [84], tendon and ligament [60], urinary bladder [85], skeletal muscle [86], heart [87], and esophagus [58]. Xenogeneic ECMs used clinically are typically of porcine origin; they are usually derived from the urinary bladder, small intestine, or dermis. In vitro, the biologic effect of different source ECMs can be diverse, which can be partly explained by the various methods used to decellularize different tissues [88]. In clinical settings, certain source ECMs are approved for specific applications and not others, although in a clinical trial for volumetric muscle loss including 13 patients, no distinguishable differences were found between patients receiving ECM from urinary bladder, small intestine, or dermis [89]. The exact mechanisms of action for ECM’s favorable host response are unknown, although several observations provide a partial explanation. ECM has been shown to attract and/or facilitate differentiation of endogenous stem or progenitor cells [72], enhance mitogenesis in certain cell populations [90], and shift the innate immune response away from inflammation and toward remodeling [73]. In the clinical trial mentioned earlier, it was noted that the ECM scaffolds attracted muscle and nerve progenitor cells to the wound area, thus assembling the key players for muscle regeneration. In another study, an ECM hydrogel was applied to the volumetric lesion of a stroke infarct and within 24 h the host cell infiltrate was characterized. Approximately 60% of the cells were found to be of brainderived phenotypes, including neural progenitor cells, which upon histological staining appeared to be differentiating into neuronal networks [81]. Another 30% of infiltrating cells were peripheral macrophages; importantly, the ECM hydrogel directed their phenotype away from the M1, classical proinflammatory activation and toward the M2, alternatively activated anti-inflammatory phenotype, putatively aiding in remodeling. It is known that ECM and their degradation products have various effects on other immune cells as well. For example, ECM components such as entactin, nidogen, and fragments of laminin, elastin, and type IV collagen have been shown to attract neutrophils, a microbicidal cell type [91,92]. It was also found that tumor necrosis factor-a linked to fibronectin affected the signal transduction and cell adhesive properties of CD4þ T cells, possibly helping to direct these immune cells toward inflammatory sites [93]. Another study showed that a 3D collagen matrix was essential for naive T cells to interface meaningfully with antigen-presenting dendritic cells, leading to signal induction and T-cell activation [94]. These studies illustrate how ECM bioscaffolds promote regenerative conditions and tissue development by priming the local immune response.

Whole-Organ Scaffolds In 2008, perfusion of detergents through an intact donor heart was used to generate a 3D heart ECM scaffold [50]. This detergent-perfusion decellularization technique has since been applied to liver, lung, pancreas, kidney, and even a whole limb [95e99]. Whole-organ engineering using porcine-derived organ scaffolds seeded with

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patient-derived stem cells provides an alternative to traditional transplantation that circumvents the need for immunosuppression and the shortage of organs available for traditional transplantation. Recellularization of organ scaffolds is the current focus of the field. To facilitate recellularization, bioreactors have been developed that can provide perfusion, monitoring, and organ-specific stimuli such as mechanical stretch, pressure gradients, or electric impulses (reviewed in Scaritt et al. [100]). These bioreactors permit the ex vivo culture of whole-organ scaffolds seeded with cells such as primary parenchymal cells, endothelial cells, or stem/progenitor cells. Bioengineered rodent organs that have been generated using this bioreactor-assisted recellularization method have been orthotopically implanted by several groups. These bioengineered organs were able to participate in organ function to a limited extent, but hemorrhage and/or edema occurred within the bioengineered organ as a result of incomplete recellularization [95,98]. Regardless of the challenges ahead, the promise and progress of whole-organ engineering have prompted companies such as United Therapeutics Corporation and Miromatrix Medical Incorporated to invest in developing whole-organ decellularization-recellularization technology toward clinical application using porcine organ scaffolds.

CLINICAL AND COMMERCIAL APPLICATIONS As the science behind the generation and characterization of ECM scaffolds have expanded, the applications for these scaffolds have exponentially increased. Most commercially available ECM scaffolds are used as a surgical mesh for soft tissue reconstruction or topical wound healing (Table 35.1). Preclinical research has demonstrated the promise of using ECM scaffolds (and recellularized ECM scaffolds) for cardiac patches, vascular cuffs, heart valve replacement, tracheal reconstruction, and potentially even organ transplantation, among numerous other uses.

Regulatory Considerations for Extracellular Matrix Scaffolds ECM scaffolds have a long history of commercial and clinical application. Most of these ECM scaffolds have been classified as a medical device by the US Food and Drug Administration (FDA). In 1997, the FDA created the Tissue Reference Group (TRG) to assist in categorizing products as a device, biologic, or combination, specifically in the realm of regulating human cells, tissues, and cellular- and tissue-based products (HCT/P). The TRG is composed of representatives from the FDA’s Center for Biologics Evaluation and Research, the Center for Devices and Radiological Health, the Office of Combination Products, and the Office of the Chief Counsel. The TRG continues to update recommendations on the classification of products. As of this writing, the TRG indicated that secreted or extracted human products such as collagen are not considered an HCT/P (http://www.fda.gov/ biologicsbloodvaccines/tissuetissueproducts/regulationoftissues/ucm152857.htm). In 2014, ground, defatted, decellularized adipose tissue was determined not to be an HCT/P because it is more than minimally manipulated. As a final example, in 2012, allogeneic, processed acellular dermis for breast tissue defects was determined not to be an HCT/P because it is used in a nonhomologous site. It is apparent that FDA regulation of commercial or clinical products is changing as the field of tissue engineering and regenerative medicine advances. For the present, it is hoped that new ECM scaffold-based products will be classified based on their predicates.

CONCLUSIONS Biologic scaffold materials represent one facet of the multiple strategies used in regenerative medicine to construct functional tissue. Biologic scaffolds are derived from mammalian tissues and thus contain countless structural and functional moieties that have been shown to be necessary for tissue development, homeostasis, and response to injury. Such scaffold materials are currently used in clinical medicine and their use will likely expand as regenerative medicine strategies evolve.

List of Acronyms and Abbreviations b-FGF Basic fibroblast growth factor DNA Deoxyribonucleic acid ECM Extracellular matrix FACITs Fibril-associated collagens with interrupted triple helices

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FDA Food and Drug Administration GAGs Glycosaminoglycan HA Hyaluronic acid HCT/P Human cells, tissues, and cellular- and tissue-based products MACITs Membrane-associated collagens with interrupted triple helices MBV Matrix-bound vesicle PLGA poly(D,L-lactide-co-glycolide) RGD Arginine-glycine-aspartate SIS-ECM Porcine small intestinal submucosa extracellular matrix TGF-b Transforming growth factor-b TRG Tissue Reference Group UBM-ECM Porcine urinary bladder matrix extracellular matrix VEGF Vascular endothelial growth factor

Glossary Decellularization The process by which cells are removed from a tissue or organ to isolate the extracellular matrix scaffold of the tissue or organ for use in engineering or regenerating new tissue Recellularization The process by which cells are delivered to and cultured within a tissue or organ extracellular matrix scaffold Xenogeneic Relating to tissues, cells, or materials belonging to different species.

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[25] Mithieux SM, Weiss AS. Elastin. Adv Protein Chem 2005;70:437e61. [26] Yamauchi Y, Tsuruga E, Nakashima K, Sawa Y, Ishikawa H. Fibulin-4 and -5, but not Fibulin-2, are associated with tropoelastin deposition in elastin-producing cell culture. Acta Histochem Cytochem 2010;43(6):131e8. [27] Maki JM, Sormunen R, Lippo S, Kaarteenaho-Wiik R, Soininen R, Myllyharju J. Lysyl oxidase is essential for normal development and function of the respiratory system and for the integrity of elastic and collagen fibers in various tissues. Am J Pathol 2005;167(4):927e36. [28] Machula H, Ensley B, Kellar R. Electrospun tropoelastin for delivery of therapeutic adipose-derived stem cells to full-thickness dermal wounds. Adv Wound Care 2014;3(5):367e75. [29] Silbert JE, Sugumaran G. Biosynthesis of chondroitin/dermatan sulfate. IUBMB Life 2002;54(4):177e86. [30] Reichenbach S, Sterchi R, Scherer M, Trelle S, Burgi E, Burgi U, et al. Meta-analysis: chondroitin for osteoarthritis of the knee or hip. Ann Intern Med 2007;146(8):580e90. [31] Trowbridge JM, Gallo RL. Dermatan sulfate: new functions from an old glycosaminoglycan. Glycobiology 2002;12(9):117re25r. [32] Klintworth GK, Meyer R, Dennis R, Hewitt AT, Stock EL, Lenz ME, et al. Macular corneal dystrophy. Lack of keratan sulfate in serum and cornea. Ophthalmic Paediatr Genet 1986;7(3):139e43. [33] Groah SL, Libin A, Spungen M, Nguyen KL, Woods E, Nabili M, et al. Regenerating matrix-based therapy for chronic wound healing: a prospective within-subject pilot study. Int Wound J 2011;8(1):85e95. [34] van Neck J, Tuk B, Barritault D, Tong M. Heparan sulfate proteoglycan mimetics promote tissue regeneration: an overview. In: Davies J, editor. Tissue regenerationdfrom basic biology to clinical application. InTech; 2012. [35] Schulz T, Schumacher U, Prehm P. Hyaluronan export by the ABC transporter MRP5 and its modulation by intracellular cGMP. J Biol Chem 2007;282(29):20999e1004. 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An investigation of the long-term bioactivity of endogenous growth factor in OASIS Wound Matrix. J Wound Care 2005;14(1):23e5. [47] Hodde JP, Record RD, Liang HA, Badylak SF. Vascular endothelial growth factor in porcine-derived extracellular matrix. Endothelium 2001; 8(1):11e24. [48] Sellaro TL, Ravindra AK, Stolz DB, Badylak SF. Maintenance of hepatic sinusoidal endothelial cell phenotype in vitro using organ-specific extracellular matrix scaffolds. Tissue Eng 2007;13(9):2301e10. [49] Crapo PM, Gilbert TW, Badylak SF. An overview of tissue and whole organ decellularization processes. Biomaterials 2011;32(12):3233e43. [50] Ott HC, Matthiesen TS, Goh SK, Black LD, Kren SM, Netoff TI, et al. Perfusion-decellularized matrix: using nature’s platform to engineer a bioartificial heart. Nat Med 2008;14(2):213e21. [51] Gilbert TW, Sellaro TL, Badylak SF. Decellularization of tissues and organs. Biomaterials 2006;27:3675e83. England. 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Natural anti-galactose alpha1,3 galactose antibodies delay, but do not prevent the acceptance of extracellular matrix xenografts. Transpl Immunol 2002;10(1):15e24. [58] Badylak SF, Vorp DA, Spievack AR, Simmons-Byrd A, Hanke J, Freytes DO, et al. Esophageal reconstruction with ECM and muscle tissue in a dog model. J Surg Res 2005;128(1):87e97. [59] Nieponice A, Gilbert TW, Badylak SF. Reinforcement of esophageal anastomoses with an extracellular matrix scaffold in a canine model. Ann Thorac Surg 2006;82(6):2050e8. [60] Dejardin LM, Arnoczky SP, Ewers BJ, Haut RC, Clarke RB. Tissue-engineered rotator cuff tendon using porcine small intestine submucosa. Histologic and mechanical evaluation in dogs. Am J Sports Med 2001;29(2):175e84. [61] Freytes DO, Badylak SF, Webster TJ, Geddes LA, Rundell AE. Biaxial strength of multilaminated extracellular matrix scaffolds. Biomaterials 2004;25(12):2353e61.

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36 Hydrogels in Regenerative Medicine David F. Williams Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

INTRODUCTION To discuss the role of hydrogels in regenerative medicine without extraneous and irrelevant information, it is necessary to set out the meaning and boundaries of regenerative medicine itself. The author of the current described regenerative medicine in 2014, especially in the context of the use of biomaterials in medical technologies [1]. Regenerative medicine addresses a major nonpharmacological, approach to treat disease and injury through the de novo development of functional tissue rather than the replacement of tissue through synthetic devices. Although implantable medical devices can give good performance in treating many conditions, they will always be limited to situations that involve mechanical or physical functions and will not in themselves be able to provide biological solutions to replacing tissue structure and function. Thus, the techniques of regenerative medicine rely on tools that result in regenerating the patient’s own tissue; regenerative medicine involves alternative therapies to treat disease and injury by the regeneration of functional tissues or organs instead of replacement by medical devices, transplantation of viable structures or palliative care through the use of pharmaceuticals. As depicted in Fig. 36.1, there are three main strands of regenerative medicine. The first, usually called cell therapy, involves using groups of cells derived from the patient or elsewhere, which can be injected or otherwise placed at the site of disease or injury in the expectation that they will facilitate the spontaneous regeneration of the required tissue. At its conceptually simplest level, cell therapy does not involve conventional biomaterials, but there are situations in which they may have a supportive role; as will be shown, hydrogels feature prominently here. The second strand is that of gene therapy, in which specific genes are inserted into specific cells to correct deficiencies in those cells, enabling certain processes in tissue expression. It is possible for some biomaterials to be involved in the delivery process, and gene and cell therapies may be combined. The third strand is that of tissue engineering, “the creation of new tissue by the deliberate and controlled stimulation of selected target cells through a systematic combination of molecular and mechanical signals” [2]. Molecular and mechanical signals do not directly imply that the use of biomaterials and tissue engineering may theoretically be carried out in the absence of biomaterials. However, three factors determine that biomaterials are most likely to be involved in tissue engineering processes. The first is that new tissue generated in this way usually needs form and structure, and in themselves, injected cells are unlikely to provide this without the assistance of biomaterials. Second, molecular signals are easily delivered with the appropriate spatial and temporal characteristics; a biomaterial that contains and delivers such signals to the required cells would be beneficial. Third, mechanical signals may be equally difficult to deliver without the sustained effects of a biomaterial support. As implied earlier, the tissue engineering, cell therapy, and gene therapy modes have some overlapping features; it is not the intention here to become involved in the semantics of these terms or the details of interfaces between them. Instead, this chapter focuses on the role of hydrogels as the preferred exemplars of biomaterials that may be used to support the mechanisms of regenerative medicine and specifically facilitate the delivery of the mechanical and molecular signals, and to assist in generating new functional tissues with appropriate morphological characteristics.

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Tissue Engineering The creation of new tissue for the therapeutic reconstruction of the human body, by the deliberate and controlled stimulation of selected target cells through a systematic combination of molecular and mechanical signals

Regenerative Medicine Alternative therapies to treat disease and injury by the regeneration of functional tissue / organ structures instead of replacement by medical devices or palliative care through pharmaceuticals

Gene Therapy The technique for correcting defective genes responsible for disease development in which copies of a gene are inserted into living cells to replace abnormal, disease-causing gene.

May be combined

Cell Therapy The process of introducing new cells into a tissue in order to treat a disease.

FIGURE 36.1 The essence of regenerative medicine and its components.

BIOMATERIALS TEMPLATES Nevertheless, it is necessary to briefly address the principles underlying the specifications for the biomaterials that can support regenerative medicine. Specifically focusing on tissue engineering, it has been common practice to describe these material constructs as scaffolds [3]. Conventional scaffolds tend to be composed of discrete porous constructs, usually of polymers or ceramics, in which appropriate cells infiltrate the pores and are intended to express new tissue within these spaces, with the biomaterial degrading and resorbing at the same time. Such constructs have usually been produced by three-dimensional (3D) techniques such as solid free-form fabrication, electrospinning, and solvent casting with porogen leaching. Typical porous scaffolds are seen in Fig. 36.2.

(A)

(B)

FIGURE 36.2 Examples of conventional porous solid polymer tissue engineering templates.

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As discussed by Edalat et al. [4], a scaffold is required to provide an environment, or niche, that favors the natural behavior of cells. The in vivo microenvironment of a cell in general is composed of the relevant extracellular matrix (ECM), homotypic or heterotypic cells surrounding that cell, and cytokines and other bioactive agents around the cells associated with endocrine, autocrine, and paracrine secretions. The microenvironment also involves topographical and architectural features and mechanical forces. It immediately becomes obvious that the type of porous scaffold epitomized by Fig. 36.2 will have considerable difficulty in replicating this type of microenvironment. This becomes even more significant when stem cells are involved, where the microenvironment is referred to as the stem cell niche. Fig. 36.3, reproduced from Scadden [5], indicates the factors that participate in regulating the system of a stem cell in its tissue state. These include “the constraints of the architectural space, physical engagement of the cell membrane with tethering molecules on neighboring cells or surfaces, signaling interactions at the interface of stem cells and niche or descendent cells, paracrine and endocrine signals from local or distant sources, neural input and metabolic products of tissue activity.” It is suggested here that conventional porous scaffolds do not represent the ideal format for a so-called tissue engineering “scaffold.” It is further suggested that, as discussed by the author, the term “scaffold” falls far short of the characterization and specification required for these supporting constructs. The overarching term “template” is preferable [6]. To determine what type of template offers better chances of success than conventional porous solids, it is necessary to define the specifications of these constructs; this has rarely been done within tissue engineering. Edalat et al. [4] have addressed this issue, stating that the engineering of what they describe as scaffolds requires close attention to the 3D microgeometry of the construct (including porosity, pore size, and interpore connectivity), mechanical parameters such as linearity or nonlinearity, elasticity, viscoelasticity, or anisotropy and the successful delivery of biologics including cells, nucleic acids, and cytokines. These generic requirements can be translated into detailed specifications, some of which are mandatory if optimal performance is to be achieved, whereas others are optional, depending on the precise application [1]. The mandatory specifications are: • The material should be capable of recapitulating the architecture of the niche of the target cells; • Because the cell niche is changeable over time, the material should be capable of adapting to the constantly changing microenvironment; • The material should have elastic properties, particularly stiffness, that favor mechanical signaling to the target cells to optimize differentiation, proliferation, and gene expression; • The material should have optimal surface or interfacial energy characteristics to facilitate cell adhesion and function;

Components of the local environment that participate in regulating stem cells in their tissue environment. These include the constraints of the architectural space, the physical engagement of the cell membrane with tethering molecules in neighboring cells, signaling interactions at the interface of stem cells, and the niche paracrine and endocrine signals, neural input, and metabolic products. Reproduced from Scadden DT. The stem cell niche as an entity of action. Nature 2006;44:1075e78, with permission of the Nature Publishing Group.

FIGURE 36.3

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• The material should be capable of orchestrating molecular signaling to the target cells by directing endogenous molecules or delivering exogenous molecules; • The material should be of a physical form that provides appropriate shape and size to the regenerated tissue; • The material should be capable of forming into an architecture that optimizes cell, nutrient, gas, and biomolecule transport ex vivo, in vivo, or both, and facilitates blood vessel and nerve development; • The material should be intrinsically noncytotoxic and nonimmunogenic, and minimally proinflammatory. Optional specifications that vary with the application are: • The material should be degradable if that is desired, with appropriate degradation kinetics and appropriate morphological and chemical degradation profiles; • The material should be injectable if that is desired, with the appropriate rheological characteristics and transformation mechanisms and kinetics; • Where necessary, the material should be compatible with the processing techniques that simultaneously pattern both the material and living cells; • Where multiple cell types are involved, the material properties should be fine-tunable to accommodate variable cellular requirements, with spatiotemporal control as appropriate; • When used in a significantly stressed in vivo environment, the material must have sufficient strength and toughness; • In situations in which the biomaterial encapsulates cells, optimal diffusion characteristics concerning critical molecules are required. If porous solids are generically unable to comply with all of the mandatory specifications concerned with the cellular microenvironment, consideration has to be given to the type of material that is able to do so; universally the solution to this problem has involved the group of materials known as hydrogels, which are 3D networks composed of cross-linked hydrophilic polymer chains that contain large amounts of water [7e10]. In consideration of these specifications, it is necessary to focus on the characteristics of the composition, structure, and properties of the various forms of ECM [7]. In its natural form, an ECM is composed of water and a mixture of glycosaminoglycans (GAGs) and fibrous proteins, which collectively self-assemble into nanofibrillar supramolecular networks. The precise composition and specific structure vary from tissue to tissue, which implies that the hydrogel for any one construct will depend on the intended tissue application. The main fibrous proteins are the collagens, elastin, fibronectin, and laminin, whereas principal GAGs are hyaluronates, dermatan sulfates, heparin and heparan sulfates, chondroitin sulfates, and keratan sulfates. In addition, many of the GAGs link to core proteins to form proteoglycans. It should be of unsurprising that many of the currently favored hydrogels are based on naturally occurring molecules found in these ECMs. The water content also varies, both with the type of tissue and regionally within any tissue type. Articular cartilage water content varies from 65% to 80% on a regional basis. Within the intervertebral disk, the ECM water content is around 75% in the annulus fibrosis and 80% in the nucleus pulposus; importantly, this decreases with disk degeneration, typically by 10% in advanced cases [11]. From a purely compositional perspective, the ECM appears to be analogous to a hydrogel, albeit a complex one; this was pointed out clearly by a number of reports [12,13]. Naturally, the presence of cells appropriate for the tissue in question and the presence of mineral phases in tissues such as bone and dentin influence the biological and physical properties. The topology, or architecture, reflects the dynamic, reciprocal relationship between the cells and the molecular microenvironment, and the topology and biochemical composition of the ECM is markedly heterogeneous [14,15]. As a form of hydrogel, various types of ECM would be expected to have gel-like mechanical properties, especially those of viscoelasticity, and indeed this is the case [16]. These characteristics vary, especially the complex shear modulus, both regionally within a tissue and with physiological variables such as aging and disease [17,18]. The ECM not only provides the structural and physical support of cells and their biomechanical stimuli in tissue, it initiates critical biochemical cues that are required for tissue function, including morphogenesis, differentiation, and homeostasis. This has to be seriously considered when designing hydrogel templates for tissue engineering. One critical aspect is adhesion between the cells and their ECM, which is mediated by receptors such as integrins. The ECM also is responsible for morphological organization and physiological function by binding growth factors and interacting with cell surface receptors to generate signal transduction and regulate gene expression [14]. The complexity of tissue function requires the ECM to be a dynamic structure that is continually being remodeled and subjected to a variety of posttranslational modifications. Hydrogels used as tissue engineering templates have to address this multitude of characteristics and properties of the ECM.

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STRUCTUREePROPERTY RELATIONSHIPS IN HYDROGELS As noted earlier, hydrogels are 3D networks composed of cross-linked hydrophilic polymer chains, which can contain large amounts of water, up to thousands of times their dry weight. The structureeproperty relationships that characterize these networks depend on a number of factors relating to their source, the polymer composition, the type and nature of the cross-links, their electrical charge, and waterswelling behavior. Hydrogels may be natural or synthetic, or possibly a blend or composite of the two groups. They may be homopolymers (based on a single species of monomer), copolymers (involving two or more different monomer species, at least one of which is hydrophilic), or interpenetrating networks (containing two independent cross-linked components). The hydrogels may be crystalline, semicrystalline, or amorphous. Different properties are seen, depending on the presence or absence of electrical charge on the chains, which may be neutral, anionic or cationic, amphoteric (with both basic and acidic groups), or zwitterionic (with both anionic and cationic groups) in each repeating unit. One of most important characteristics is the ability to swell and deswell reversibly in water on the basis of environmental stimuli; chemical stimuli that induce this behavior include the pH, solvent composition, and ionic strength, whereas physical stimuli primarily include the temperature and electrical or magnetic fields. As reviewed by Slaughter et al. [19], cross-linking is characterized by junctions, or tie points, formed by covalent or ionic bonds between polymer chains, or by physical entanglements or weak interactions, such as by hydrogen bonding (Fig. 36.4). The resulting network structure can be quantified through a number of parameters, including the polymer volume fraction in the swollen state, the average molecular weight between cross-links, and the distance between them, which is obviously a measure of the mesh size and porosity. This latter parameter is critical for determining the properties of solute transport in the hydrogel, which is obviously important for compliance with the requirement that a tissue engineering template facilitates the transport of nutrients and other molecules. Diffusion is considered to be the dominant mode of transport in tissue engineering hydrogels. This is influenced by the electrical charge, the mesh size, and environmental factors such as the pH and temperature. In complex biological media in which both the polymer and solute are likely to be ionized, interactions between them may have significant effects, generally decreasing solute transport. The precise mechanical characteristics of a hydrogel range from rapid elastic recovery from applied stress or strain to time-dependent viscous recovery; a major factor in tissue engineering hydrogels is the relationship between the glass transition temperature, Tg, and body temperature [19]. At temperatures below Tg there is a tendency for a greater viscous component, with contributions from creep and stress relaxation. However, the very high water content associated with these hydrogels tends to suppress Tg, favoring a more rubbery elastic regime. We may use the arguments of Buwalda et al. regarding the historic development of hydrogels for biomedical applications [20] to capture the increasing sophistication of hydrogel design as attempts have been made to Covalent bonds

Physical hysicaal entanglement ntaangl glement

P

Physical junction

FIGURE 36.4

Hydrogel network with physical junctions, covalent bonds, and physical entanglements.

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incorporate the ECM specifications described in the last section. As those authors noted, the first significant event in hydrogel design related to such applications was the introduction of water-swollen, cross-linked, macromolecular networks to construct poly(2-hydroxyethyl methacrylate) (pHEMA) soft contact lenses by Wichterle and Lim in 1960 [21]. The specifications for such a lens were simple and involved sufficient water content for optical transparency, the lack of extractable toxic components, resistance to degradation, good permeability to water-soluble substances (especially oxygen), and suitable mechanical properties; accordingly, the hydrogel structure that was developed was also simple. These first lenses were prepared by the free radical polymerization of 2-hydroxyethyl methacrylate (HEMA) in aqueous solution using ethylene glycol dimethacrylate as a cross-linker. Buwalda et al. [20] defined this type of simple hydrogel as being a “first-generation” hydrogel. In general, the use of “generation” descriptors is not helpful in biomaterials science [1] because such classifications are subjective, but the use of increasing complexity with specifically designed hydrogels as a marker of suitability for tissue engineering templates serves a good purpose. The baseline of simple synthetic hydrogels encompasses a small group of hydrogels prepared by the polymerization of water-soluble monomers and a group based on cross-linking watersoluble polymers. pHEMA belongs to the first group. These polymers were successful in a number of applications but were soon shown to have several limitations, especially insufficient oxygen (and other solute) transport and mechanical fragility. N-Vinylpyrrolidone (nVP) was used as an alternative water-soluble monomer; polyvinylpyrrolidone had higher hydrophilicity and overall improved biocompatibility. Nevertheless, these were not ideal, and various copolymers of HEMA, such as nVP, substituted acrylamides, vinyl acetate, and substituted methacrylates, were subsequently developed. These produced hydrogels with variations in basic properties but with no significant change in the underlying philosophy. The second group of simple hydrogels, based on the cross-linking of water-soluble synthetic polymers, largely involved poly(vinyl alcohol) (PVA) and poly(ethylene glycol) (PEG). PVA is a linear polymer produced by the free radical polymerization of vinyl acetate, followed by the hydrolysis of the acetate groups to alcohol groups [22]. Cross-linking has been achieved through the use of bifunctional or multifunctional agents that are able to react with the hydroxyl groups of the PVA molecules, including aldehydes such as glutaraldehyde, or through the use of electron beam or g-irradiation. Obtaining the optimal balance of biocompatibility (potentially influenced by residual chemical cross-linking agents) and mechanical properties has not been straightforward. PEG and its highere molecular weight counterpart, poly(ethylene oxide), are widely used hydrogels in pharmaceutical preparations; the highly water-soluble polymers are readily cross-linked by g or electron beam radiation [23]. PEG hydrogels are versatile because molecular chains can have a variety of forms, such as linear or star-shaped, and because they can be biologically modified readily. This provides a good example of how the basic synthetic hydrogels have led to more sophisticated structures that can approximate the ECM paradigm discussed previously.

INCREASING SOPHISTICATION OF SYNTHETIC HYDROGELS FOR TISSUE ENGINEERING Bioactive Forms of Poly(ethylene Glycol) as Exemplars of Increasing Sophistication The PEG hydrogels discussed in the previous section provide a good example of how modifications, especially biological ones, can improve the performance of synthetic hydrogels [24,25]. The major limitation of these hydrogels with respect to regenerative medicine applications is the lack of cell-specific adhesion. One approach to overcoming this involves the use of cell-adhesive peptides in modifying the PEG structure. These have generally been derived from the major ECM proteins fibronectin, laminin, and collagen. The most widely used peptide has been arginineglycine-aspartate (RGD) [26], which can be derived from the cell-binding domain of all three proteins; both linear and cyclic RGD have been employed, with preference for the latter in view of the greater affinity for the integrin anb3. The peptide may be linked to the PEG network by a number of different groups, such as monoacrylates. A second modification to PEG involves incorporating growth factors. These are polypeptides that transmit signals to modulate cellular activity, which is important in tissue engineering processes. However, in their free form they have short half-lives because they are susceptible to proteolytic degradation. Growth factors may be incorporated into PEG directly during hydrogel formation, but this usually results in a burst release; this is likely to be unsatisfactory because the dosage response required to facilitate tissue formation is precise and sensitive. Several strategies have been used to address these difficulties, particularly using functional groups to modify the growth factors before tethering them to the PEG network, such as by using cysteine to tether recombinant vascular endothelial growth factor using multiarm PEG vinyl sulfone [27]. Among other growth factors that have been attached to PEG are basic

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fibroblast growth factor, transforming growth factor-b, epidermal growth factor, and various bone morphogenetic proteins. Various methods have also been used to mimic the growth factor binding mechanism of GAGs in the natural ECM by the chemical modification of heparin, chondroitin sulfate, and hyaluronic acid, using reactive groups such as acrylate or thiol, and subsequent reaction with functionalized PEG derivatives [28]. With these PEG-based examples in mind, generic forms of hydrogel sophistication are now discussed.

Spatial Heterogeneity As alluded to earlier, biological processes are usually regulated by spatially dependent signals, in which gradients of molecules are able to regulate cell migration, axon extension, angiogenesis, differentiation, and other processes. As discussed by Khademhosseini et al., control of the spatial location of molecules on a surface and/or throughout a material template could be beneficial for tissue engineering [29]. One approach to generating these gradients involves releasing molecules from a source over time to form a concentration gradient as the molecule diffuses away from the source; in general, however, these gradients are unstable and it is difficult to control their shape. An alternative concept that involves conjugating biomolecules to materials has been used to increase the stability of signaling molecules in a spatially controlled manner. For example, multicomponent, spatially patterned, photocross-linkable hydrogels may be fabricated to localize growth factors within hydrogels. In addition, microfabrication approaches provide attractive technologies because of their availability and costeffectiveness. The ability to pattern fluids within microchannels has been merged with photopolymerization chemistry to form spatially oriented hydrogels. Hydrogels may be synthesized with gradients of signaling or adhesive molecules or with varying cross-linking densities across the material to direct cell behavior such as migration, adhesion, and differentiation. These concepts of controlling the microstructure and spatial compositional character in relation to the connectivity between multicomponent tissues, especially the ECM, and hydrogels was discussed by several authors [7,10,12,30e32]. Burdick and Murphy [32] referred to the motivation to introduce spatial heterogeneity into hydrogels, and discussed the role of micropatterning (Fig. 36.5). They claimed that introducing spatially specific cues in hydrogels makes multicellular constructs possible through cocultures or multilineage differentiation. One widely used micropatterning technique is photolithography, in which a hydrogel precursor material is exposed to ultraviolet light through a photomask that displays the required pattern. This provides reliable shape definition and is able to pattern multiple cells with materials to facilitate the selective adhesion of individual cell types to specific regions; photocross-linkable hydrogels are placed underneath the mask that controls the exposure to generate the structures in the shape of the mask. Soft lithography allows microfabrication at the micron scale, especially using silicon-based elastomers (e.g., polydimethylsiloxane) in microfluidic systems. This is of great potential interest in the development of constructs that have microchannels that resemble the vascular systems of tissues; so far, nutrient and metabolite diffusion has been observed only in relatively small hydrogel-based constructs owing to transport limitations, so microfluidic and nanofluidic techniques that allow for the creation of channels to overcome these limitations are immensely important [33e35]. Most techniques that have been investigated with respect to spatially heterogeneous hydrogels involve nanofibrous architecture [31]; these techniques include electrospinning, phase separation, and self-assembly. Electrospinning is an old technique; it dates back to the 1970s with respect to medical technology, but it is now considerably more sophisticated in relation to the materials used and the structures produced. For example, Sun et al. investigated the use of electrospun photocross-linkable hydrogel fibrous constructs for skin flap regeneration that possess the dual properties of a fibrous nanostructure and hydrogel softness, designed to allow cell migration into the scaffolds to develop 3D microvascular structures [36]. They hypothesized that such a hydrogel would be conducive to endothelial cell adhesion and growth, tubulogenesis, skin flap adhesion of the wound bed, and the formation of microvasculature. This should increase the number of capabilities to aid blood supply and enhance the survival rate of random skin flap after implantation. Gelatin methacryloyl (GelMA) hydrogel, fabricated by incorporating methacrylate groups to the amine-containing side groups of gelatin, was the photocross-linkable hydrogel used. The methacryloyl groups maintained the properties of gelatin and also allowed solidification from liquid to solid permanently via the chemical reaction of the methacryloyl groups. Also, by varying the polymer cross-linking density to control the hydrogel network structure, the mechanical, degradation, and biological properties could easily be fine-tuned. The study demonstrated the suitability of scaffolds for accelerated vascularization and that electrospun GelMA nanofiber scaffolds could support cell adhesion, proliferation, migration in vitro, and the formation of 3D vascular networks in vivo. The photocross-linkable gelatin exhibited controllable mechanical and degradation

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(A)

(B)

(C)

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Cell patterning and critical areas of cell adhesion on micropatterned surfaces with adhesion contrast. (A) Optical micrographs of micropatterns with arginine-glycine-aspartate (RGD) peptideegrafted gold microislands of varying diameters on poly(ethylene glycol) hydrogel. (B) Fluorescence micrograph of MC3T3-E1 cells, a preosteoblast cell line, on a micropatterned surface with cellular nuclei labeled by 40 ,6diamidino-2-phenylindole. (C) Fluorescence micrograph of adherent cells with cellular nuclei labeled in blue and F-actin in red. Dashed circles indicate the contours of the underlying RGD-grafted microislands. (D) Schematic presentation of three characteristic areas for cells adhering on adhesive microislands on a cell-resistant background: Ac1, also named A*, the critical area from apoptosis to survival; Ac2, the critical area from single-cell adhesion to multicell adhesion; and AD, the characteristic area for one more cell to adhere. Courtesy of Professor Jiandong Ding and Dr. Ce Yan of Fudan University, China.

FIGURE 36.5

properties, resulting in nanofibrous templates to provide the rapid regeneration and formation of cutaneous tissues with minimal inflammation. Spatial heterogeneity is also important in other tissues and for different reasons, including applications in musculoskeletal engineering [37]. In electrospun constructs, fiber alignment promotes the formation of long lamellipodia extensions parallel to the direction of the fibers, resulting in directional cell orientation and migration through mechanisms of contact guidance similar to that seen in native tissue environmental signaling. Cellular orientation is particularly important in the biomechanics of musculoskeletal tissue, because it determines the pattern of ECM deposition, which is an essential factor in the functionality of bone, tendon, ligament, and cartilage. In bone tissue engineering, for example, fiber alignment and consequent cellular alignment have been shown to regulate cell adhesion and migration, promote osteogenic phenotype, differentiate stem cells toward osteogenic lineage, and enhance mineralization and osteogenesis. Fiber alignment associated with electrospun fibers also closely resembles that seen with collagen fibrils in tendon tissue. Such anisotropy promotes elongated physiological cell morphology, the phenotype maintenance of tendon-derived cells, and the transdifferentiation of other cell types toward tenogenic lineage. The porous 3D nature of electrospun materials also provides a good environment for chondrogenic phenotype maintenance, the chondrogenic differentiation of stem cells, and new tissue formation in vivo in both cartilage and osteochondral defects. Fiber size as well as alignment is an important variable in maintaining cell phenotype and function in cartilage and tendon engineering. One interesting strategy to enhancing the complexity of constructs is based on the use of emulsions or multiaxial nozzles to produce multicomponent coreesheath fibers with multiple, often immiscible components. Such systems have been used extensively for bone, cartilage, and tendon repair.

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Electrospinning has also been combined successfully with other fabrication technologies with optimal outputs for various clinical targets. For example, multiphasic scaffolds can be fabricated using electrospinning and additive manufacturing techniques, yielding constructs with large size pores essential for cell and mass transportation, together with fibrous components that provide suitable substrates for cell attachment. Molecular self-assembly, a ubiquitous biological process, is defined as an autonomous organization of components into patterns or structure without human intervention, and is identified across length scales ranging from DNA base pairing to microtubule fabrication, and up to macroscale tissue morphogenesis [31,37]; unsurprisingly, self-assembly features powerfully in the concepts of structural heterogeneity for the creation of hydrogel-based ECM mimics. The natural extracellular environment is synthesized and organized through self-assembly in hierarchical motifs in which, in a dynamic equilibrium, the structural and chemical milieu needed to promote a range of physiological functions are controlled, including cell morphology, proliferation, attachment, migration, and tissue morphogenesis. Self-assembled template fabrication aims to replicate the sophistication of nature to produce hierarchical 3D tissue equivalents, in which hydrogels have a dominant role. Molecules that are structurally conformable and have chemical complementarity can spontaneously self-assemble into supramolecular architectures under appropriate conditions of temperature, pH, and anionic strength. These will be held together by reversible, noncovalent bonds, which, although they are inherently weak individually, yield strong and stable complexes through their collective interactions. Self-assembled natural, synthetic, or peptide-derived hydrogels have the ability to capture and deliver living cells while controlling their fate. They may also immobilize and control the release of potent biological and bioactive molecules, preserving their molecular conformation and bioactivity. By way of example, collagen nanotextured microfibers, produced by extrusion into a series of neutral phosphate buffers and cross-linking or functionalization solutions at 37 C, represent a significant advance in the recapitulation of the hierarchical architectural organization of musculoskeletal tissues. The self-assembly of peptides can result in the formation of nanofibers with very high aspect ratios, which may be able to mimic the physical microenvironment of cells, such as wrapping around cells and acting as ties between adjacent structures. One prominent area involves peptide amphiphiles. The development of tissue-engineered nerve conduits used in the setting of complex nerve injury has seen interesting developments with peptide amphiphile nanofibrous constructs. These should mimic the aligned architecture of native nerves to support directional axonal regeneration, but also provide the bioactivity found in the native ECM that facilitates Schwann cell attachment, proliferation, migration, and function. Stupp and colleagues used a series of aligned nanofiber gels formed by self-assembling peptide amphiphiles in peripheral nerve regeneration [38]. Because of their molecular design, these nanofibers can mimic the internal fascicular architecture of peripheral nerves, allowing for the incorporation of Schwann cells vital for peripheral nerve and inducing cellular and neurite alignment and guiding cell migration. They can also be engineered to possess bioactivity that is relevant to nerve regeneration; a peptide presenting the amino acid sequence IKVAV (derived from laminin) was shown to induce neural stem cell differentiation, stimulate neurite outgrowth, and result in functional improvement in acute spinal cord injury.

Matrix Mechanics The biophysical properties of the ECM are important determinants of many biological processes. As discussed by Gattazzo et al. [39], every cell in its anatomical location has to balance external forces dictated by the mechanical properties of its environment resulting from compression exerted by neighboring cells and the stiffness of the surrounding ECM. The cells have to regulate their own cytoskeleton, generating internal forces that are transmitted to the environment by adhesion sites. These focal adhesion complexes include integrins and signaling proteins, and these physically link the cytoskeleton to the ECM. Mechanical forces are exerted on and by each cell, and this interplay generates a tension within the cytoskeleton that allows maintenance of cell shape and the dynamic response to external forces. This response to mechanical stimuli is referred to as mechanotransduction. The current author explained elsewhere the pivotal role that mechanotransduction has in many biocompatibility phenomena [40]. Included here are interactions between hydrogels and other template materials with stem cells, as discussed a few years ago by Lutolf et al. [41]. Mechanotransduction profoundly affects the behavior of stem cells, both under natural circumstances and within tissue engineering systems, such as in in vitro bioreactors. The forcedependent cell signaling processes in stem cell differentiation were reviewed by Yim and Sheetz [42], with a special emphasis on focal adhesions, mechanosensitive ion channels, cytoskeletal contractility, Rho guanosine triphosphatase (GTPase) signaling, calcium signaling, and nuclear regulation. Many individual components of the various

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pathways in these systems are clearly force-dependent, including the binding of vinculin to talin during initial stages of focal adhesion assembly [43] and the activation of RhoA and Cdc42 in neurogenesis in neural stem cells [44]. Dealing first with in vitro bioreactor-based tissue engineering, two separate types of mechanical cue influence stem cell behavior, only one of which is concerned with a biomaterial property. These aspects were discussed by Steward and Kelly with respect to the mechanical regulation of mesenchymal stem cell (MSC) differentiation [45]. The first type refers to the shear stress system imposed by the mechanics of the bioreactor, which include spinner flasks, rotating wall bioreactors, and perfusion bioreactors, as described by Yeatts et al. [46]. Each of these provides different stress systems and dynamic variations in shear stresses. A primary shear stressedriven signaling pathway in the differentiation of MSCs in both osteogenesis and chondrogenesis is that of mitogen-activated protein kinases. Mechanical stresses are involved in pathway activation and in the upregulation of the proteins on which the pathways depend. Although the physical characteristics of any biomaterial template, including porosity, have some influence on fluid flow, they are not the primary determinant of the shear stresses that affect the cells. The second type of mechanical cue is that of structural stresses, perhaps best seen in cell-seeded scaffolds in static culture, in which hydrostatic pressure results in stress transfer between biomaterial surfaces and cell membranes. The precise nature of the stresses at these interfaces, including magnitude and type (especially tensile or compressive), has a strong influence on the gene expression of the cells and the differentiation pathway down which they are directed. The mechanisms here are likely to reflect the normal processes of stem cellematrix interactions within the microenvironment of cell niche; the material property most likely to influence the cell fate is substrate stiffness, or elasticity. In particular, MSCs clearly respond to 3D hydrogel stiffness. They are modulated by integrin binding through the reorganization of ligand presentation at the nanoscale; matrices of 11e30 kPa stiffness induce MSC osteogenic differentiation whereas those of 2.5e5 kPa show adipogenesis. At this stage, there is a lack of consistency in the details of the causal relationship between stiffness and cell fate when considering all types of cells and all practical conditions, largely because of the interactivity among different mechanisms, but it is clear that mechanotransduction is a primary controlling factor in the phenomena of biomaterialebioreactor-induced stem cell differentiation. The situation is similar with in vivo tissue engineering, in which much evidence points to a role of mechanical stress in tissue regeneration associated with injectable scaffolds. Myocardial tissue engineering provides a good example. The disparity between the stiffness of myocardium and injectable hydrogels and the importance of associated stress fields were addressed by Reis et al. [47]. When cardiovascular progenitor cells are contained in cardiac ECMefibrin hybrid scaffolds, their differentiation is affected by the stiffness as well as the composition of the hydrogel. For example, von Willebrand factor gene expression is upregulated with increasing gel stiffness, although such effects vary between adult and neonatal cell sources and with other relevant variables. The potential role of biomaterials as stem cell regulators was extensively analyzed by Murphy et al. [48]. It is relevant to repeat here a major part of their conclusions: “Although there are many mechanisms at play at the cell/material interface, the fundamental interaction that all cells must have is a link between the cytoskeleton and the material. The consequences of this interaction include a cascade of events in the cell, all of which are initiated by the cytoskeleton or by structures that link it to the material . the cytoskeletal protein actin and its molecular motor myosin II bind and slide past one another to contract the cell. This mechanism is highly organized in muscle, yet it is present in all adherent cell types and in stem cells it enables them to ‘feel’ the stiffness and topography of the environment, as well as to control their size, shape and polarity. Although such inherent properties of the material may seem disparate, they are united by a common contractility-based mechanism that directs stem cells towards specific lineages based on the degree of activation.” Two further points are worth mentioning. First, some tissues that are prime targets for tissue engineering solutions may be remarkably heterogeneous and/or anisotropic with respect to elasticity, so that replication of the matrix mechanics may not be easy. Within the myocardium, for example, even healthy ventricular ECM can show an effective Young’s modulus ranging from 30 to 75 kPa in different regions; infarcted myocardium can be much stiffer than this [18]. Second, it may be difficult to decouple matrix mechanics from inherent porosity and permeability [49], so that altering hydrophilicity and cross-link density, for example, can have varying effects on stiffness; this has to be taken into account in hydrogel design.

Hydrogel Degradation It is usually considered necessary for tissue engineering templates to be degradable so that they are replaced as new tissue develops. A few general points need to be made about in vivo polymer degradation in relation to the

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mechanisms of degradation and the biological consequences of degradation processes [1]. First, decades of development have shown that it is possible to design biomaterials that are essentially resistant to degradation in tissue environments (such as polytetrafluoroethylene) or are rapidly degraded (such as poly[glycolic acid]) and many that have characteristics in between. The principal mechanism of degradation is hydrolysis, so that susceptibility to degradation is largely controlled by hydrophilicity and the presence of hydrolyzable bonds. By definition, hydrogels are hydrophilic, so their degradation behavior will be determined by the hydrolyzable bonds. If these are numerous and homogeneously distributed, the hydrogel should be rapidly and uniformly degraded. It is also possible under some circumstances for polymers to be degraded by other mechanisms such as oxidation. Polyolefins such as polypropylene may experience oxidative degradation under some in vivo conditions, which is why implantable devices made of such polymers usually contain antioxidants. More important, many biological processes that occur in the vicinity of biomaterials, including inflammation, involve reactive oxygen species, which may become associated with degradation. Other reactions involve tissue enzymes, which may also be able to influence polymer degradation. Because these processes are not necessarily uniform, either spatially or temporally, it follows that degradation effects may be heterogeneous. As noted subsequently, this is an important point in the design of complex heterogeneous or anisotropic hydrogel templates. Second, degradation processes result in the generation of by-products. It is essential that these by-products be compatible with the tissue engineering environment. Ideally, the hydrogel should be bioresorbable: that is, the by-products should be metabolized and harmlessly assimilated into the tissue, such as water and carbon dioxide. It is very important that the by-products not be proinflammatory or cytotoxic. Third, degradation processes may be influenced by mechanical stress, which should be taken into account in template design. As pointed out by Burdick on several occasions [32,49], uniform hydrogel constructs with homogeneous degradation profiles considerably oversimplify the complexity of the temporal dynamics that are present in tissue development. As implied in the earlier parts of this chapter, the natural ECM undergoes dynamic remodeling through a combination of matrix assembly and degradation, and this characteristic should be replicated in a hydrogel template. These processes occur through the effects of molecules such as proteolytic enzymes, such as matrix metalloproteinases (MMPs), which are produced by cells during migration and signaling. If hydrogels are simple uniform structures in which degradation is solely controlled by cross-link density, there will be no local control over degradation behavior. In general, hydrogel degradation rates can be fine-tuned by manipulating network connectivity and mesh size. Increased the cross-linking density usually leads to a smaller mesh size, an increased elastic modulus, and slower degradation, because there will be an increased number of cleavable bonds that have to be broken for network mass loss and erosion. Decreased the mesh size also limits accessibility of the degradable moiety to larger molecules, including enzymes, because of a reduced diffusion rate. In addition, encapsulated cells, cell-secreted enzymes, and growth media can influence the degradation rates for chemically or physically cross-linked hydrogels [50]. Hydrogels can degrade by bulk degradation, surface erosion, or a combination of the two (Fig. 36.6). At high cross-link density, restricted diffusion of water and enzymes preferentially leads to surface erosion. In hydrogels with high water content and high diffusivity, bulk degradation occurs when cleavable groups are present throughout the bulk and may degrade simultaneously. Physically cross-linked hydrogels can degrade by processes that reverse the gelation mechanism or disturb the noncovalent interactions of the cross-links. Chemically cross-linked hydrogels degrade via several mechanisms, including cleavage of the backbone chain, cross-linker, or pendant groups. Also, hydrogels prepared using polymers with degradable functional groups within the backbone chain may be degraded into smaller segments, depending on the location of the degradable groups. Many hydrogels include degradable cross-linkers, such as peptides, proteins, or polymers with chemically labile moieties. These networks degrade into highemolecular weight polymer backbone chains, with degradation products derived from the cross-linker. Polymer chains also can be end-capped with degradable functional groups followed by the addition of reactive functionalities, thus creating cross-linkable degradable macromers. Chemically cross-linked hydrogels may be degraded through hydrolysis, enzymatic cleavage, reversible click reactions, or photolytic degradation. Obviously, to engineer hydrogel degradability, it is necessary to understand the types of cleavable groups and modes of degradation, their by-products, and factors affecting degradation rates.

Polymerization Mechanisms These sections have referred to the general mechanisms of hydrogel formation and the effects of, for example, different cross-linking processes on the resulting complexity and properties of the hydrogel. The development of

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FIGURE 36.6 Degradation mechanisms for hydrogels, involving erosion, bulk degradation, or a combination of these. (A) Polymers subject to degradation at the cross-links. (B) Degradation of the backbone. (C) Release of pendant side groups.

new techniques of cross-link chemistry has led to new structures that are able to mimic the ECM better. Lin discussed such advances with respect to PEG hydrogels [51]; these examples are instructive. It will be recalled that although PEG-based hydrogels have a high water content, tissue-mimicking elasticity, solute permeability, and cytocompatibility, the lack of biological recognition sites is a significant disadvantage. Therefore, PEG hydrogels have been developed that are functionalized through copolymerization with peptides during network cross-linking. For example, acrylated or methacrylate peptides can be copolymerized within PEG-diacrylate or PEG-dimethacrylate hydrogels through chain-growth homopolymerization. PEG-based hydrogels can also be prepared by stepgrowth photopolymerization, giving a more homogeneous network structure and better mechanical properties. Photopolymerization has the advantage of excellent spatiotemporal control of gelation characteristics. Several bioactive motifs can be incorporated into the hydrogel using this technique. Click chemistry has been embraced in hydrogel preparation [52] because of the high reactivity and selectivity that can be achieved, as well as the mild conditions that are involved. Click chemistry produces highly efficient, quantitative, orthogonal reactions between mutually reactive functional groups. With reference to PEG, MMP-sensitive peptide sequences can be incorporated into the hydrogels using nucleophilic Michael-type addition reactions between these cross-linkers with terminal cysteines, and multiarm PEG-vinyl sulfone [53]. Cell-adhesive ligands may be readily conjugated using similar Michael-type addition reactions. One major advantage of these types of nucleophilic additions reactions is that they avoid the generation of radicals, which often compromise the biocompatibility of the systems. Biomimetic hydrogels with good biocompatibility may also be prepared by macromolecular or supramolecular self-assembly, especially where gelation is produced by physical processes, which again avoids the use of radicals. Amphiphilic cyclodextrins are widely used in this context, with physical interactions between their inner hydrophobic cavity and hydrophobic molecules; the hydrophilic outer surfaces facilitate dissolution in physiological environments.

Injectable Systems In the context of different polymerization mechanisms, mention must be made of the possibility of forming in situ or in vivo gelling systems [54]. The attraction is obvious of being able to inject a viscous sol into the tissue at the site where regeneration is required, under ambient physiological conditions and in a short time (Fig. 36.7).

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Normal Heart

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Injectable hydrogel Myocardial Infarction administered at the in heart infarct site

Infarct site

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FIGURE 36.7 Example of potential benefits of injectable hydrogels in regenerative medicine. Epicardial delivery of hydrogel-based solution carrying cells and biomolecular signals, which upon administration forms three-dimensional hydrogel over the infarct site owing to cross-linked networks. Reproduced from Radhakrishnan et al. Biotech Adv 2014;32(2):449e61, with permission of Elsevier.

In situeforming hydrogels may be prepared using noncytotoxic cross-linking agents, physical interactions, or selfassembly supramolecular chemistry. These resulting gels should be able to take on the shape and architecture of the site where they are injected and preferably adhere to the local tissue. However, success has been elusive owing to the narrow range of biologically acceptable stimuli for triggering the physical interactions, the generally low mechanical properties, the difficulty of incorporating bioactive agents, and often poor stability. The range of options for gelling reactions follows the systems discussed in the previous section on general polymerization mechanisms. Much attention has been given to Michael-type addition reactions, especially in PEG, collagen, hyaluronic acid, and heparin systems. Click chemistry is also employed.

NATURAL BIOPOLYMERS AS EXTRACELLULAR MATRIXeANALOG HYDROGELS It is not the intention in the next two sections to provide detailed catalogs of tissue engineering hydrogels, but rather to give a summary of the salient features of the various species that are attracting attention. In both categories, i.e., natural and synthetic biomaterials, categories can exist in solid polymer, elastomer, and nanostructured forms as well as hydrogels. Biopolymer or naturally derived hydrogels include those based on hyaluronic acid, alginates, collagen, fibrin, and peptides. They tend to be considered superior to synthetic gels with respect to biocompatibility because they may offer better molecular and morphological cues to cells. However, they have potential disadvantages associated with the sourcing of raw materials from natural origins where purity and consistency may be less than ideal for a quality product used in health care.

Polysaccharides Generically, a polysaccharide is a complex carbohydrate composed of a chain of monosaccharide units linked by glycosidic groups. Sugars are monosaccharides or disaccharides; polysaccharides have much larger molecules. They clearly have some similarities, but structural differences between different types mean that it is difficult to find common ground that distinguishes the group as a whole from other materials, and they are best dealt with individually. Some of those discussed here are GAGs (mucopolysaccharides), which are long, linear, unbranched polysaccharides of repeating disaccharide units. Some, such as hyaluronan (HA), are nonsulfated, but several, including heparan sulfate, dermatan sulfate, and keratan sulfate, are sulfated. Hyaluronic Acid HA, otherwise known as hyaluronic acid, is a linear glycosaminoglycan that has a molecular mass of 106e107 Da, with very long molecules that consist of linear chains of repeating units of disaccharides of glucuronic acid and N-acetylglucosamine. Two main features of HA have contributed to its attractiveness as a biomaterial [55]. First, it is contained in the properties of many tissues in the human body and contributes to them, which suggests the possibility of

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recapitulating some of these within therapeutic products. Among the specific biological properties of HA are its role in embryonic development and wound healing. It is present in high concentrations in synovial fluid and in the ECM of cartilage, so it has a significant role in the functioning of articulating joints. HA interacts with some cell surface receptors and is involved in angiogenesis, cell migration and motility, and tissue organization. It has a role in inflammation and the stimulation of cytokine activity. It can also be functionalized and chemically modified to present a wide range of physical characteristics, with wide-ranging solubility and mechanical properties. The carboxyl group can be functionalized through amidation, esterification, or oxidation. The hydroxyl group may undergo esterification or etherification and the acetyl group may be reacted by deacetylation or amidation. HA can be derived from a number of sources; one of the most common is rooster comb. It is highly soluble, especially at a low pH, and has a high rate of turnover in human tissue. For the purposes of creating a practical biomaterial, HA is normally cross-linked, for which a number of methods are available, and it may be modified with other substances. Covalent cross-linking provides the opportunity to achieve hydrogels, sponges, and other solid forms while maintaining biological functionality. Cross-linking may take place using water-soluble agents such as a carbodiimide or by the use of photocross-linking using glycidylmethacrylate or methacrylic anhydride. As a gel, it has high viscoelasticity, a major factor in its use in ophthalmic surgery and in therapies for osteoarthritis. It can be degraded by reactive oxygen intermediates; unmodified HA is rapidly degraded and cleared from the site of administration. To reduce the rate of degradation, cross-links may be introduced. It may also be modified with peptides to enhance cell attachment, spreading, and proliferation. For example, thiol-modified HA can be functionalized with the RGD sequence. These peptide-functionalized gels may also be used as in situ gelling injectable constructs for in vivo tissue engineering. The areas of tissue engineering most relevant to hyaluronic acid templates are brain and neural regeneration, cardiovascular tissue engineering, skin regeneration, retinal regeneration, and cartilage repair [55]. Several factors have limited the clinical applications so far, including nonspecific protein adsorption and cell adhesion, which can lead to inflammation and the accumulation of degradation products at the site of application, which inhibits stem cell differentiation. Alginate Algae are living organisms that are mostly found in water; they can be harvested and provide substances for many industrial uses [56]. These uses are mainly based on their polysaccharide content, although they are often also rich in amino acids such as proline, glycine, and lysine, which accounts for their widespread application as food additives. Seaweeds constitute an important source of harvested algae-based substances because their cell walls contain polysaccharides, which can be readily extracted. Some seaweeds produce agar (from red seaweed); others produce carrageenans. Brown algae (Phaeophyta) produce alginates and several other polysaccharides. The alginates are probably the most important seaweed-derived products and certainly the most significant from a biomaterials perspective. These are harvested in their wild state; cultivation is too expensive, which accounts for some variability in the extracted alginates, because there is some species and seasonal dependence. Alginates are extracted from the seaweed using sodium carbonate and precipitated as either sodium or calcium alginate. This is treated with diluted HCl to produce alginic acid, which us purified and reconstituted into different ionic forms, depending on the application. Alginates are linear block copolymers of 1,4-linked b-D-mannuronic acid (M) and a-L-guluronic acid (G). The properties of the material will depend on the M:G ratio. Alginates form gels and readily retain water. Different ionic forms have different solubilities; the transition from sol to gel can easily be achieved by conversion between calcium and sodium forms, a process that was extensively used in dentistry to produce elastomeric impression materials. In the solid state, alginates can form films and fibers of good structural quality. When prepared as meshes of nonwoven fibers, they make useful wound dressings, because they are able to absorb large amounts of exudate from the wound bed. Alginate products are used as food additives, moisturizing components of cosmetics, and ingestible preparations for treating inflamed mucosal surfaces. The most attention has been paid in tissue engineering and cell therapies. It is particularly notable that alginates present one of the best options for cell encapsulation, because the hydrated material will allow the diffusion of small molecules essential for metabolic activity but not the immunoglobulins that would attack the cells. The viscosity of alginates and the stiffness after gelling depend on the concentration of the polymer and its molecular weight distribution. Cross-linking between polymer chains can be arranged through multivalent cations (especially calcium) and with carboxylic acid groups in the sugars. Alginates have generally good biocompatibility and can be prepared as an injectable ionic solution. They have poorly controlled degradation and variable cell adhesion characteristics. g-Irradiation may be used to break highemolecular weight chains to allow faster degradation and clearance in vivo, whereas partial oxidation such as with sodium periodate

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increases susceptibility to hydrolysis. There is also interest in derivatization reactions on the polysaccharide backbone that enable, for example, hydroxyapatite nucleation and growth for bone tissue engineering applications and anticoagulation properties through the attachment of heparin [57]. Agarose Agarose is also a polysaccharide that originates from sea algae; it is composed of repeating units of 3,6anhydrous-L-galactose and D-galactose [58]. The agarose gels involve a rigid network that is porous and may be used for drug and gene delivery [59]. They can polymerize in situ and are attractive in the context of therapies for central nervous system injury and disease, where they support neurite extension [60]. It is also significant that agarose can form interpenetrating networks with some synthetic hydrogels, such as PEG diacrylate, into which bioactive molecules such as RGD peptides or aggrecan can be covalently immobilized or physically entrapped [61]. Chitin and Its Derivatives Chitin was first prepared from the cuticle of beetles, from which it derives its name. It is widely distributed in both animals and plants, and is in the shells of crustaceans and mollusks, the backbone of squid, the cell wall of many fungi, within marine diatoms, and so on. Chitin is a linear polysaccharide of b-(1e4)-2-acetamido-2-deoxy-Dglucopyranose, where all the residues are composed entirely of N-acetyl-glucosamine: that is, it is fully acetylated. Chitosan is a derivative of chitin, which is a linear polymer of b-(1e4)-2-amino-2-deoxy-D-glucopyranose, in which all of the residues are composed entirely of N-glucosamine, which is fully deacetylated. In nature, it is rare for the material to exist as either pure chitin or pure chitosan, and the natural biopolymer will be a copolymer of the two. Generally, when the number of acetamide groups exceeds 50%, the material is referred to as chitin, and the actual percentage is termed the degree of acetylation. Conversely, when the amino groups dominate, the material is referred to as chitosan. The dry shells of animal sources such as crabs and lobsters contain 20%e40% chitin; the remainder is proteins and calcium carbonate. Demineralization and deproteinization steps are used in the process to prepare raw chitin products. Chitosan can be prepared from the chitin by deacetylation methods involving sodium hydroxide. Moreover, chitosan is able to form a gel by itself without the need of additives. That may happen via hydrogen bonds, hydrophobic interactions, and chitosan crystallites. These hydrogels can also be formed by blending chitosan with other water-soluble nonionic polymers [62] or polyol salts. Because it is polycationic in nature under acidic conditions, chitosan can also form hydrogels through interaction with negatively charged molecules [63]. In addition, the gelation of chitosan could also be obtained through covalent bonding between polymer chains. These bonds make the hydrogel more stable because the gelation is irreversible. Chitin can exist in three polymorphic states (the a, b, and g forms); a-chitin is the most common. The biostability varies with the source of the material, the crystallinity, and the degree of acetylation. Chitosan hydrogels are pHsensitive; they are soluble in dilute aqueous conditions and precipitate into a gel at neutral pH. Generally, chitosan is susceptible to enzymatic degradation. Within animal species, a variety of chitinases are able to break the chitin down into oligosaccharides, which can then be degraded by enzymes such as b-N-acetyl-glucosaminidase to yield N-acetyl-glucosamine. Similar mechanisms exist for the degradation of chitosan to N-glucosamine. In neuronal repair, chitosan is commonly used to produce tubular structures, such as in the peripheral nervous system [64]. However, chitosan hydrogels have also been applied in neural tissue engineering. For instance, the use of chitosaneglycerophosphate salt hydrogels showed that this type of gel provides a suitable 3D environment for neurons. The addition of peptides such as poly-D-lysine may improve scaffold biocompatibility and nerve cell affinity for chitosan materials (Fig. 36.8). Cellulose Cellulose is a linear polymer of b-(1,4)-D-glucose units. It is the main component of the primary cell wall of plants. It forms as crystalline microfibrils that encapsulate the cell with a mesh-like structure, and controls, along with hemicellulose, pectin, and lignin, the mechanical properties of the plants. There is interest in using some derivatives of cellulose as biomaterials in scaffolds and drug delivery systems, for example. Microbial cellulose is a polymer that is synthesized be Acetobacter xylinum, a simple gram-negative bacterium. During the synthesis, various carbon compounds are used by the bacteria, polymerized into single linear b-1,4-glucan chains, and then secreted through pores to the cell exterior. These chains then self-assemble into subfibrils and then microfibrils and bundles, which yields a highly 3D crystalline structure with considerable mechanical strength. This nanostructure results in a large surface area that can hold a large amount of water. It can be prepared as a gelatinous membrane that is highly nanoporous. It may be treated with strong bases at elevated temperatures to remove the cells that are embedded in the cellulose

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FIGURE 36.8 Schematic diagram of chitosan-based nerve guide conduit (NGC) with intraluminal fillers supporting cell migration of Schwann cells and guiding axonal growth after implantation to bridge a peripheral nerve gap. Courtesy of Professor Xiaosong Gu, Jiangsu Key Laboratory of Neuroregeneration, Nantong University, China.

net; it is free of lignin and hemicelluloses so that the final material should be nonpyrogenic, noncytotoxic, and nonimmunogenic. It has shown considerable potential, with support from clinical evidence, for uses in wound healing and possibly as a tissue engineering scaffold. Economic large-scale fermentation systems have been difficult to optimize, which has restricted commercial applications so far.

Proteins and Peptides Notwithstanding the challenges of purity and consistency mentioned earlier, a wide variety of proteins and their derivatives have been developed as biomaterials, which may be used in several different ways. Most proteins that are used as biomaterials are based on those found in mammalian tissue. These include the structural proteins collagen and elastin, and also some that are derived from plasma proteins, including fibrin and fibrinogen. These structural proteins could be used in what is essentially their natural form: that is, as the mammalian tissues themselves, with varying degrees of processing. Alternatively, they may be extracted from such tissues and subjected to some form of purification and reconstitution. Also, because of the inherent variability in such products, it may be possible to prepare the materials by recombinant technologies. There are some generic and important differences between the reconstituted and recombinant forms of protein biomaterials. In the former case, there is always a risk of contamination with prions and viruses, but this risk is not present with recombinant proteins. The latter are fully characterized, consistent, and reproducible, whereas the former depend on the source quality and batchto-batch variability may be high. From the perspective of commercial manufacturing, recombinant proteins will usually be expensive but are amenable to proprietary processes that can be protected by patents and trademarks, which is not as readily applicable to the natural products. Collagen and Its Derivatives Collagen may be prepared in various forms of gel for tissue engineering applications, including those with the ability to form in situ. Many of these applications involve unmodified collagen; chemical cross-linkers can be used to inhibit degradation and resorption when necessary. Collagen in one form or another has been used as a biomaterial for many years. Catgut sutures, which are rarely used now, were the mainstay of surgical wound closure for a long time and were made from the collagen of bovine intestines. Purified collagen has been used in injectable form as a tissue filler for both functional (vocal chords and urethra) and cosmetic (facial) effects. Some of these products are xenogeneic; they are derived from bovine dermis, for example, whereas others are allogeneic and are obtained from cultured human dermal fibroblasts or from human cadavers. Formulations vary considerably. Some are simple suspensions (at less than 5%) in buffered saline. Others

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are prepared as micronized particles contained in syringes that are hydrated just before use. Depending on the source of the collagen, these products have the potential to be immunogenic. A small number of patients show immunological intolerance. Procedures usually involve prior testing of sensitivity. The collagen filler becomes incorporated into the patient’s tissue, but some volume loss will occur over time. The precise biochemical characteristics of the collagens used in these products are not significant because the material simply acts as a space filler. The applications of collagen in tissue engineering, wound healing, and drug delivery are more dependent on these characteristics. The structural order of collagen occurs at several different levels, and there are many different forms of the protein in mammalian tissue. Of relevance here is the need to balance the mechanical properties and biological activity of the products; this balance will change with the specific type of collagen. At this stage, collagen types I and III, alone or combined, are most commonly used in products of regenerative medicine. Conduits for peripheral nerve repair are good examples. There are products made from type I or mixed types I and III in clinical use. These have different structures; some conduits are homogeneous fibrillar structure whereas others have heterogeneous structures with concentric cylindrical structures. The materials are processed in ways to minimize antigenicity, such as by cross-linking or the enzymatic removal of antigenic nonhelical telopeptides. Degradation rates vary from a few months to a year or so. In some situations, the collagen may be copolymerized, such as with GAG molecules. Collagen type I is also used in a number of products in bone tissue engineering, although not on its own. It is usually combined with hydroxyapatite or tricalcium phosphate, either as a reinforced composite or as phosphate-coated collagen fibrils. It was noted earlier that recombinant collagen may have some advantages over animal- or human-derived materials. A number of methods are available for recombinant collagen production. This occurs in bioreactorbased eukaryotic systems, mammalian cell culture, insect cell culture, and many other systems. However, the best results and the formation of collagens most appropriate for human regenerative medicine applications are obtained with mammalian cells transfected with collagen genes, in which hydroxylated full-length collagens are produced. Recombinant human collagen of types I, II, and III can be reconstituted into fibrils that can be processed into forms such as fleeces, 3D gels, and sponges, in which it is anticipated that type I will be used in bone tissue engineering, type II in cartilage, and type III in vascular tissues. In general, collagen products used in medical technology are structural materials rather than gels [65]. Antoine et al. reviewed the specific features of type I collagen that are associated with its use as a hydrogel in tissue environments [66], especially in relation to the encapsulation of viable cells in the constructs. As alluded to earlier, the properties of collagen biomaterials are variable, depending on a large number of fabrication parameters, and this is especially relevant for collagen hydrogels. The collagen source is particularly influential, with marked differences between animal sources (e.g., murine, porcine, bovine) and tissue types (tendon, skin, etc.) The method by which the collagen is extracted from the tissue is also critical. Acid solubilization is usually used for minimally crosslinked collagens whereas a combination of neutral salt solution with proteolytic digestion is needed for highly cross-linked collagens to denature them fully. The polymerization temperature affects the properties of the resulting hydrogel, with more rapid self-assembly and therefore less ordering as the temperature increases; the temperature used will reflect the need to avoid cell damage. Collagen concentration is also important, although correlations between this and resulting mechanical properties are not clear. As with the temperature, the pH is influential; cell viability will be negatively affected if the pH is outside the range 7.4e8.4, so the better properties that may be achieved at the extremes of pH are not relevant. Gelatin is obtained by the partial hydrolysis of collagen obtained from bone and other connective tissues of animals. When obtained under acid conditions, it is known as type A gelatin, and under alkaline conditions as type B. It is used in foodstuffs and in many pharmaceutical formulations. It is also used as a gel to provide an initial seal within vascular grafts, in which situation it may also be a drug carrier, such as in the delivery of antibiotics. Elastin Derivatives Elastin is the dominant protein of elastic tissue fibers; as such, it is an important component of the ECM of tissues, such as those of the lungs, skin, and blood vessels, which depend on elasticity for their function. It is derived in vivo from cross-linking of the tropoelastin monomer and is essentially insoluble. The tropoelastin/elastin molecules have hydrophobic regions from which the resulting elasticity is derived, and hydrophilic regions, which provide sites for amine-dependent cross-linking and biological signaling. The elastin possesses a number of peptide motifs that are able to influence cell behavior, including proliferation and differentiation. Such interactions take place through several cell-surface receptors, including the elastin-laminin receptor. The signaling of cells in wound healing by elastin controls the relative activity of dermal fibroblasts and contractile myofibroblasts and hence determines the mechanical properties of the subsequent repaired skin.

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The insolubility of elastin, which restricts processing, and its relatively poor strength, has limited the practical applications of this protein in biomedical applications. Thus, elastin-like or elastin-based materials have attracted more attention [67]. There are solubilized forms of elastin, including a-elastin, which is obtained under acid conditions, and k-elastin, which is obtained under alkaline conditions. Recombinant techniques have allowed the preparation of materials that are able to mimic the important and desirable regions of the elastin. Synthetic analogs of elastin may also be prepared from the aqueous processing of replicas of tropoelastin. Many of these substances are amenable to scaleup manufacture, giving a range of elastin-derived products for biomedical uses. For tissue engineering applications, many of these derivatives may be processed by electrospinning or prepared as hydrogel matrices. It is also possible to prepare blends involving elastin derivatives and other proteins, such as silk and collagen. Desai et al. discussed critical issues with elastin-based hydrogels [68] and noted that many applications requiring large and repeated deformations can benefit from rubber-like elastomeric hydrogels; there are examples in regenerative medicine that come into this category. However, hydrogels generally lack sufficient strength and elastic extensibility owing to cross-link inhomogeneity in the networks. Short, uneven intercross-link distances and poor chain extensibility limit the ability of the hydrogels to stretch. Various strategies have been used to improve hydrogel properties, such as extensibility and toughness through physical bonds. However, time-dependent recovery of physical cross-links in these hydrogels results in high hysteresis and stress softening that can lead to a loss of performance under repeated deformation. Creating a rubber-like elastomeric hydrogel requires control over the architecture of the cross-linked network, which should be homogeneous, and a careful choice of polymer. The chain length between cross-links should be constant to avoid local concentrations of stress, whereas the cross-linking scheme should ensure proper chain incorporation to avoid dangling chains that do not participate in the elastic network. Ideally, the polymer should behave like an entropic spring that recoils once unloaded, in which the driving force restoring an entropic spring is an increase in entropy as the chain goes from a stretched state with limited movements to a coiled state. Protein-based polymers may provide opportunities for engineering hydrogels with these network and polymer properties. To create elastomers, elastin-like polypeptide (ELP) sequences provide a natural choice because they have been characterized as entropic springs. ELPs are built on a repetition of the pentapeptide Val-Pro-Gly-X-Gly with a guest residue “X” that cannot be proline. They are thermoresponsive and undergo the process of inverse temperature transition in which they transition reversibly from a hydrophilic to a hydrophobic state. Thermodynamic factors including the hydrophobic interactions within the chains keep the polypeptides in an unstructured coil form after the transition and have a role in the elastomeric nature of the polypeptides. These elastomeric hydrogels are clearly interesting for tissue engineering applications. Fibrin Derivatives Fibrinogen is a soluble protein in blood that is converted to an insoluble fibrin network in the presence of thrombin during coagulation. The fibrinogen has three pairs of polypeptide chains that are joined by disulfide groups. The central domain contains fibrinopeptides A and B, which are cleaved in the presence of thrombin to form the fibrin monomer. These monomeric units form two-stranded fibrils that undergo covalent cross-linking to form the fibrin network, a process that is facilitated by CaCl2. This process can be recapitulated artificially in biomaterial products, generally known as fibrin glues or fibrin sealants, in which separate preparations containing fibrinogen and thrombin are mixed just before application at sites of surgical injury, such as during cardiopulmonary bypass, to assist in sealing tissue defects such as fistulae and facilitate tissue repair as in peripheral nerves. It is possible for fibrin glue to be derived from autogenous or allogeneic blood, although the latter is more common in commercial products. Blood components are obtained from the blood and undergo various screening and purification procedures, especially when pooled blood is involved. Fibrinogen may be isolated from the blood using centrifugation and cryoprecipitation. Typically, sodium citrate solution is added to the blood to anticoagulate it before centrifugation. The cryoprecipitation procedure may be carried out by a variety of regimes, usually using temperatures between 20 C and 80 C. CaCl2 and aprotinin are added to the solutions to optimize the gelation of the fibrinogen and thrombin solutions once mixed. The role that fibrin has in response to the injury of vascularized tissue may also be recapitulated in the formation of tissue engineering scaffolds, which have been used in many experimental systems involving several different tissues. Fibrin hydrogels prepared, as discussed earlier, from commercially purified allogeneic fibrinogen and thrombin have some attractive features but are not ideal. Mechanically, they are weak and not stiff, and they have a tendency to degrade quickly. The gels also undergo considerable shrinkage during formation. In addition, the fibrin is not particularly active biologically in this format. Several procedures may be used to modify the gel to obtain better performance. Shrinkage may be minimized by incorporating poly-L-lysine. Using a variety of synthetic polymers or calcium phosphate additions, composite or copolymer scaffolds may improve the mechanical

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properties. Cross-linking agents may improve stability. In addition, biologically active peptides or therapeutic proteins can be incorporated into the hydrogels, such as by incorporating functionalized PEG. Najjat et al. showed that it is possible to engineer fibrin gels incorporating proangiogenic growth factors that are able in a murine model to promote the engraftment of pancreatic islets in extrahepatic sites [69]. Silk Silks are proteins that are synthesized by Lepidoptera larvae such as silkworms and spiders. They are biosynthesized in epithelial cells and secreted into the lumen of specialized glands, where they are stored and subsequently spun into fibers. The properties of silk vary considerably with their source, with different amino acid sequences and mechanical properties that are fine-tuned to their specific function. The most widely used and investigated silks with respect to biomaterials applications are derived from the domesticated silkworm, Bombyx mori, and from spiders such as Nephila clavipes and Araneus diadematus. Silkworm silk is popular and has been used for medical devices such as sutures for many years; for general textiles, it has been used even longer. This silk has two major fibroin proteins, light (25-kDa) and heavy kDa chains, where the core sequence repeats include alanine-glycine with serine or tyrosine. These core fibers are encased in the glue-like sericin protein. Spider silk proteins range from 70 to 700 kDa; many such silks are characterized by polyalanine and glycine regions. Spider silk is not easy to harvest, and much emphasis has been placed on the use of genetic engineering techniques to produce synthetic versions. Cloning and expression of silks has been achieved in a number of host systems, because the sequences of complementary DNA and genomic clones encoding spider silks show highly repetitive structures that can be used to construct genetically engineered spider silkelike proteins. Silk fibers have significant hydrophobic regions and high crystallinity with extensive hydrogen bonding, which give good environmental stability and mechanical properties. They are insoluble in most solvents including water. The crystallinity results from the presence of small b -sheets within the fibers. B. mori silk can have an ultimate tensile strength of 740 MPa, a Young’s modulus of 10 GPa, and a 20% strain at break. Spider silk may have values of 950 MPa, 12 GPa, and 18% respectively. The biocompatibility of silk products varies, largely because of the varying levels of nonfibroin components such as the sericin. When used as a suture material, silk elicits a greater inflammatory response than do most synthetic polymers. Although it is classified as nonabsorbable, it degrades slowly. Silk fibers losing their tensile strength over a year or so. Proteases such as chymotrypsin can cleave proteins to peptides, especially in amorphous regions. Various silks have found utility in tissue engineering applications because they are often able to support cell growth [70]. As discussed by Melke et al., silk fibroin can be used in several different formats in tissue engineering [71]. Sponges made by freeze-drying or porogen leaching techniques are easily produced and give good porosity, although the precise design of architecture is difficult. Fibers may be prepared by electrospinning, although reproducibility is problematic. 3D printing is used, with controllable geometry and cell encapsulation, although with low resolution. The hydrogel format (Fig. 36.9) is probably the most versatile; there are opportunities for in vivo injectable cell-encapsulation applications, although the small pore size is a disadvantage. Self-assembled Peptides An increasingly important class of hydrogels for regenerative medicine is those made from self-assembled peptides [72]. These are polypeptides that assemble under specific conditions to form nanoscale structures. One prominent example is the class of self-assembled peptides made from amphiphilic molecules, derived from polypeptides linked to a polycarbon chain. The polypeptide region is typically hydrophilic whereas the hydrocarbon chain is hydrophobic. They can self-assemble into rod structures because of the arrangement of the hydrophobic regions as well as the charge shielding of the hydrophilic end groups by ionic molecules in the solution. These molecules can be decorated with functional groups to facilitate cellular adhesion and signaling. A number of other self-assembling peptides have been produced with advantages such as the ease in which gels are formed and functionalization. They tend to be mechanically weak. The significance of these engineered peptide hydrogels is that they epitomize this direction toward materials that can replicate cell niches, referred to earlier as the most important specification for tissue engineering templates. Considering the stem cell niche in particular, these contain ECM components such as laminin and hyaluronan, which present cell-adhesion ligands, and soluble factors such as cytokines and growth factors, with a constantly replenished supply of differentiation cues. Peptide materials can be designed at a molecular level to bestow combined structural and biological activity characteristics that start to address these niche characteristics. These engineered, self-assembled peptides contain relatively short chains of amino acids. Through the careful choice of amino acid monomer sequences, the peptides can fold into secondary structures such as b-sheets, which themselves

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Fibroin Larvae (5th instar)

Cocoons Film

Samia ricini

Antheraea assamensis

Antheraea mylitta

Bombyx mori

Scaffold

FIGURE 36.9 Different silkworms and silk-derived biomaterials. Sericin can also be fabricated into different biomaterials such as scaffolds, films, hydrogels, nanofibers, and particles. Images courtesy of Professor Kundu, Indian Institute of Technology Kharagpur, India.

self-assemble into hierarchical structures such as fibers and micelles. These fibrous hydrogels replicate the required cell niches far better than do other materials. It is possible that such structures may be reinforced by other nanoscale structures to give robust templates. Several forms of self-assembled peptides have reached advanced stages of development, with good biocompatibility and degradation properties and without immunogenicity [72]. B-strand peptides, which assemble into discrete b-sheets, appear to be favored. For example, templates consisting of alternating amino acids that contain 50% charged residues may be prepared. The b-sheets have distinct polar and nonpolar surfaces; included here are several self-assembling peptides such as RAD16-I and RAD16-II, in which stable macroscopic matrix structures can be fabricated through the spontaneous self-assembly of aqueous peptide solutions introduced into physiological solutions. Such peptide structures support the cell attachment of a variety of mammalian cells.

SYNTHETIC HYDROGELS FOR TISSUE ENGINEERING TEMPLATES An informed discussion about the current situation with synthetic hydrogels in tissue engineering is problematic; despite a vast amount of research and development in this area, and although such hydrogels have a reasonably good record in drug delivery systems, there has been little progress in introducing them into clinical tissue engineering products.

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One important consideration is that of intrinsic biocompatibility. It is recognized that the chemicals used in preparing hydrogels may have some toxicity, and care has to be taken if the degree of conversion is not 100%; initiators, organic solvents, stabilizers, emulsifiers, unreacted monomers, cross-linking agents, and other substances have to be considered in this light and such chemicals may need to be removed. Synthesis typically should be followed by purification processes such as solvent washing or dialysis. When any of the materials are derived from natural sources, they may carry the risk of batch-to-batch variation, which also has to be taken into account. Synthetic hydrogels have some important advantages over natural biopolymer-based hydrogels, including easier large-scale production and fine-tunable and consistent properties. However, many are made using harsh synthetic chemistry, which requires care to ensure that contaminants and unreacted reagents present during synthesis are then removed. Reaction schemes for these hydrogels often rely on multifunctional cross-linking agents. Free radical polymerization is widely used with many tissue engineering systems that require in situ formation. The main difficulty, as pointed out several times in this chapter, is that truly synthetic hydrogels usually do not have the biological functionality to provide for the specifications that have been defined for template materials. Thus, many of the newer developments have addressed this major deficiency by introducing functionalization methods and using hybrid or composite structures, often through blending or copolymerization with one of more of the biopolymers discussed earlier. Following the pattern of the Bioactive Forms of Poly(ethylene Glycol) as Exemplars of Increasing Sophisticationsection, the experiences with PEG may be used as an example. PEG is a diol with two hydroxy end groups; it is formed by linking repeating units of ethylene glycol. The versatility arises from two facts; first, it has linear and branched forms’ and second, these end groups can readily be functionalized, such as through carboxyl, amine, thiol, or azide groups. These functional end groups can be symmetric or asymmetric. The latter option is effective because it enables simultaneous but different properties to be achieved. A popular form of PEG is the four-arm-PEG, but other geometries are possible. The versatility is also facilitated by the possibility of cross-linking by different methods under different conditions. Free radical polymerization, condensation reactions, enzymatic reactions, and click chemistry can all be used, but the most common cross-linking procedure is photopolymerization. This allows the conversion of liquid to solid state under ambient physiological conditions in situ, with good spatial and temporal control and with the possibility of the simultaneous incorporation of biological species or active agents. Best results are normally achieved through the use of acrylates such as diacrylates or dimethacrylates as the macromers. PEG is normally nonbiodegradable and has little intrinsic biological reactivity; a major use of PEG is as a coating to minimize protein adsorption to surfaces. Unmodified PEG is unattractive for tissue engineering applications because of the lack of support for cell function. PEG itself, in a lowemolecular weight, unmodified form, has a number of mundane medical uses, largely as over-the-counter preparations such as laxatives and skin moisturizing agents. The real attraction of PEG as a biomaterial is associated with its combination with other molecules. There are two main scenarios here: either the PEG is used to surface modify a structure to provide or hide properties, or other molecules are used to modify PEG hydrogels to capitalize on the hydrogel characteristics but provide some specific biological activity. The first of these options is referred to as pegylation. There are many examples ranging from pharmaceutical molecules to nanoparticles. It has long been known that pegylation of protein and polypeptide drugs can alter both their pharmacokinetic and pharmacodynamic properties, such as by increasing water solubility, minimizing cytotoxicity, and reducing renal clearance. At its simplest, PEG makes the drug molecule larger; each ethylene glycol subunit is tightly associated with two or three water molecules, which makes the pegylated molecules appear up to 10 times larger than the unmodified molecule. The PEG acts a shield around the molecule, protecting it from degradation and rapid clearance. The effect may be much more subtle, depending on the chemistry involved. For example, the bond between the PEG and the drug molecule may be intentionally unstable to improve targeting, such as when the bond is cleavable by enzymes within the endosomal compartment of cells, which releases the peptide or protein molecule within the cell. Bioactive modification of PEG hydrogels for tissue engineering applications may facilitate cellular function. A variety of ECM protein-derived cell-adhesive molecules may be incorporated into the hydrogel to increase cell adhesiveness. The hydrogels may be made degradable by incorporating hydrolytically or enzymatically susceptible segments. Growth factors that have been functionalized may be covalently attached to the PEG hydrogels, especially those that involve multithiol or multiacrylate groups. PEG has been combined with a variety of nonhydrogel biodegradable polymers such as poly(glycolic acid) [73] and reinforced with fibrous collagen [74]. Carbon nanotubes (CNTs) have also been used to create hydrogel composites [75]. CNTs can impart electroconductivity to otherwise insulating materials, improve mechanical stability, guide

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FIGURE 36.10 Principles of cell sheet engineering. Conventional techniques to release cells from biomaterials surfaces using enzymes result in disruption of the cells. Techniques of cell sheet engineering, which involve a change in the hydrophobicehydrophilic balance of a thermally responsive hydrogel such as poly(N-isopropylacrylamide) (PIPAAm), produce the release of sheets of cells without disruption. Courtesy of Professor Teruo Okano, Institute of Advanced Biomedical Engineering and Science, Tokyo Womens Medical University, Japan.

Cell Sheet Manipulation by Enzyme Treatment or Temperature Change Disruption of structure and function

Cell Sheet formation On Culture Dishes

Hydrophobic Surface

Maintains structure and functions

Temperature Responsive Polymer, PIPAAm Hydrophilic Surface

neuronal cell behavior, and elicit axon regeneration. A PEG hydrogel composite was prepared in which the CNTs were entrapped in the hydrogel phase during gelation. The hydrogel cross-linking reaction was based on Michael-type addition, which is ideal for in situ cell and protein encapsulation, and in which sonication and surfactants were used to disperse the highly hydrophobic CNTs in the aqueous polymer solution. The inclusion of the CNTs impeded hydrogel crosslinking leading to longer gelation times, higher swelling and porosity, and lower storage modulus above a threshold CNT concentration. Unlike the PEG hydrogel alone, the PEGeCNT hydrogel composite was capable of supporting high neural cell viability in which the CNTs provided sites for cell attachment. As a final point, mention should be made of thermoresponsive hydrogels, which have potential uses in regenerative medicine. Poly(N-isopropylacrylamide) (pNIPAAm) is a good example [76]; it is a polymer with lower critical solution behavior in aqueous solvents and a lower critical solution temperature (LCST) around 32 C. During the phase transition, the polymer chains undergo a change in conformation from an extended coil to a globular structure. The LCST of N-isopropylacrylamide (NIPAAm)-based polymers can be increased with copolymerization with more hydrophilic monomers, resulting in a transition temperature close to physiological, which can be useful in medical applications, including drug delivery, tissue engineering, and cell sheet engineering. The homopolymer of NIPAAm is not biodegradable, and many studies have focused on imparting biodegradation by adding various monomers into the polymer structure, such as by copolymerizing NIPAAm with benzomethylene dioxepane. The development of in situeforming materials with a dual solidification mechanism, a physical gelation imparted by NIPAAm and a covalent cross-linking reaction, is interesting. The dual gelation mechanism results in hydrogels with significantly enhanced mechanical properties compared with hydrogels solidified only by physical or chemical means. Covalent cross-linking mechanisms include Michael-type addition between thiols and vinyl groups, functionalization of thermogelling macromers with (meth)acrylate groups, cross-linking with a thermal initiator, and epoxy functionalities reacting with amines. Substrates based on NIPAAm have been also used to fabricate cell sheets. When heated above the transition temperature, the hydrophobic nature of pNIPAAm favors cell attachment. The LCST of the pNIPAAm-based polymer can be fine-tuned to be below cell culture temperatures. When the temperature is lowered below the LCST, the polymer surface becomes more hydrophilic and cells can be harvested in a confluent monolayer (Fig. 36.10). This process eliminates the use of proteolytic enzymes or mechanical means for cell detachment from cell culture surfaces [77].

CONCLUSIONS Much progress has been made in the development of hydrogels as templates in tissue engineering and regenerative medicine. However, there is a long way to go. The ultimate aim is to create templates that can meet all of the specifications set out in the early parts of this chapter, especially concentrating on replicating all of the characteristics of the ECM. Although hydrogels are superior to porous solid biomaterials in this respect, the architectural and

REFERENCES

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functional requirements are exacting. Most synthetic hydrogels are too far removed from ECM characteristics for us to have serious expectations for their scientific and clinical success. Natural biopolymers have many advantages, but they still have limitations. Hybrid, composite, and functionalized structures are showing attractive properties in many situations.

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C H A P T E R

37 Surface Modification of Biomaterials Rachit Agarwal1, Andre´s J. Garcı´a2 1

Centre for BioSystems Science and Engineering, Indian Institute of Science, Bangalore, India; 2Woodruff School of Mechanical Engineering and Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States

INTRODUCTION Biomaterial Interfaces in Regenerative Medicine Biomaterials, whether synthetic (e.g., polymers, metals, ceramics) or natural (e.g., proteins polysaccharides), have central roles in tissue engineering and regenerative medicine applications by providing three-dimensional scaffolds to support cellular activities, matrices for the delivery of therapeutic agents (e.g., drugs, proteins, DNA, small interfering RNA), and functional device components (e.g., mechanical supports, sensing/stimulating elements, nonthrombogenic surfaces, diffusional barriers). The bulk properties of the biomaterial are critical determinants of the biological performance of the material [1]. For example, the mechanical properties of a vascular substitute, including elastic modulus, ultimate tensile stress, and compliance, dictate the ability of this tissue construct to support the applied mechanical loads associated with blood flow. On the other hand, the biological response to a biomaterial is governed by the material surface properties, primarily surface chemistry and structure. Protein adsorption or activation and cell adhesion, events that regulate host responses to materials, occur at the biomaterialetissue interface, and the physicochemical properties of the material surface modulate these biological events [2,3]. For instance, the chemical properties of the surface of a vascular substitute control blood compatibility (i.e., protein adsorption, platelet adhesion, thrombogenicity, patency). Hence, modification of biomaterial surfaces represents a promising route to engineer biofunctionality at the materialetissue interface to modulate biological responses without altering material bulk properties.

Overview of Surface Modification Strategies Numerous surface modification approaches have been developed for all classes of materials to modulate biological responses and improve device performance. Applications include the reduction of protein adsorption and thrombogenicity, control of cell adhesion, growth and differentiation, modulation of fibrous encapsulation and osseointegration, improved wear and/or corrosion resistance, and potentiation of electrical conductivity [1]. Surface modifications fall into two general categories: (1) physicochemical modifications involving alterations to the atoms, compounds, or molecules or topography on the surface; and (2) surface coatings consisting of a different material from the underlying support. Physicochemical modifications include chemical reactions (e.g., oxidation, reduction, silanization, acetylation), etching, and mechanical roughening or polishing and patterning (Fig. 37.1). Overcoating alterations are composed of grafting (including tethering of biomolecules), noncovalent and covalent coatings, and thin-film deposition (Fig. 37.2). Whereas the specific requirements of the surface modification approach vary with application, several characteristics are generally desirable. Thin surface modifications are preferred for most applications because thicker coatings often negatively influence the mechanical and functional properties of the material. Ideally, the ˚ ), but in practice, thicker surface modification should be confined to the outermost molecular layer (w10e15 A layers (10e100 nm) are used to ensure uniformity, durability, and functionality. Stability of the modified surface Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00037-0

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37. SURFACE MODIFICATION OF BIOMATERIALS

FIGURE 37.1 Schematic representations of common physicochemical surface modifications of biomaterials.

FIGURE 37.2

Schematic representations of common overcoating technologies for surface modification.

PHYSICOCHEMICAL SURFACE MODIFICATIONS

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is a critical requirement for adequate biological performance. Surface stability refers to mechanical durability (i.e., resistance to cracking, delamination, and debonding) but also to chemical stability, especially in aggressive, chemically active environments such as the biological milieu. Several types of surface rearrangements, such as translation of surface atoms or molecules in response to environmental factors and mobility of bulk molecules to the surface, and vice versa, readily occur in polymers and ceramics after exposure to biological fluids. Given the uniquely reactive nature and mobility or rearrangement of surfaces, as well as the tendency of surfaces to contaminate readily, rigorous analyses of surface treatments are essential to surface modification strategies. Surface analyses technologies generally focus on characterizing topography, chemistry or composition, and surface energy [4] (Table 37.1). Important considerations for these surface analyses technologies include operational principles (impact of high-energy particles or X-rays under ultrahigh vacuum, adsorption, or emission spectroscopies), depth of analysis, sensitivity, and resolution. For most applications, several analyses techniques must be used to obtain a complete description of the surface.

PHYSICOCHEMICAL SURFACE MODIFICATIONS Physicochemical modifications involve alterations to the atoms, compounds, or molecules and/or topography on the material surface (Fig. 37.1).

Chemical Modifications Countless chemical reactions, including UV and laser irradiation and etching reactions to clean, alter, or cross-link surface groups, have been developed to modify biomaterial surfaces [1]. Nonspecific reactions yield a distribution of chemically distinct groups at the surface; the resulting surface is complex and difficult to characterize owing to the presence of different chemical species in various concentrations. Nevertheless, nonspecific chemical reactions are widely used in biomaterials processing. Examples of nonspecific reactions include radio-frequency glow discharge in different plasmas (e.g., oxygen, nitrogen, argon), corona discharge in air, oxidation of metals, and acidebase treatments of polymers. In contrast, specific chemical reactions target particular chemical moieties on the surface to convert them into another functional group with minimal side (unwanted) reactions. Acetylation, fluorination of hydroxylated surfaces via trifluoroacetic anhydrides, silanization of hydroxylated surfaces, and incorporation of glycidyl groups into polysiloxanes are examples of specific chemical reactions. In addition, various chemical methods exist to tether biomacromolecules onto available anchoring groups on surfaces, as described in the Biological Modification of Surfaces section. The reaction of metal surfaces to produce an oxide-rich layer that conveys corrosion resistance, passivation, and improved wear and adhesive properties (also referred to as conversion coatings) is a common surface modification in metallic biomaterials. For example, nitric acid treatment of titanium and titanium alloys to generate titanium oxide layers is regularly performed on titanium-based medical devices, and the excellent biocompatibility properties of titanium are attributed to this oxide layer [5]. Implantation of ions into surfaces by beaming accelerated ions has been applied to modify the surface properties of metals and ceramics. For example, ion beam implantation of nitrogen into titanium and boron and carbon into stainless steel improves wear resistance and fatigue life, respectively [6]. In addition, evidence suggests that ion beam implantation of silicone and silver can enhance the blood compatibility and infection resistance of silicone rubber catheters [7,8].

Topographical Modifications The size and shape of topographical features on a surface influence cellular and host responses to the material. For example, surface macrotexture and microtexture alter cell adhesion, spreading, and alignment [9,10] and can regulate cell phenotypic activities, including neurite extension and osteoblastic differentiation [11,12]. Moreover, surface topography can have significant in vivo effects. For instance, the implant porosity modulates bone and soft tissue ingrowth [13,14], and the surface texture alters epithelial downgrowth responses to percutaneous devices and inflammatory reactions and fibrous encapsulation to materials implanted subcutaneously [15e17]. Although specific surface texture parameters that elicit particular biological responses have been identified in several cases, the mechanisms generating these behaviors remain poorly understood. Methods for generating surface texture can be grouped into approaches for engineering either roughness or topography (Fig. 37.3). Surface roughness indicates a random or complex pattern of features of varying amplitude and spacing, typically on a scale smaller than a cell (10e20 mm). On the other hand, surface topography refers to

TABLE 37.1

Common Surface Analysis Techniques Chemical Composition

Spatial Resolution

Depth

Sensitivity

Texture

Principle

Operation

Contact angle

Liquid wetting of surfaces

Air Liquid

NA

˚ 3e20 A

NA

Indirect

Atomic force microscopy

Records interatomic forces between tip and sample

Air Aqueous

Atomic

NA

Single atom

Yes

No

No

No

Scanning electron microscopy

Secondary electron emission caused by electron bombardment is imaged

Vacuum

˚ 40 A

˚ 5e10 A

High

Yes

No

No

No

Energy-dispersive X-ray spectroscopy

X-ray emission caused by electron bombardment

Vacuum

˚ 40 A

1 mm

10

No

Z>5

No

No

Auger electron spectroscopy

Auger electron emission caused by electron bombardment

Vacuum

˚ 100 A

˚ 15e50 A

10 10 g/cm2 0.1 atm%

No

Z>3

Chemical shift

No

X-ray photoelectron spectroscopy

X-rays cause emission of photoelectrons with characteristic energies

Vacuum

10 mm

˚ 10e150 A

10 10 g/cm2 0.1 atm%

No

Z>3

Chemical shift (excellent)

No

Secondary-ion mass spectrometry

Ion bombardment causes secondary ion emission

Vacuum

3-10 mm

˚ 10 A

10

No

All

Yes

Yes

Fourier transform infrared (IR) resonance

Molecular vibrations resulting from adsorption of IR radiation

Air Aqueous (attenuated total reflection)

10 mm

70%) with average pore sizes larger than 300 mm [31]. However, in skin regeneration, successful scaffolds need only to exhibit pore sizes of 20e125 mm [32]. This discrepancy can be explained by the low vascular requirements of the skin and its convenient juxtaposition to an ample supply of atmospheric oxygen. However, there is an upper limit in the porosity and pore size set by constraints associated with mechanical properties. An increase in the void volume results in a reduction in the mechanical strength of the scaffold, which can be detrimental in applications where regenerated tissues must support significant mechanical loads (e.g., long bones, heart valves, and articular cartilage) [13]. The extent to which the porosity of a scaffold can be increased while allowing it to meet tissue mechanical requirements depends on many factors, including the intrinsic makeup of the biomaterial and the processing conditions used in fabrication [31]. As histogenesis progresses and gives way to organogenesis, the impact of the scaffold pore structure on material degradation and tissue vascularization becomes apparent. The size and distribution of pores within a scaffold greatly influence the manner and rate of in vivo degradation [33], which can affect tissue formation and construct mechanical integrity. In materials susceptible to hydrolytic cleavage, for example, the access of water molecules to the interior of a scaffold is limited by porosity. Similar parallels exist for matrices subject to enzymatic degradation, which rely on interaction with cell-secreted molecules for dissolution. A final, important consideration is the influence of pore structure on the establishment of a blood supply in newly developing tissue. In early stages of histogenesis, nutrients, metabolites and other factors essential to cell survival pass freely through scaffold pores. As these pores fill with new tissue, however, a functioning vasculature is necessary. New “designer” tissueengineered scaffolds are composed of precisely controlled porous architectures that support and guide vessel ingrowth during tissue development [34,35]. The concepts of degradation and microvascularization are discussed in more detail in the following sections.

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Degradation Although nondegradable biomaterials have had success in many medical devices, complication primarily caused by chronic foreign body responses have yet to be overcome. For this reason, the ideal regeneration construct is one that can eventually be replaced by native tissues. Furthermore, the degradation rate of the construct is intrinsic to the success of the implant. This means that the material dissolution should complement tissue synthesis to ensure suitable mechanical stability during the process of histogenesis. The necessary scaffold residence time is tissue specific and must account for the time required for cells to populate the scaffold adequately and deposit a stable ECM. If a biomaterial degrades before sufficient ECM deposition has occurred, cells will lose important physiochemical factors for tissue regeneration and repair is likely to occur, resulting in scar formation. However, if the scaffold residence time is too long, ECM deposition and cell proliferation will be suppressed. An important balance must be established for successful regeneration. In most currently employed scaffold materials, degradation in the in vivo aqueous environment occurs via hydrolysis of chemical bonds in the base material. Chemical functionalities, molecular weight, and the degree of crosslinking determine the degradation characteristics. For example, higheremolecular weight materials tend to degrade more slowly over time as do materials with a higher hydrophobicity and crystallinity. Using a combination of these factors, predictable degradation profiles can be developed to match expected tissue formation rates. However, the consequences of material dissolution must be considered. As mentioned earlier, scaffolds that undergo bulk erosion can rapidly become unstable owing to the formation of large pores with low mechanical stability [36]. In addition, the degradation products of some scaffolds can be toxic not only to cells of the surrounding tissue, but also to vital organs of the lymphatic system. For example, degradation of the frequently studied polylactic acid (PLA) and polyglycolic acid (PGA) scaffolds results in a marked drop in pH in the local vicinity because of the release of acidic degradation products [36,37]. The decrease in pH can be detrimental to cells and organs; over time, it can lead to an inflammatory response with possible capsule formation and tissue necrosis [38]. As an alternative to hydrolytic degradation, many investigators are developing smart materials that can be dynamically remodeled during histogenesis via cell-mediated processes. These scaffolds are designed to mimic the degradation of natural ECM proteins, which are subject to matrix metalloproteinases (MMPs) and serine proteases that are either secreted or activated by most cell types. Because proteolysis-induced degradation is required for cell migration and invasion, researchers have had success in introducing synthetic hydrogels that are sensitive to cell proteases. Hydrogels containing amino acid sequences that can be degraded by plasmin [39], MMPs [40], or both of these protease families [41e43] exhibit sustained degradation upon cellular infiltration. In our own laboratory, we have fabricated MMP-degradable hydrogels that become fluorescent when degraded by cell proteases [44]. These ¼ polyethylene glycol (PEG)-based hydrogels are synthesized with MMP-degradable segments in the polymer backbone that are labeled with fluorescent, self-quenching tags. Thus, intact substrates show no fluorescence, but upon degradation by cell proteases, quantifiable fluorescence is emitted. Cells seeded in these fluorogenic substrates are able to cleave the degradable hydrogel matrix, as visualized by a marked increase in fluorescence in the areas immediately around the cell (Fig. 38.5). In addition, cell migration trails could be seen in the hydrogels. It is believed these materials will contribute to an understanding of cell migration and cell-mediated scaffold degradation.

Fibroblast encapsulated within a fluorogenic substrate. Differential interference contrast image (left) and fluorescent image (right) showing green fluorescence generated by material degradation around cell (red).

FIGURE 38.5

DESIGN PARAMETERS FOR HISTOGENESIS

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Biomolecular Factors In many cases, seeding cells inside a porous scaffold is insufficient to induce tissue regeneration because the material does not contain the chemical cues necessary to promote cellular remodeling events. Thus, researchers have attempted to modify biomaterials actively at the molecular level by incorporating cell-specific biomolecules. One strategy is to encapsulate molecules such as peptides or proteins into biomaterial carriers so that these molecules can be released from the material to trigger or modulate new tissue formation [8]. Another approach involves physically or chemically modifying scaffolds with specific cell-binding peptides to increase cellular interaction with the substrate. Cell-binding peptides are short amino acid sequences derived from much longer native ECM proteins that have been identified as able to incur specific, predictable interactions with cell receptors. Essentially the peptides function to mimic the signaling dynamic between the ECM and cells, and because many synthetic scaffold materials are not inherently adhesive to cells, the introduction of such sequences can be critical to encouraging cell retention and subsequent tissue formation [45]. The most well-studied cell adhesion peptide, arginine-glycineaspartic acid-serine (RGDS) has been widely used to encourage cells of various types to interact with otherwise nonadhesive synthetic matrices [46] (Fig. 38.6). Other amino acid sequences have been found to promote adhesion by specific cell phenotypes including endothelial cells [29,47e49], smooth muscle cells [47], neural cells [50], and osteoblasts [51]. Various growth factors have also been employed in efforts to enhance the process of histogenesis. Because they have key roles in tissue differentiation and repair, molecules such as epidermal growth factor (EGF), platelet-derived growth factor (PDGF), and transforming growth factor-b1 (TGF-b1) are popular for use in these applications. The release of PDGF from a polyurethane scaffold has been shown to encourage wound healing of in a rat model [52], whereas Griffith and colleagues used tethered EGF to promote osteogenic differentiation in an effort to improve connective tissue regeneration [53]. TGF-b1 has many roles in histogenesis [54] and has received particular attention for aiding the differentiation of stem cells and the development of new vasculature in vivo [55]. In other work, basic fibroblast growth factor (bFGF) and nerve growth factor were immobilized in fibrin scaffolds to facilitate cellular recruitment and differentiation [56]. Biomolecules have also been covalently coupled to PEG-based materials [57e60] with vascular endothelial growth factor and showed the potential to drive endothelial cell tubulogenesis. Furthermore, Delong and West formed gradients of bFGF and observed cellular alignment and migration that was directly influenced by growth factor presentation [61] (Fig. 38.7).

Importance of Microvasculature One of the biggest limitations to histogenesis in 3D scaffolds is the lack of a functional vascular system. The most successful engineered materials to date have been for tissues such as skin, which are thin enough to be supported by diffusion from the host vascular, and cartilage, which is relatively avascular and as such contains cells that are tolerant of anoxic conditions. All other tissues require some form of vascular system to permit the long-term survival of cells within the material. There are two primary strategies for establishing a blood supply in implanted 3D scaffolds. Vessels can grow into the construct from host tissue or they can be preformed in vitro and interconnect with host vascular upon implantation.

FIGURE 38.6 Peptide modification promotes cell adhesion. The nonadhesive nature of polyethylene glycol hydrogels (left) can be significantly altered by inclusion of an arginine-glycine-aspartic acid-serine peptide to promote cell attachment (right).

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FIGURE 38.7 Gradient of basic fibroblast growth factor (bFGF) immobilized in a hydrogel scaffold. Cells seeded on hydrogels containing a bFGF gradient aligned and migrated along the axis of growth factor immobilization.

The premise behind the first strategy is to encourage vessels to enter an avascular construct from the host tissue by stimulating angiogenesis. To encourage this process, scaffolds can be fabricated with precisely designed pore structures or surface chemistries that support ingrowth [34,62]. In some cases, proangiogenic factors are immobilized on or released from implants to encourage the ingrowth of vessels from host tissue. Work from our laboratory showed extensive infiltration of functional vessels into a cell-adhesive, proteolytically degradable, PEG-based hydrogel that had been implanted in the mouse cornea [63e66]. The limitation of these strategies is the time required for vessels to extend into the entirety of the engineered construct. At an extension rate of approximately 5 mm/h [67], new vessels will not reach the center of large scaffolds until several days after implant, leaving any cells at these locations without sufficient supplies of oxygen and nutrients. A second option is to create preformed vascular networks in vitro that are capable of anastomosing with host vascular upon implantation. This process of connecting two independent vascular networks is called inosculation and is the mechanism primarily responsible for the successes in plastic surgery and skin transplantations. To generate microvascular networks in vitro, researchers seed scaffolds with cells known to participate in vasculogenesis, including endothelial cells, stem cells, and pericytes. With appropriate biochemical and/or physical stimulation, these cells self-assemble into capillary-like structures. As an example, successful vascular networks have been formed by endothelial cells in fibrin scaffolds [68] and by adult and cord bloodederived progenitor cells in Matrigel [69]. In each study, the preformed vessels were functional upon implantation in vivo; and in the case of the endothelial cells, immature capillaries were further stabilized by host mural cells. Other work along these lines suggests that providing relevant physiomechanical stimulation in vitro will aid in developing functional prevascularized networks in engineered constructs [70e72]. A final approach exploits the body’s own ability to form blood vessels. A variety of scaffold materials have been implanted in highly vascularized anatomical sites, where they are incubated for up to 3 weeks as host vessels infiltrate. Once the vessel network is established, the scaffold is explanted, loaded with cells, and finally implanted into the site targeted for regeneration [73]. The need for adequate blood flow in 3D scaffolds is readily apparent and of concern to all researchers interested in regeneration strategies. Advances are promising, although much work remains to be done. Inosculation of preformed microvascular networks with host vascular, for instance, provides functional transport of engineered materials much more rapidly than some other methods, but it is still too slow for sensitive tissue applications. A combination of this approach with proangiogenic strategies may encourage connection in a shorter time, thus leading to the better long-term regeneration of target tissue.

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SYNTHETIC MATERIALS FOR HISTOGENESIS OF NEW ORGANS Biomaterials investigated as scaffolds for histogenesis include natural polymers such as collagen and fibrin, as well as a range of synthetic substrates. Although natural matrices have certain advantages in that their chemical composition is generally amenable to cell growth, batch-to-batch variations in substrate quality and performance can make their use in clinical regeneration applications problematic. As such, the control and flexibility of synthetic materials make them attractive alternatives. As alluded to previously, control is an important element in tailoring the scaffold’s material properties for appropriate cellematerial and matrixematerial interactions. The following sections highlight two popular classes of synthetic materials: hydrolytically degradable polymers and hydrogels.

Hydrolytically Degradable Polymers The most widely used polymers for cellular scaffold materials are PLA, PGA, or a combination of these two polymers (poly[lactic-co-glycolic acid] [PLGA]). PLA, PGA, and PLGA are aliphatic esters that possess good biocompatibility [74] and have been used as drug delivery materials to administer biomolecules during tissue regeneration [75,76]. These polymers are also among the few synthetic polymers approved by the US Food and Drug Administration (FDA) for certain human clinical applications. PGA is extremely hydrophilic in nature; consequently, it will lose its mechanical strength within 2e4 weeks of implantation [77]. PLA, however, contains an additional methyl group and as a result is more hydrophobic. Degradation of PLA scaffolds can take from months to years [75,78]. In addition, the degradation rates of these polymers can be tailored by using copolymer blends (PLGA), which give distinct degradation profiles [75,79]. However, these scaffolds undergo acid-catalyzed hydrolysis and bulk erosion, which have the potential to result in structural instability and interruption of the regeneration process [80]. Polyanhydrides have been synthesized for a number of biomedical applications including tissue engineering and drug delivery [81]. Polyanhydride scaffolds exhibit excellent biocompatibility and contain a large aliphatic component that possesses an ester group that makes the material subject to surface erosion [82]. This deliberate surface erosion is mechanistically different from bulk hydrolysis and can be exploited to synthesize biomaterials scaffolds that have predictable degradation profiles. In addition, the erosion of only the surface of the material allows anhydrides to maintain structural integrity in support of histogenesis. Because they exhibit mechanical properties similar to bone and are ideal scaffolds for tissue infiltration, anhydrides have been widely employed as scaffolds for in vivo bone regeneration [83e85]. Polyanhydride networks can also be combined with other polymers to change their degradation and structural characteristics. Jiang and Zhu [86] showed that anhydride polymers could be polymerized in the presence of PEG to form a cross-linked network with both hydrophobic and hydrophilic components. The hydrophilic PEG chains increase the uptake of water to drive the hydrolysis of the ester bond in the hydrophobic anhydride. As such, the degradation properties can be tailored by altering the amount of PEG in the scaffold material.

Hydrogels Hydrogels, which contain up to 90% water, are another widely studied class of materials for tissue regeneration. These materials are appealing because the polymer properties are controllable and reproducible [87] and the large water uptake promotes excellent biocompatibility. In many cases, hydrogel mechanical properties resemble those of native tissue and can be systematically controlled for specific applications. In addition, several hydrogel monomers contain vinyl chemical moieties, which are conducive to various free radicaleinitiated polymerizations schemes that can be employed to generate solid substrate materials. Photoinitiation, for example, enables polymers to be formed using specific wavelengths of light. Using this method, many researchers have had success forming complex 3D structures of varying stiffnesses. For example, polyacrylamide hydrogels have been shown to induce regeneration of soft tissue in facial defects [88], and 2-hydroxyethylmethacrylate has had good success as a fibrillar support for nerve regeneration [89]. Among the most well-studied hydrogel materials is cross-linked PEG, which has been approved by the FDA for use in certain medical applications [90]. As with other hydrogels, the hydrophilic nature of PEG discourages cell and

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protein adhesion and therefore results in a low instance of immunorejection by the host. By changing the monomer chain length, adding biological molecules, or using copolymers, researchers have generated a wide array of PEG hydrogel formulations suitable for many different tissue engineering applications. To render an otherwise blank slate amenable to histogenesis, various biomimetic peptides and growth factors have been incorporated into the PEG hydrogel matrix [57,59,91e93]. These modifications have been successful in achieving selective cell adhesion and promoting the accumulation of secreted tissue matrix. As mentioned previously, similar methods have been employed to encourage vasculogenesis and angiogenesis in PEG hydrogel materials. In addition, these hydrogel materials show promise as small-diameter vascular grafts [94]. Incorporation of peptides subject to proteolytic cleavage in the backbone of the PEG monomer renders the scaffolds subject to cell-mediated remodeling, giving these materials an additional advantage as histogenesis conduits.

FUTURE DIRECTIONS IN THREE-DIMENSIONAL SCAFFOLDS: THREE-DIMENSIONAL MICROFABRICATION Advances in the biomaterial field are providing tissue engineers with the means to generate complex and highly specialized 3D scaffolds. One of the earliest examples of such architectures was developed by Griffith and colleagues for hepatocyte culture and liver regeneration. Using a rapid printing technique, microporous PLGA scaffolds were fabricated by directing solvent streams onto polymer granules in a precisely controlled manner [95]. The hepatocytes seeded upon these constructs exhibited increased metabolic rates that mimicked cells in vivo more closely. In other work, 3D, microporous PLGA foams were prepared by drilling with dies of a specific size. The dimensions of these cylindrical scaffolds were reproducible with millimeter precision, and when placed in vivo, the materials supported bone regeneration in nonhealing defect models [96,97]. Porous scaffolds have also been micropatterned for vascular tissue engineering applications [35,98]. Several researchers have used photopolymerization techniques to mold and pattern hydrogel scaffolds for better control of cellesubstrate interactions [99]. Peppas reported micropatterning of PEG hydrogels using UV polymerization to generate many different substrate morphologies on the order of 100 mm [100]. Liu and Bhatia also photopatterned PEG hydrogels using a layer-by-layer method to generate a 3D scaffold for hepatocytes [101]. Laser-based patterning of hydrogels is a relatively new technique for generating complex 3D microenvironments inside hydrogel materials and natural constructs. Liu et al. used a laser ablation technique to form lines, holes, and interconnected grids in collagen matrices [102], whereas growth factors and peptides were patterned by Roy and colleagues using laser-based stereolithography [103] inside a PEG hydrogel. Biomolecules have also been patterned inside agarose hydrogels [104]. In this case, RGDS peptides were patterned in cylindrical shapes within the hydrogel material. After 3 days, neuronal cells seeded on the surface of the materials were shown to have migrated into the hydrogel in only the selectively patterned areas. Additional studies of cell migration in hydrogel materials were conducted with micropatterned, PEG-based materials functionalized with several different bioactive moieties [60,105,106]. In the process of laser scanning lithography, photosensitive peptides or proteins are covalently incorporated into 3D hydrogels with the precision of a confocal microscope laser (Fig. 38.8). The technique is capable of generating features from 1 mm to 1 mm and can be extended to include multiple bioactive moieties in a single substrate. The image in Fig. 38.9 illustrates the 3D nature of a patterned ligand. These precisely fabricated regeneration matrices provide great opportunities for controlled tissue growth.

CONCLUSIONS The need for replacement tissues and organs is driving tissue engineers to develop materials and strategies capable of generating biologically functional substitutes. The study of natural processes such as wound healing has provided insights into the complex mechanisms of tissue regeneration and allowed researchers to prioritize design parameters for 3D scaffolds. At the same time, advances in biomaterial synthesis and modification, as well as a better understanding of the signaling molecules important in tissue synthesis, are providing a wealth of tools for regeneration strategies. In a systematic approach to histogenesis, Nettles et al. developed a method of neural network analysis in which a self-organizing map delineates the relationships between scaffold parameters, such as cross-link density, and tissue outcomes [107]. The investigators employed this tool with the goal of optimizing and

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FIGURE 38.8 Laser scanning lithography patterning of polyethylene glycol (PEG) hydrogels. Precisely defined patterned areas are generated in three-dimensional (3D) hydrogels by using a confocal microscope laser to cross-link photosensitive materials. DA, diacrylate; RGDS, arginine-glycine-aspartic acid-serine. Schematic courtesy of Joseph C. Hoffman, Rice University.

FIGURE 38.9 Laser scanning lithography pattern of fluorescently labeled arginine-glycine-aspartic acid-serine in a polyethylene glycol hydrogel. The fluorescent peptide (red) is visible in the bulk hydrogel (black) after patterning. Scale bar ¼ 10 mm. Image courtesy of Joseph C. Hoffman, Rice University.

accelerating the design of a cartilage tissue substitute. Tools such as these will focus the work of tissue engineers going forward. Good success has been seen in developing substitutes for skin and cartilage. Advances in scaffold microvascularization techniques will aid in progressing the field to larger, more complex target tissues.

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[81] Burkoth AK, Anseth KS. A review of photocrosslinked polyanhydrides:: in situ forming degradable networks. Biomaterials 2000a;21(23): 2395e404. [82] Davis KA, Burdick JA, Anseth KS. Photoinitiated crosslinked degradable copolymer networks for tissue engineering applications. Biomaterials 2003;24(14):2485e95. [83] Anseth KS, Shastri VR, Langer R. Photopolymerizable degradable polyanhydrides with osteocompatibility. Nat Biotechnol 1999;17(2): 156e9. [84] Burkoth AK, Anseth KS. A review of photocrosslinked polyanhydrides:: in situ forming degradable networks. Biomaterials 2000b;21(23): 2395. [85] Muggli DS, Burkoth AK, Anseth KS. Crosslinked polyanhydrides for use in orthopedic applications: degradation behavior and mechanics. J Biomed Mater Res 1999;46(2):271e8. [86] Jiang HL, Zhu KJ. Preparation, characterization and degradation characteristics of polyanhydrides containing poly(ethylene glycol). Polym Int 1999;48(1):47e52. [87] Peppas NA. Devices based on intelligent biopolymers for oral protein delivery. Int J Pharm 2004;277(1e2):11e7. [88] von Buelow S, von Heimburg D, Pallua N. Efficacy and safety of polyacrylamide hydrogel for facial soft-tissue augmentation. Plast Reconstr Surg 2005;116(4):1137e46. [89] Flynn L, Dalton PD, Shoichet MS. Fiber templating of poly(2-hydroxyethyl methacrylate) for neural tissue engineering. Biomaterials 2003; 24(23):4265e72. [90] Drury JL, Mooney DJ. Hydrogels for tissue engineering: scaffold design variables and applications. Biomaterials 2003;24(24):4337e51. [91] DeLong SA, Gobin AS, West JL. Covalent immobilization of RGDS on hydrogel surfaces to direct cell alignment and migration. J Contr Release 2005a;109(1e3):139. [92] Gonzalez AL, Gobin AS, West JL, McIntire LV, Smith CW. Integrin interactions with immobilized peptides in polyethylene glycol diacrylate hydrogels. Tissue Eng 2004;10(11e12):1775e86. [93] Mann BK, Schmedlen RH, West JL. Tethered-TGF-[beta] increases extracellular matrix production of vascular smooth muscle cells. Biomaterials 2001b;22(5):439e44. [94] Hahn MS, McHale MK, Wang E, Schmedlen RH, West JL. Physiologic pulsatile flow bioreactor conditioning of poly(ethylene glycol)-based tissue engineered vascular grafts. Ann Biomed Eng 2007;35(2):190e200. [95] Kim SS, Utsunomiya H, Koski JA, et al. Survival and function of hepatocytes on a novel three-dimensional synthetic biodegradable polymer scaffold with an intrinsic network of channels. Ann Surg 1998;228(1):8e13. [96] Karp JM, Rzeszutek K, Shoichet MS, Davies JE. Fabrication of precise cylindrical three-dimensional tissue engineering scaffolds for in vitro and in vivo bone engineering applications. J Craniofac Surg 2003;14(3):317e23. [97] Karp JM, Sarraf F, Shoichet MS, Davies JE. Fibrin-filled scaffolds for bone-tissue engineering: an in vivo study. J Biomed Mater Res 2004; 71A(1):162e71. [98] Sarkar S, Lee GY, Wong JY, Desai TA. Development and characterization of a porous micro-patterned scaffold for vascular tissue engineering applications. Biomaterials 2006;27(27):4775e82. [99] Bryant SJ, Cuy JL, Hauch KD, Ratner BD. Photo-patterning of porous hydrogels for tissue engineering. Biomaterials 2007;28(19):2978e86. [100] Peppas NA, Ward JH. Biomimetic materials and micropatterned structures using iniferters. Adv Drug Deliv Rev 2004;56(11):1587e97. [101] Tsang VL, Chen AA, Cho LM, et al. Fabrication of 3D hepatic tissues by additive photopatterning of cellular hydrogels. FASEB J 2007;21(3): 790e801. [102] Liu YM, Sun S, Singha S, Cho MR, Gordon RJ. 3D femtosecond laser patterning of collagen for directed cell attachment. Biomaterials 2005; 26(22):4597e605. [103] Mapili G, Lu Y, Chen SC, Roy K. Laser-layered microfabrication of spatially patterned functionalized tissue-engineering scaffolds. J Biomed Mater Res B Appl Biomater 2005;75B(2):414e24. [104] Luo Y, Shoichet MS. A photolabile hydrogel for guided three-dimensional cell growth and migration. Nat Mater 2004;3(4):249e53. [105] Hahn MS, Miller JS, West JL. Laser scanning lithography for surface micropatterning on hydrogels. Adv Mater 2005;17(24):2939. [106] Lee SH, Moon JJ, West JL. Three-dimensional micropatterning of bioactive hydrogels via two-photon laser scanning photolithography for guided 3D cell migration. Biomaterials 2008;29(20):2962e8. [107] Nettles DL, Haider MA, Chilkoti A, Setton LA. Neural network analysis identifies scaffold properties necessary for in vitro chondrogenesis in elastin-like polypeptide biopolymer scaffolds. Tissue Eng Part A 2009;16(1):11e20.

C H A P T E R

39 Biocompatibility and Bioresponse to Biomaterials James M. Anderson Case Western Reserve University, Cleveland, OH, United States

INTRODUCTION Biocompatibility is generally defined as the ability of a biomaterial or medical device to perform with an appropriate host response in a specific application. A bioresponse or biocompatibility assessment (i.e., an evaluation of biological responses) is considered to be a measure of the magnitude and duration of the adverse alterations in homeostatic mechanisms that determine the host response. From a practical point of view, the evaluation of biological responses to a medical device is carried out to determine whether the medical device performs as intended and presents no significant harm to the patient. The goal of bioresponse evaluation is to predict whether a biomaterial or medical device presents potential harm to the patient. In regenerative medicine, biomaterials are used in a wide variety of ways ranging from carriers of genetic material to tissue-engineered implants that may contain autologous, allogeneic, or xenogeneic genetic materials, cells, and scaffold materials. Scaffolds may be composed of synthetic or modified-natural materials. A tissue-engineered implant is a biologicalebiomaterial combination in which some component of tissue has been combined with a biomaterial to create a device to restore or modify tissue or organ function. Thus, tissue-engineered devices with a biological component(s) require an expanded perspective and understanding of biocompatibility and biological response evaluation. The purpose of this chapter is to provide an overview of this expanded perspective. Each unique tissue-engineered device requires a distinctive set of experiments to determine its biological responses and biocompatibility. This chapter presents an overview of host responses that must be considered in determining the biocompatibility of tissue-engineered devices that employ biomaterials. The three major responses that must be considered for biocompatibility assessment are inflammation, wound healing, and immunological reactions or immunity. For the purposes of biological response evaluation, immunological reactions or immunity are considered to be immunotoxicity. Pathologists use the terminology of inflammation and immunity to describe adverse tissue reactions, whereas immunologists commonly refer to inflammation as innate immunity and activation of the immune system as being acquired immunity. Tissueematerial interactions are a series of responses that are initiated by the implantation procedure, as well as by the presence of the biomaterial, medical device, or tissue-engineered device. In this chapter, we divide the series of tissueematerial responses into inflammation (innate immunity) and wound healing, and immunotoxicity. After implantation, early, transient tissueematerial responses include injury (implantation), bloodematerials interactions, provisional matrix formation, and the temporal sequence of inflammation and wound healing including acute inflammation, chronic inflammation, granulation tissue development, foreign body reaction, and ultimately fibrosis or fibrous capsule (scar) development. Immunotoxicity is any adverse effect on the function or structure of the immune system or other systems as a result of an immune system dysfunction. Two significant failure mechanisms of tissue-engineered devices are fibrosis or fibrous capsule (scar) development surrounding and infiltrating the tissue-engineered device, or the initiation of acquired or cellular immunity by the biological component of the tissue-engineered device. Also, the biological component and the biomaterial component in a tissueengineered device may act in concert or synergistically to facilitate either of these failure mechanisms.

Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00039-4

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Copyright © 2019 Elsevier Inc. All rights reserved.

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INFLAMMATION (INNATE IMMUNITY) AND WOUND HEALING The process of implanting a biomaterial or tissue-engineered device results in injury to tissues or organs [1e8]. This injury and the subsequent perturbation of homeostatic mechanisms lead to inflammatory responses, foreign body reaction, and wound healing. The response to injury depends on multiple factors that include the extent of injury, loss of basement membrane structures, bloodematerial interactions, provisional matrix formation, the extent or degree of cellular necrosis, and the extent of the inflammatory response. The organ or tissue undergoing implantation may have a significant role in the response. These events, in turn, may affect the extent or degree of granulation tissue formation, foreign body reaction, and fibrosis or fibrous capsule (scar) development [9]. These events are summarized in Table 39.1. These host reactions for biocompatible biomaterials are considered to be normal. They are also tissue-, organ-, and species-dependent. These dependencies thus provide perspectives on the biological response evaluation and the ultimate determination of biocompatibility. The reactions occur or are initiated early (that is, within 2e3 weeks of the time of implantation) and undergo resolution quickly, leading to fibrosis or fibrous capsule formation.

BloodeMaterial Interactions and Initiation of the Inflammatory Response Bloodematerial interactions and the inflammatory response are intimately linked; in fact, early responses to injury involve mainly blood and the vasculature [1e8]. Regardless of the tissue into which a biomaterial is implanted, the initial inflammatory response is activated by injury to vascularized connective tissue. Because blood and its components are involved in the initial inflammatory responses, thrombus, blood clot, or both also form. Thrombus formation involves activation of the extrinsic and intrinsic coagulation systems, the complement system, the fibrinolytic system, the kinin-generating system, and platelets. Thrombus or blood clot formation on the surface of a biomaterial is related to the well-known Vroman effect of protein adsorption. From a wound healing perspective, blood protein deposition on a biomaterial surface is described as provisional matrix formation. Although injury initiates the inflammatory response, released chemicals from plasma, cells, and injured tissue mediate the response [4,6,10,11]. Important classes of chemical mediators of inflammation are presented in Table 39.2. Several important points must be noted to understand the inflammatory response and how it relates to biomaterials. First, although chemical mediators are classified on a structural or functional basis, different mediator systems interact and provide a system of checks and balances regarding their respective activities and functions. Second, chemical mediators are quickly inactivated or destroyed, which suggests that their action is predominantly local (i.e., at the implant site). Third, generally acid, lysosomal proteases, and oxygen-derived free radicals produce the most significant damage or injury. These chemical mediators are also important in the degradation of biomaterials. Phagolysosomes in macrophages can have acidity as low as pH 4, and direct microelectrode studies of this acidic environment have determined pH levels to be as low as 3.5. Moreover, only several hours are necessary to achieve these acid levels after adhesion of macrophages [12e14]. The predominant cell type present in the inflammatory response varies with the age of the injury. In general, neutrophils, commonly called polymorphonuclear leukocytes or polys, predominate during the first several days after injury and then are replaced by monocytes as the predominant cell type. Three factors account for this change in cell type [1]: Neutrophils are short-lived and disintegrate and disappear after 24e48 h; neutrophil emigration is of TABLE 39.1

Sequence of Host Reactions

Injury Bloodematerial interactions Provisional matrix formation Acute inflammation Chronic inflammation Granulation tissue Foreign body reaction Fibrosis/fibrous capsule development

INFLAMMATION (INNATE IMMUNITY) AND WOUND HEALING

TABLE 39.2

677

Important Chemical Mediators of Inflammation Derived From Plasma, Cells, or Injured Tissue

Mediators

Examples

Vasoactive agents

Histamine, serotonin, adenosine, endothelial-derived relaxing factor, prostacyclin, endothelin, thromboxane a2

Plasma proteases Kinin system

Bradykinin, kallikrein

Complement system

C3a, C5a, C3b, C5beC9

Coagulation/fibrinolytic system

Fibrin degradation products, activated Hageman factor (FXIIA), tissue plasminogen activator

Leukotrienes

Leukotriene B4, hydroxyeicosatetraenoic acid

Lysosomal proteases

Collagenase, elastase

Oxygen-derived free radicals

H2O2, superoxide anion, nitric oxide

Platelet-activating factors

Cell membrane lipids

Cytokines

Interleukin-1, tumor necrosis factor

Growth factors

Platelet-derived growth factor, fibroblast growth factor, transforming growth factor-a or b, epithelial growth factor

short duration because chemotactic factors for neutrophil migration are activated early in the inflammatory response [2]. After emigration from the vasculature, monocytes differentiate into macrophages, and these cells are very long-lived (up to months) [3]. Monocyte emigration may continue for days to weeks, depending on the injury and implanted biomaterial, and chemotactic factors for monocytes are activated over longer periods of time.

Provisional Matrix Formation Injury to vascularized tissue in the implantation procedure leads to immediate development of the provisional matrix at the implant site. This provisional matrix consists of fibrin, produced by activation of the coagulative and thrombosis systems, and inflammatory products released by the complement system, activated platelets, inflammatory cells, and endothelial cells [15e18]. These events occur early, within minutes to hours after implantation of a medical device. Components within or released from the provisional matrix (that is, fibrin network [thrombosis or clot]) initiate the resolution, reorganization, and repair processes such as inflammatory cell and fibroblast recruitment. Platelets activated during the fibrin network formation release platelet factor 4, platelet-derived growth factor (PDGF), and transforming growth factor b (TGF-b), which contribute to fibroblast recruitment [19,20]. Upon activation, monocytes and lymphocytes generate additional chemotactic factors including leukotriene B4 (LTB4), PDGF, and TGF-b to recruit fibroblasts. The provisional matrix is composed of adhesive molecules such as fibronectin and thrombospondin bound to fibrin as well as platelet granule components released during platelet aggregation. Platelet granule components include thrombospondin, released from the platelet a-granule, and cytokines including TGF-a, TGF-b, PDGF, platelet factor 4, and platelet-derived endothelial cell growth factor. The provisional matrix is stabilized by the cross-linking of fibrin by factor XIIIa. The provisional matrix appears to provide both structural and biochemical components to the process of wound healing. The complex three-dimensional structure of the fibrin network with attached adhesive proteins provides a substrate for cell adhesion and migration. The presence of mitogens, chemoattractants, cytokines, and growth factors within the provisional matrix provides for a rich milieu of activating and inhibiting substances for various cellular proliferative and synthetic processes. The provisional matrix may be viewed as a naturally derived, biodegradable, sustained release system in which mitogens, chemoattractants, cytokines, and growth factors are released to control subsequent wound healing processes [21e27]. Despite the rapid increase in our knowledge of the provisional matrix and its capabilities, our knowledge of the control of the formation of the provisional matrix and its effect on subsequent wound healing events is poor.

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Temporal Sequence of Inflammation and Wound Healing Inflammation is generally defined as the reaction of vascularized living tissue to local injury. Inflammation serves to contain, neutralize, dilute, or wall off the injurious agent or process. In addition, it sets into motion a series of events that may heal and reconstitute the implant site by replacing the injured tissue by regenerating native parenchymal cells, forming fibroblastic scar tissue, or a combination of these processes [4,6]. The sequence of events after implantation of a biomaterial is illustrated in Fig. 39.1. The size, shape, and chemical and physical properties of the biomaterial and the physical dimensions and properties of the prosthesis or device may be responsible for variations in the intensity and time duration of the inflammatory and wound healing processes. Thus, the intensity and/or time duration of the inflammatory reaction may characterize the biocompatibility of a biomaterial or device. Classically, the biocompatibility of an implanted material has been described in terms of the morphological appearance of the inflammatory reaction to the material; however, the inflammatory response is a series of complex reactions involving various types of cells, the densities, activities, and functions of which are controlled by various endogenous and autacoid mediators. The simplistic view of the acute inflammatory response progressing to the chronic inflammatory response may be misleading with respect to biocompatibility studies and the inflammatory response to implants. In vivo studies using the cage implant system show that monocytes and macrophages are present in highest concentrations when neutrophils are also at their highest concentrations: that is, the acute inflammatory response [28,29]. Neutrophils have short lifetimes (hours to days) and disappear from the exudates more rapidly than do macrophages, which have lifetimes of days to weeks to months. Eventually, macrophages become the predominant cell type in the exudates, resulting in a chronic inflammatory response. Monocytes rapidly differentiate into macrophages, the cells principally responsible for normal wound healing in the foreign body reaction. Classically, the development of granulation tissue has been considered to be part of chronic inflammation, but because of unique tissueematerial interactions, it is preferable to differentiate the foreign body reaction, with its varying degree of granulation tissue development, including macrophages, fibroblasts, and capillary formation, from chronic inflammation.

Sequence of events involved in inflammatory and wound healing responses leading to foreign body giant cell (FBGC) formation. This shows the potential importance of mast cells in the acute inflammatory phase and T-helper 2 (Th2) lymphocytes in the transient chronic inflammatory phase with the production of interleukin (IL)-4 and IL-13, which can induce monocyte/macrophage fusion to form FBGCs. PMN, polymorphonuclear.

FIGURE 39.1

INFLAMMATION (INNATE IMMUNITY) AND WOUND HEALING

679

Acute Inflammation Acute inflammation is of relatively short duration, lasting from minutes to days, depending on the extent of injury. The main characteristics of acute inflammation are the exudation of fluid and plasma proteins (edema) and the emigration of leukocytes (predominantly neutrophils). Neutrophils and other motile white cells emigrate or move from the blood vessels to the perivascular tissues and the injury (implant) site [30e32]. The accumulation of leukocytes, in particular neutrophils and monocytes, is the most important feature of the inflammatory reaction. Leukocytes accumulate through a series of processes including margination, adhesion, emigration, phagocytosis, and extracellular release of leukocyte products [33]. Increased leukocytic adhesion in inflammation involves specific interactions between complementary “adhesion molecules” present on the leukocyte and endothelial surfaces [34,35]. The surface expression of these adhesion molecules is modulated by inflammatory agents; mechanisms of interaction include stimulation of leukocyte adhesion molecules (C5a and LTB4), stimulation of endothelial adhesion molecules (interleukin-1 [IL-1]), or both effects of tumor necrosis factor-a (TNF-a). Integrins are composed of a family of transmembrane glycoproteins that modulate cellematrix and cellecell relationships by acting as receptors to extracellular protein ligands and also as direct adhesion molecules [36]. An important group of integrins (adhesion molecules) on leukocytes include the CD11/CD18 family of adhesion molecules. Inflammatory mediators (i.e., cytokines) stimulate a rapid increase in these adhesion molecules on the leukocyte surface as well as increased leukocyte adhesion to endothelium. Leukocyteeendothelial cell interactions are also controlled by endothelialeleukocyte adhesion molecules (also known as E-selectins) or intracellular adhesion molecules (ICAMs) such as ICAM-1 and ICAM-2, and vascular cell adhesion molecules on endothelial cells [37]. Inflammatory cell emigration is controlled in part by chemotaxis, which is the unidirectional migration of cells along a chemical gradient. A wide variety of exogenous and endogenous substances have been identified as chemotactic agents [11,30e41]. Important to the emigration or movement of leukocytes is the presence of specific receptors for chemotactic agents on the cell membranes of leukocytes. These and other receptors also may have a role in activating leukocytes. After localization of leukocytes at the injury (implant) site, phagocytosis and the release of enzymes occur following activation of neutrophils and macrophages. The major role of the neutrophils in acute inflammation is to phagocytose microorganisms and foreign materials. Phagocytosis is seen as a three-step process in which the injurious agent undergoes recognition and neutrophil attachment, engulfment, and killing or degradation. With regard to biomaterials, engulfment and degradation may or may not occur, depending on the properties of the biomaterial. Although biomaterials are not generally phagocytosed by neutrophils or macrophages because of the size disparity (i.e., the surface of the biomaterial is greater than the size of the cell), certain events may occur in phagocytosis. The process of recognition and attachment is expedited when the injurious agent is coated by naturally occurring serum factors called opsonins. The two major opsonins are immunoglobulin (Ig)G and the complement-activated fragment, C3b. Both of these plasma-derived proteins are known to adsorb to biomaterials, and neutrophils and macrophages have corresponding cell membrane receptors for these opsonization proteins. These receptors may also have a role in activating the attached neutrophil or macrophage. Because of the size disparity between the biomaterial surface and the attached cell, “frustrated phagocytosis” may occur [38,39]. This process does not involve engulfing the biomaterial but it causes the extracellular release of leukocyte products in an attempt to degrade the biomaterial. Neutrophils adherent to complement-coated and Ig-coated nonphagocytosable surfaces may release enzymes by direct extrusion or exocytosis from the cell [38,39]. The amount of enzyme released during this process depends on the size of the polymer particle, with larger particles inducing greater amounts of enzyme release. This suggests that the specific mode of cell activation in the inflammatory response in tissue depends on the size of the implant and that a material in a phagocytosable form (e.g., powder or particulate) may provoke a degree of inflammatory response different from that of the same material in a nonphagocytosable form (e.g., film). Tissue-engineered constructs containing biomaterial scaffolds alone or with cells and/or chemokines, growth factors, or other biological components are thus subjected to an aggressive microenvironment that may quickly compromise the intended function of the construct [42].

Chronic Inflammation Chronic inflammation is less uniform histologically than is acute inflammation. In general, chronic inflammation is characterized by the presence of monocytes and lymphocytes with the early proliferation of blood vessels and connective tissue [4,6,43e45]. Many factors modify the course and histological appearance of chronic inflammation.

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Persistent inflammatory stimuli lead to chronic inflammation. Although the chemical and physical properties of the biomaterial may lead to chronic inflammation, motion in the implant site by the biomaterial may also produce chronic inflammation. The chronic inflammatory response to biomaterials is confined to the implant site. Inflammation with the presence of mononuclear cells, including lymphocytes and plasma cells, is given the designation chronic inflammation, whereas the foreign body reaction with granulation tissue development is considered the normal wound healing response to implanted biomaterials (i.e., the normal foreign body reaction). Chronic inflammation with biocompatible materials is usually of short duration (i.e., a few days). Lymphocytes and plasma cells are involved principally in immune reactions and are important mediators of antibody production and delayed hypersensitivity responses. Their roles in nonimmunological injuries and inflammation are largely unknown [46e48]. Little is known regarding immune responses and cell-mediated immunity to synthetic biomaterials. The role of macrophages must be considered in the possible development of immune responses to synthetic biomaterials. Macrophages and dendritic cells process and present the antigen to immunocompetent cells and thus are critical mediators in the development of immune reactions. The macrophage is probably the most important cell in chronic inflammation because of the great number of biologically active products it produces [44]. Important classes of products yielded and secreted by macrophages include neutral proteases, chemotactic factors, arachidonic acid metabolites, reactive oxygen metabolites; complement components, coagulation factors, growth-promoting factors, and cytokines [49e51]. Growth factors such as PDGF, fibroblast growth factor, TGF-b, TGF-a/epithelial growth factor, and IL-1 or TNF are important to the growth of fibroblasts and blood vessels and the regeneration of epithelial cells. Growth factors released by activated cells stimulate the production of a wide variety of cells; initiate cell migration, differentiation, and tissue remodeling; and may be involved in various stages of wound healing [19,52e56]. There is a lack of information regarding interaction and synergy among various cytokines and growth factors and their abilities to exhibit chemotactic, mitogenic, and angiogenic properties.

Granulation Tissue Within 1 day after implantation of a biomaterial (i.e., injury), the healing response is initiated by the action of monocytes and macrophages, followed by the proliferation of fibroblasts and vascular endothelial cells at the implant site, leading to the formation of granulation tissue, the hallmark of healing inflammation. Granulation tissue derives its name from the pink, soft, granular appearance on the surface of healing wounds; its characteristic histological features include the proliferation of new small blood vessels and fibroblasts. Depending on the extent of injury, granulation tissue may be seen as early as 3e5 days after the implantation of a biomaterial. The new small blood vessels are formed by budding or sprouting of preexisting vessels in a process known as neovascularization or angiogenesis [45,57e60]. This process involves proliferation, maturation, and organization of endothelial cells into capillary tubes. Fibroblasts also proliferate in developing granulation tissue and are active in synthesizing collagen and proteoglycans. In the early stages of granulation tissue development, proteoglycans predominate; later, however, collagen (especially type I collagen) predominates and forms the fibrous capsule. Some fibroblasts in developing granulation tissue may have features of smooth muscle cells. These cells are called myofibroblasts and are considered to be responsible for the wound contraction seen during the development of granulation tissue.

Macrophage Interactions The inflammatory and immune systems overlap considerably through the activity and phenotypic expression of macrophages that are derived from blood-borne monocytes. Monocytes and macrophages belong to the mononuclear phagocytic system (MPS) (Table 39.3). Cells in the MPS may be considered resident macrophages in the respective tissues that take on specialized functions that depend on their tissue environment. From this perspective, the host defense system may be seen as blood-borne or circulating inflammatory and immune cells as well as mononuclear phagocytic cells that reside in specific tissues with specialized functions. In the inflammatory and immune responses, the macrophage has a pivotal role in both the induction and effector phases of these responses. Two factors that have a role in monocyte or macrophage adhesion and activation and foreign body giant cell (FBGC) formation are the surface chemistry of the substrate onto which the cells adhere and the protein adsorption that occurs before cell adhesion. These two factors have been hypothesized to have significant roles in the inflammatory and wound healing responses to biomaterials and medical devices in vivo.

681

INFLAMMATION (INNATE IMMUNITY) AND WOUND HEALING

TABLE 39.3

Mononuclear Phagocytic System

Tissues

Cells

Implant sites

Inflammatory macrophages, foreign body giant cells

Liver

Kupffer cells

Lung

Alveolar macrophages

Connective tissue

Histiocytes

Bone marrow

Macrophages

Spleen and lymph nodes

Fixed and free macrophages

Serous cavities

Pleural and peritoneal macrophages

Nervous system

Microglial cells

Bone

Osteoclasts

Skin

Langerhans cells and dendritic cells

Lymphoid tissue

Dendritic cells

MACROPHAGE

MONOCYTE BLOOD

CHEMOTAXIS MIGRATION

TISSUE

TISSUE/BIOMATERIAL

CHEMOTAXIS MIGRATION ADHESION DIFFERENTIATION

FOREIGN BODY GIANT CELL BIOMATERIAL

ADHESION DIFFERENTIATION SIGNAL TRANSDUCTION ACTIVATION

ACTIVITY PHENOTYPIC EXPRESSION

FIGURE 39.2 In vivo transition from blood-borne monocyte to biomaterial-adherent monocyte/macrophage to foreign body giant cell (FBGC) at the tissueebiomaterial interface. Little is known regarding the indicated biological responses that are considered to have important roles in the transition to FBGC development.

Macrophage interactions with biomaterials are initiated when blood-borne monocytes in the early transient responses migrate to the implant site and adhere to the blood protein adsorbed biomaterial through monocytee integrin interactions. After adhesion, adherent monocytes differentiate into macrophages that then may fuse to form FBGCs. Fig. 39.2 demonstrates the progression from circulating blood monocyte to tissue macrophage to FBGC development that is most commonly observed. Because of the progression of monocytes to macrophages to FBGCs (Fig. 39.2), the following discussion of macrophage interactions includes perspectives on how macrophages are formed (i.e., monocyte adhesion) and what happens to macrophages on biomaterial surfaces (i.e., FBGC formation) [61,62]. Material surface property-dependent blood protein adsorption occurs immediately upon surgical implantation of a biomaterial; it is the protein-modified biomaterial that inflammatory cells subsequently encounter. Monocytes express receptors for various blood components, but they recognize naturally occurring foreign surfaces by receptors for opsonins such as fragments of complement component C3. Complement activation by biomaterials has been well-documented [63]. Exposure to blood during biomaterial implantation may permit extensive opsonization with the labile fragment C3b and the rapid conversion of C3b to its hemolytically inactive but nevertheless opsonic and more stable form, C3bi. C3b is bound by the CD35 receptor, but C3bi is recognized by distinct receptors, CD11b/CD18 and CD11c/CD18, on monocytes [61]. Fibrinogen, a major plasma protein that adsorbs to biomaterials, is another ligand for these receptors that, together with CD11a/CD18, constitutes a subfamily of integrins that is restricted to leukocytes [61,62]. Studies with monoclonal antibodies to their common b2 subunit (CD18) and distinct a-chains have implicated CD11b/CD18 and CD11c/CD18 in monocyte/macrophage responses. Other potential adhesion-mediating proteins that adsorb to biomaterials include IgG, which may interact with monocytes via various receptors and fibronectin, for which monocytes also express multiple types of receptors [64e67].

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39. BIOCOMPATIBILITY AND BIORESPONSE TO BIOMATERIALS

Foreign Body Giant Cell Formation and Interactions The foreign body reaction is composed of FBGCs and the components of granulation tissue, which consist of macrophages, fibroblasts, and capillaries in varying amounts, depending on the form and topography of the implanted material [68]. Relatively flat and smooth surfaces, such as those found on breast prostheses, have a foreign body reaction that is composed of a layer of macrophages one to two cells thick. Relatively rough surfaces, such as those found on the outer surfaces of expanded poly(tetrafluoroethylene) vascular prostheses or poly(methyl methacrylate) bone cement, have a foreign body reaction composed of several layers of macrophages and FBGCs at the surface. Fabric materials generally have a surface response composed of macrophages and FBGCs with varying degrees of granulation tissue subjacent to the surface response. As previously discussed, the form and topography of the surface of the biomaterial determine the composition of the foreign body reaction. With biocompatible materials, the composition of the foreign body reaction in the implant site may be controlled by the surface properties of the biomaterial, the form of the implant, and the relationship between the surface area of the biomaterial and the volume of the implant. For example, high surface-to-volume implants such as fabrics or porous materials will have higher ratios of macrophages and FBGCs in the implant site than will smooth-surface implants, which will have fibrosis as a significant component of the implant site. The foreign body reaction consisting mainly of macrophages and/or FBGCs may persist at the tissueeimplant interface for the lifetime of the implant [1,2,69e71]. Generally, fibrosis (i.e., fibrous encapsulation) surrounds the biomaterial or implant with its interfacial foreign body reaction, isolating the implant and foreign body reaction from the local tissue environment. Early in the inflammatory and wound healing response, the macrophages are activated upon adherence to the material surface [50]. Although the chemical and physical properties of the biomaterial are generally considered to be responsible for macrophage activation, the nature of subsequent events regarding the activity of macrophages at the surface is not clear. Tissue macrophages, derived from circulating blood monocytes, may coalesce to form multinucleated FBGCs. FBGCs containing large numbers of nuclei are typically present on the surface of biomaterials. Although these FBGCs may persist for the lifetime of the implant, it is not known whether they remain activated, releasing their lysosomal constituents, or become quiescent. FBGCs have been implicated in the biodegradation of polymeric medical devices [72e74]. Fig. 39.1 demonstrates the sequence of events involved in inflammation and wound healing when medical devices are implanted. In general, the neutrophil (polymorphonuclear) predominant acute inflammatory response and the lymphocyte/monocyte predominant chronic inflammatory response resolve quickly (i.e., within 2 weeks), depending on the type and location of the implant. Studies using IL-4 demonstrate the role for type 2 (Th2) helper lymphocyte and mast cells in the development of the foreign body reaction at the tissueematerial interface [75]. Th2 helper lymphocytes have been described as “antiinflammatory” based on their cytokine profile, of which IL-4 is a significant component. Th2 helper lymphocytes also produce IL-13, which has an effect similar to that of IL-4 on FBGC formation. In this regard, anti-IL-4 antibody does not inhibit IL-13einduced FBGC formation, nor does anti-IL-13 antibody inhibit IL-4einduced FBGC formation. In IL-4 and IL-13 FBGC culture systems, the macrophage mannose receptor (MMR) has been identified as critical to the fusion of macrophages in the formation of FBGC [76,77]. FBGC formation can be prevented by competitive inhibitors of MMR activity (i.e., a-mannan) or inhibitors of glycoprotein processing that restrict MMR surface expression. Regarding the effect of lymphocytes on the foreign body reaction, studies have demonstrated that interactions with lymphocytes enhance adherent macrophage and FBGC production of proinflammatory cytokines. Interactions through indirect (paracrine) signaling showed a significant proinflammatory effect in enhancing adherent macrophage/FBGC at early time points, whereas interactions via direct (juxtacrine) mechanisms dominated at later points. Furthermore, lymphocytes prefer interactions with adherent macrophages and FBGCs, as opposed to biomaterial surfaces, resulting in lymphocyte activation [78e81]. In vivo studies using clinically synthetic biomaterials have demonstrated that there is a quantitative increase in T cells after secondary biomaterial exposure, but the T-cell subset distribution does not change, indicating nonspecific recruitment of T cells rather than an adaptive immune response. Studies in T celledeficient mice have shown no change in foreign body giant cell formation. In vitro studies with clinical synthetic biomaterials showed no expression of the activation markers CD69 and CD25 and lymphocyte proliferation was not identified [82e85]. Results from these in vivo and in vitro studies do not suggest an adaptive immune response with clinically relevant biomaterials, because Tcell markers CD25 and CD69 were not upregulated and T cell cytokines IL-2 and interferon-gamma were not present. FBGC formation at the surface of implanted biomaterials occurs quickly and within days to weeks after implantation. Fig. 39.3 shows scanning electron micrographs of monocytes, macrophages, and FBGCs at various times after implantation of a biocompatible polyurethane material in the in vivo rat cage system. Monocyte adhesion

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FIGURE 39.3 In vivo time-dependent processes of monocyte adhesion (A), macrophage development (B), early macrophage fusion to form foreign body giant cells (C), and late macrophage and foreign body giant cell fusion to form large foreign body giant cells (D).

(Fig. 39.3A) occurs quickly, at 2e4 days. Macrophage development is seen within 1 week after implantation (Fig. 39.3B). Macrophage fusion to form FBGCs can occur within 1e2 weeks after implantation (Fig. 39.3C); interestingly, FBGCs themselves can fuse to produce exceptionally large giant cells on the surface (Fig. 39.3D). Of interest in the use of porous scaffolds and other forms of topographical variations of implanted tissue-engineered systems, these large giant cells may be significantly large enough to cover the porous structure and thus inhibit cellular infiltration into the respective pores as well as inhibit diffusion from these porous structures. The formation of large FBGCs formed late (i.e., weeks) may compromise the function of tissue-engineered medical devices in which porosity and diffusion are important design criteria for the function of the device. Examples of such behavior are the development of biosensors and of complex tissue-engineered scaffolds in which proteins and other important modifying agents depend on diffusion from the device into the microenvironment. The potential for these events should be considered in the design of in vivo biocompatibility studies.

FIBROSIS AND FIBROUS ENCAPSULATION The end-stage healing response to biomaterials is generally fibrosis or fibrous encapsulation. However, tissueengineered devices may be exceptions to this general statement (e.g., porous materials inoculated with parenchymal cells or porous materials implanted into bone). Repair of implant sites involves two distinct processes: regeneration, which is the replacement of injury tissue by parenchymal cells of the same type, or replacement by connective tissue that constitutes the fibrous capsule. These processes are generally controlled by either the proliferative capacity of the cells in the tissue receiving the implant and the extent of injury as it relates to the destruction or persistence of the tissue framework of the implant site. The regenerative capacity of cells permits classification into three groups: labile, stable (or expanding), and permanent (or static) cells. Labile cells continue to proliferate throughout life, stable cells retain this capacity but do not normally replicate, and permanent cells cannot reproduce themselves after birth. Perfect repair with restitution of normal structure theoretically occurs only in tissue consisting of stable and labile cells, whereas all injuries to tissues composed of permanent cells may result in fibrosis and fibrous capsule formation with little restitution of the normal tissue or organ structure. Tissues composed of permanent cells (e.g., nerve cells, skeletal muscle cells, and cardiac muscle cells) most commonly undergo an organization of the inflammatory exudates, leading to fibrosis. Tissues composed of stable cells (e.g., parenchymal cells of the liver, kidney, and pancreas), mesenchymal cells

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39. BIOCOMPATIBILITY AND BIORESPONSE TO BIOMATERIALS

(e.g., fibroblasts, smooth muscle cells, osteoblasts, and chondroblasts), and vascular endothelial and labile cells (e.g., epithelial cells and lymphoid and hematopoietic cells) may also follow this pathway to fibrosis or may undergo resolution of the inflammatory exudates, leading to restitution of the normal tissue structure. The condition of the underlying framework or supporting stroma of the parenchymal cells after an injury has an important role in restoring normal tissue structure. Retention of the framework may lead to restitution of the normal tissue structure, whereas destruction of the framework most commonly leads to fibrosis. It is important to consider the species-dependent nature of the regenerative capacity of cells. For example, cells from the same organ or tissue but from different species may exhibit different regenerative capacities and/or connective tissue repair. The extent of provisional matrix formation is an important factor because it is related to wound healing by first or second intention. First intention (primary union) wound healing occurs when there is minimal to no space between the tissue and device, whereas second intention (secondary union) wound healing occurs when a large space is present, providing for extensive provisional matrix formation. Obviously, inappropriate or inadequate preparation of the implant site leading to extensive provisional matrix formation may predispose the implant to failure through mechanisms related to fibrous capsule formation. The inflammatory (innate) and immune (adaptive) responses have common components. It is possible to have inflammatory responses only with no adaptive immune response. In this situation, both humoral and cellular components that are shared by both types of responses may only participate in the inflammatory response. Table 39.4 indicates the common components of the inflammatory (innate) and immune (adaptive) responses. Macrophages and dendritic cells are known as professional antigen-presenting cells responsible for the initiation of the adaptive immune response. Many investigators have considered macrophages and dendritic cells to be distinctly different antigen-presenting cells (APCs). Hume summarized evidence that dendritic cells are part of the mononuclear phagocyte system and are derived from the same common macrophage precursor; they are responsive to the same growth factors, express the same surface markers, and have no unique adaptation for antigen presentation that is not shared by other macrophages [86].

IMMUNOTOXICITY (ACQUIRED IMMUNITY) The acquired or adaptive immune system acts to protect the host from foreign agents or materials and usually is initiated through specific recognition mechanisms and the ability of humoral and cellular components to recognize the foreign agent or material as being “nonself” [18,87e92]. Generally, the adaptive immune system may be considered to have two components: humoral or cellular. Humoral components include antibodies, complement components, cytokines, chemokines, growth factors, and other soluble mediators. These components are synthesized by cells of the immune response; in turn, they function to regulate the activity of these cells and provide for communication between different cells in the cellular component of the adaptive immune response. Cells of the immune system arise from stem cells in the bone marrow (B lymphocytes) or the thymus (T lymphocytes) and differ from each other in morphology, function, and the expression of cell-surface antigens. They share the common features of maintaining cell-surface receptors that assist in recognizing and/or eliminating foreign materials. Regarding tissue-engineered devices, the adaptive immune response may recognize the biological components, modifications TABLE 39.4

Common Components of Inflammatory (Innate) and Immune (Adaptive) Responses

Components Complement cascade components Immunoglobulins Cellular components Macrophages Natural killer cells Dendritic cells Cells with dual phagocytic and antigen-presenting capabilities

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of the biological components, or degradation products of the biological components, commonly known as antigens, and initiate immune response through humoral or cellular mechanisms. Components of the humoral immune system have important roles in the inflammatory responses to foreign materials. Antibodies and complement components C3b and C3bi adhere to foreign materials, act as opsonins, and facilitate phagocytosis of the foreign materials by neutrophils and macrophages that have cell-surface receptors for C3b. Complement component C5a is a chemotactic agent for neutrophils, monocytes, and other inflammatory cells and facilitates the immigration of these cells to the implant site. The complement system is composed of classic and alternative pathways that eventuate in a common pathway to produce the membrane attack complex, which is capable of lysing microbial agents. The complement system (i.e., complement cascade) is closely controlled by protein inhibitors in the host cell membrane that may prevent damage to host cells. This inhibitory mechanism may not function when non-host cells are used in tissue-engineered devices. T (thymus-derived) lymphocytes are significant cells in the cell-mediated adaptive immune response and their cell-adhesion molecules have a significant role in lymphocyte migration, activation, and effector function. The specific interaction of cell membrane adhesion molecules, sometimes also called ligands or antigens, with APCs produces specific types of lymphocytes with specific functions. Table 39.5 indicates cell types and function in the adaptive immune response. Obviously, the functions of these cells are more numerous than those indicated in the table, but the major function of these cells is provided to indicate similarities and differences in the interaction and responsiveness of these cells. Effector T cells (Table 39.6) are produced when their antigen-specific receptors and either the CD4 or the CD8 co-receptors bind to peptideemajor histocompatibility complexes (MHCs). A second, costimulatory signal is also required, which is provided by the interaction of the CD28 receptor on the T cell and the B7.1 and B7.2 glycoproteins of the Ig superfamily present on APCs. B lymphocytes bind soluble antigens through their cell-surface Ig and thus can function as professional APCs by internalizing the soluble antigens and presenting peptide fragments of these antigens as MHCepeptide complexes. Once activated, T cells can synthesize the T-cell growth factor IL-2 and its receptor. Thus, activated T cells secrete and respond to IL-2 to promote T cell growth in an autocrine fashion.

TABLE 39.5

Cell Types and Function in Adaptive Immune System

Cell Type

Motor Function

Macrophages (antigen-presenting cell [APC])

Process and present antigen to immunocompetent T cells Phagocytosis Activated by cytokines (i.e., interferon-gamma) from other immune cells

T cells

Interact with APCs and are activated through two required cell membrane interactions Facilitate target cell apoptosis Participate in transplant rejection (type IV hypersensitivity)

B cells

Form plasma cells that secrete immunoglobulins (IgGs) (IgA and IgE) Participate in antigeneantibody complex mediated tissue damage (type III hypersensitivity)

Dendritic cells (APC)

Process and present antigen to immunocompetent T cells Use Fc receptors for IgG to trap antigeneantibody complexes

Natural killer cells (non-T, non-B lymphocytes)

Innate ability to lyse tumor, virus-infected, and other cells without previous sensitization Mediates T- and B-cell function by secretion of interferongamma

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TABLE 39.6

Effector T Lymphocytes in Adaptive Immunity

T-Helper 1 Cells

CD41 Proinflammatory Activation of macrophages Produces IL-2, IFN-gamma, IL-3, TNF-a, GM-CSF, macrophage chemotactic factor, and migration inhibitor factor Induces IgG2a

Th2 helper cells

CD41 Antiinflammatory Activation of B cells to make antibodies Produces IL-4, IL-5, IL-6, IL-10, IL-3, GM-CSF, and IL-13 Induces IgG1

Cytotoxic T cells

CD81 Induces apoptosis of target cells Produces IFN-gamma, TNF-b, and TNF-a Releases cytotoxic proteins

GM-CSF, granulocyte macrophageecolony-stimulating factor; IFN, interferon; Ig, immunoglobulin; IL, interleukin; Th, T helper; TNF, tumor necrosis factor.

Cytokines are the messenger molecules of the immune system. Most cytokines have a wide spectrum of effects, reacting with many different cell types, and some are produced by several different cell types. Table 39.7 presents common categories of cytokines and lists some of their general properties. Whereas cytokines can be subdivided into functional groups, many cytokines such as IL-1, TNF-a, and IFN-gamma are pleotropic in their effects and regulate, mediate, and activate numerous responses by various cells. Immunotoxicity is any adverse effect on the function or structure of the immune system or other systems as a result of an immune system dysfunction [93]. Adverse or immunotoxic effects occur when humoral or cellular immunity needed by the host to defend itself against infections or neoplastic disease (immunosuppression) or unnecessary tissue damage (chronic inflammation, hypersensitivity, or autoimmunity) is compromised. Potential immunological effects and responses that may be associated with one or more of these effects are presented in Table 39.8. Hypersensitivity responses are classified on the basis of the immunological mechanism that mediates the response. There are four types: I (anaphylactic), II (cytotoxic), III (immune complex), and IV (cell-mediated delayed hypersensitivity). Hypersensitivity is considered to be increased reactivity to an antigen to which a human or animal has been previously exposed, with an adverse rather than a protective effect. Hypersensitivity is a synonym for allergy. Type I (anaphylactic) reactions and type IV (cell-mediated delayed hypersensitivity) reactions are the most common [94]. Types II and III reactions are relatively rare and are less likely to occur with medical devices and biomaterials; however, with tissue-engineered devices containing potential antigens (i.e., proteins), extracellular matrix (ECM) components, and/or cells, types II and III reactions must be considered in biological response evaluations. Type I (anaphylactic) hypersensitivity reactions are mediated by IgE antibodies, which are cytotropic and affect the immediate release of basoactive amines and other mediators from basophils and mast cells followed by the recruitment of other inflammatory cells. Type IV cell-mediated (delayed) hypersensitivity responses involve sensitized T lymphocytes that release cytokines and other mediators that lead to cellular and tissue injury. Type IV hypersensitivity (cell-mediated) reactions are initiated by specifically sensitized T lymphocytes. This reaction includes the classic delayed-type hypersensitivity reaction initiated by CD4þ T cells and direct cell cytotoxicity mediated by CD8þ T cells. The less common type II (cytotoxic) hypersensitivity involves the formation and binding of IgG and/or IgM to antigens on target cell surfaces that facilitate phagocytosis of the target cell or lysis of the target cell by activated complement components. Type II hypersensitivity (cytotoxic) is mediated by antibodies directed toward antigens present on the surface of cells or other tissue components. Three different antibody-dependent mechanisms may be involved in this type of reaction: complement-dependent reactions, antibody-dependent

IMMUNOTOXICITY (ACQUIRED IMMUNITY)

TABLE 39.7

687

Selected Cytokines and Their Effects

Cytokine

Effect

IL-1, TNF-a, INF-gamma, and IL-6

Mediate natural immunity

IL-1, TNF-a, and IL-6

Initiate nonspecific inflammatory responses

IL-2, IL-4, IL-5, IL-12, IL-15, and TGF-b

Regulate lymphocyte growth, activation, and differentiation

IL-2 and IL-4

Promote lymphocyte growth and differentiation

IL-10 and TGF-b

Downregulate immune responses

IL-1, INF-gamma, TNF-a, and migration inhibitor factor

Activate inflammatory cells

IL-8

Produced by activated macrophages and endothelial cells Chemoattractant for neutrophils

Monocyte chemoattractant protein-1, macrophage inflammatory protein-a, and Regulated on Activation, Normal T Expressed and Secreted

Chemoattractant for monocytes and lymphocytes

GM-CSF and G-CSF

Stimulate hematopoiesis

IL-4 and IL-13

Promote macrophage fusion and foreign body giant cell formation

G-CSF, granulocyteecolony-stimulating factor; GM-CSF, granulocyte macrophageecolony-stimulating factor; IFN, interferon; IL, interleukin; TGF, transforming growth factor; TNF, tumor necrosis factor.

TABLE 39.8

Potential Immunological Effects and Responses

Effects

Responses

Hypersensitivity

Histopathological changes

Type I: anaphylactic

Humoral responses

Type II: cytotoxic

Host resistance

Type III: immune complex

Clinical symptoms

Type IV: cell-mediated (delayed)

Cellular responses

Chronic inflammation

T cells

Immunosuppression

Natural killer cells

Immunostimulation

Macrophages

Autoimmunity

Granulocytes

reactions, cell-mediated cytotoxicity, or antibody-mediated cellular dysfunction. Type III immune complex hypersensitivity is present when circulating antigeneantibody complexes activate complement whose components are chemotactic for neutrophils that release enzymes and other toxic moieties and mediators leading to cellular and tissue injury. Immunological reactions that occur with organ transplant rejection also offer insight into potential immune responses to tissue-engineered devices. Mechanisms involved in organ transplant rejection include T cellemediated reactions by direct and indirect pathways and antibody-mediated reactions. Immune responses may be avoided or diminished by using autologous or isogeneic cells in cellepolymer scaffold constructs. The use of allogeneic or xenogeneic cells incorporated into the device requires the prevention of immune rejection by immune suppression of the host, induction of tolerance in the host, or immunomodulation of the tissue-engineered construct. The development of tissue-engineered constructs by immunoisolation using polymer membranes and the use of non-host cells have been compromised by immune responses. In this concept, a polymer membrane is used to encapsulate non-host cells or tissues, thus separating them from the host immune system. However, antigens shed by encapsulated cells were released from the device and initiated immune responses [42,95,96].

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39. BIOCOMPATIBILITY AND BIORESPONSE TO BIOMATERIALS

Although exceptionally minimal and superficial in its presentation, the previously discussed humoral and cell-mediated immune responses demonstrate the possibility that any known tissue-engineered construct may undergo immunological tissue injury. To date, our understanding of immune mechanisms and their interactions with tissue-engineered constructs is markedly limited. An obvious problem is that preliminary studies are generally carried out with nonhuman tissues, and immune reactions result when tissue-engineered constructs from one species are used to test the device in another species. Ideally, tissue-engineered constructs would be prepared from cells and tissues of a given species and subsequently tested in that species. Although this approach does not guarantee that immune responses will not be present, the probability of immune responses in this type of situation is markedly decreased. The following examples provide perspective to these issues. They further demonstrate the detailed and in-depth approach that must be taken to evaluate tissue-engineered constructs or devices and their potential adverse responses appropriately and adequately. The inflammatory response considered to be immunotoxic is persistent chronic inflammation. With biomaterials, controlled release systems, and tissue-engineered devices, potential antigens capable of stimulating the immune response may be present and these agents may facilitate a chronic inflammatory response that is of extended duration (weeks or months). Regarding immunotoxicity, this persistent chronic inflammation is of concern because immune granuloma formation and other serious immunological reactions such as autoimmune disease may occur. Thus, in biological response evaluation, it is important to discriminate between the short-lived chronic inflammation that is a component of the normal inflammatory and healing responses versus long-term, persistent chronic inflammation that may indicate an adverse immunological response. Immunosuppression may occur when antibody and T-cell responses (adaptive immune response) are inhibited. Potentially significant consequences of this type of response are frequent, and serious infections result from reduced host defense. Edelman and colleagues demonstrated that incorporating endothelial cells into three-dimensional collage matrices has a downregulating effect on the humeral and cellular immune response elicited by the endothelial cells [97]. The strong MHC-dominant immune response that occurs when endothelial cells are the primary component of an implant can be significantly reduced when the endothelial cells are embedded in the threedimensional collagen matrix. Although they retain many of the characteristics of quiescent endothelial cells, the endothelial cells evoke no significant humeral or cellular immune response in immunocompetent animals and also reduce the memory response to previous free endothelial cell implants. These studies are significant and demonstrate the influence of spatial matrix formation as well as matrix composition on endothelial cell immunophenotype. Thus, modulation of the matrix structure may be helpful in designing vascular conduits for tissueengineered devices. Using ECM scaffolds prepared under different conditions, Badylak and colleagues determined the participation of different macrophage phenotypes in the degradation and remodeling of the ECM scaffolds, demonstrating that the properties of the matrix can control the innate and possibly acquired immune responses to ECM scaffolds [98e100]. Immunostimulation may occur when unintended or inappropriate antigen-specific or nonspecific activation of the immune system is present. From a biomaterial and controlled-release system perspective, antibody and/or cellular immune responses to a foreign protein may lead to unintended immunogenicity. Enhancement of the immune response to an antigen by a biomaterial with which it is mixed ex vivo or in situ may lead to adjuvancy, which is a form of immunostimulation. This effect must be considered when biodegradable controlled-release systems are designed and developed for use as vaccines [101e103]. Patients implanted with polyurethanes used for left ventricular assist devices (LVADs) experience B-cell hyperreactivity and allosensitization [104]. There is evidence that T lymphocytes can be activated in response to biomaterials. T lymphocytes cultured in the presence of polyurethane particles from the flexible diaphragms of LVADs resulted in intracellular calcium flux, CD40 ligand expression, and nuclear translocation of nuclear factor (NFAT) of activated T cells. NFAT translocation was reduced by a calcineurin inhibitor and CD40 ligand expression was reduced by both a calcineurin inhibitor and CD25 blockade indicating IL-2edependent activation pathways [104,105]. In response to polyurethane particles, T lymphocytes exhibited classic activation indicators: calcium flux, translocation of transcription factors, upregulation of activation cell surface markers, and proliferation. Differences in human leukocyte antigen (HLA) gene inheritance can result in MHC diversity. MHC loci are among the most genetically variable loci in humans. The MHC class II proteins (DP, DQ, and DR) are found on APCs. Diversity in MHCII proteins results in individual variability in antigen presentation and thus immune responses. Because of this diversity, individuals mount immune responses to different epitopes of pathogens. LVAD recipients who are predisposed to develop B-cell hyperreactivity have HLA-DR3 expression, which indicates

IMMUNOTOXICITY (ACQUIRED IMMUNITY)

689

that lymphocyte responses to biomaterials are variable and depend on the individual’s genetic profile [106]. It is possible that only individuals with certain MHCII receptors can interact with biomaterials in a mechanism that results in a lymphocyte response. Autoimmunity is the immune response to the body’s own constituents, which are considered in this response to be autoantigens. An autoimmune response, indicated by the presence of autoantibodies or T lymphocytes that are reactive with host tissue or cellular antigens, may (or may not) result in autoimmune disease with chronic, debilitating, and sometimes life-threatening tissue and organ injury. Representative tests to evaluate for immune responses are given in Table 39.9. The table is not all-inclusive; other tests may be applicable. The examples presented represent only the large number of tests that are currently available [87,88,92]. The table is informative but incomplete, because in the future, direct and indirect markers of immune response may be validated and their predictive value documented, thus providing new tests for immunotoxicity. Direct measures of immune system activity by functional assays are the most important types of tests for immunotoxicity. Functional assays are generally more important than are tests for soluble mediators, which are more important than phenotyping. Signs of illness may be important in in vivo experiments, but symptoms may also have a significant role in studies of immune function in clinical trials and postmarket studies. As with any type of test for biological response evaluation, immunotoxicity tests should be valid; they have been shown to provide accurate, reproducible results that indicate the effect being studied and are useful in a statistical analysis. This implies that appropriate control groups are also included in the study design. Immunogenicity involving a specific immune response to a biomaterial is an important consideration because it may lead to serious adverse effects. For example, a foreign, nonhuman protein may induce IgE antibodies that cause an anaphylactic (type I) hypersensitivity reaction. An example of this type of response is latex protein found in latex gloves. Lowemolecular weight compounds such as chemical accelerators used to manufacture latex gloves may also induce a T cellemediated (type IV) reaction, resulting in contact dermatitis. Tests for type I (e.g., antigen-specific IgE) and type IV (e.g., guinea pig) maximization hypersensitivity should be considered for materials with the potential to cause these allergic reactions. In addition to hypersensitivity reactions, a device may elicit autoimmune responses (i.e., antibodies or T cells) that react with the body’s own constituents. An autoimmune response may lead to the pathological consequences of an autoimmune disease. For example, a foreign protein may induce IgG or IgM

TABLE 39.9

Representative Tests for Evaluation of Immune Responses

Functional Assays

Soluble Mediators

Skin testing

Antibodies

Immunoassays (e.g., enzyme-linked immunosorbent assays)

Complement

Lymphocyte proliferation

Immune complexes

Plaque-forming cells

Cytokine patterns (T-cell subsets)

Local lymph node assay

Cytokines (IL-1, IL-1ra, TNF-a, IL-6, TGF-b, IL-4, and IL-13)

Mixed lymphocyte reaction

Chemokines

Tumor cytotoxicity

Basoactive amines

Antigen presentation

Signs of illness

Phagocytosis Degranulation

Allergy

Resistance to bacteria, viruses, and tumors

Skin rash

Phenotyping

Urticaria Edema

Cell-surface markers

Lymphadenopathy

Major histocompatibility complex markers IL, interleukin; TNF, tumor necrosis factor; TGF, transforming growth factor.

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39. BIOCOMPATIBILITY AND BIORESPONSE TO BIOMATERIALS

antibodies that cross-react with a human protein and cause tissue damage by activating the complement system. In a similar fashion, a biomaterial or controlled-release system that has a gel or oil constituent may act as an adjuvant, leading to the induction of an autoimmune response. Even if an autoimmune response (autoantibodies and/or autoreactive T lymphocytes) is suggested in preclinical testing, it is difficult to obtain convincing evidence that a biomaterial or controlled-release system causes autoimmune disease in animals. Therefore, routine testing for the induction of autoimmune disease in animal models is not recommended. Babensee and coworkers tested the hypothesis that by promoting an inflammatory response, the biomaterial component of a medical device can recruit APCs (e.g., macrophages and dendritic cells) and induce their activation, thus acting as an adjuvant in the immune response to foreign antigens originating from the histological component of the device [107e109]. Using polystyrene and polylactic-glycolic acid microparticles and polylactic-glycolic scaffolds together with their model antigen, ovalbumin, in a mouse model for 18 weeks, the researchers demonstrated that a persistent humoral immune response that was Th2, T-cell dependent, as determined by IgG1, was present. These findings indicated that activation of CD41 T cells and the proliferation and isotype switching of B cells had occurred. A Th1 immune response characterized by the presence of IgG2a was not identified. Moreover, the humoral immune responses for all three types of microparticles were similar, indicating that the production of antigen-specific antibodies was not material chemistryedependent in this model. Babensee et al. suggested that the presence of the biomaterial functions as an adjuvant for initiation and promotion of the immune response and augments the phagocytosis of the antigen with expression of MHCII and costimulatory molecules on APCs with the presentation of antigen to CD41 T cells. Babensee and coworkers identified differential levels of dendritic cell maturation on different biomaterials used in combination products [110e114]. The effect of biomaterials on dendritic cell maturation, and the associated adjuvant effect, are novel biocompatibility selection and design criteria for biomaterials to be used in combination products in which immune consequences are potential complications or outcomes. Badylak and colleagues carried out extensive studies on the use of xenogeneic ECM as a scaffold for tissue reconstruction [115e117]. Use of the small intestinal submucosa (SIS) ECM in animals indicated a restricted Th2-type immune response. The presence of natural antibodies to the terminal galactose-a1,3-galactose (a-gal) epitope is considered to be a major barrier to xenotransplantation in humans. Cell membranes of all animals except humans express this epitope, and naturally occurring antibodies mediate hyperacute or delayed rejection of transplanted organs through complement fixation or antibody dependence cell-mediated cytotoxicity. Whereas ECM derived from porcine tissues, SIS, contains small amounts of the galactose epitope, it appears that the quantity or distribution of this epitope and/or the subtype of Ig response to the epitope is such that complement activation does not occur [118]. In addition, the resorbable characteristics of this nonchemically cross-linked ECM scaffold demonstrate constructive tissue remodeling and deposition of new matrix whereas chemically cross-linked ECM leads to active inflammation and eventually scar formation. The role of Th1 and Th2 lymphocytes in cell-mediated immune responses to xenografts has been examined. Activation of the Th1 pathway leads to macrophage activation, stimulation of complement fixing antibody isotypes, and differentiation of CD8þ cells to a cytotoxic type phenotype that is associated with both allogeneic and xenogeneic transplant rejection. The Th2 lymphocyte response does not activate macrophages and leads to production of noncomplement fixing antibody isotypes; it is usually associated with transplant acceptance. The use of appropriate animal models is an important consideration in the safety evaluation of controlled-release systems that may contain potential immunoreactive materials [91,119,120]. A study involving the in vivo evaluation of recombinant human growth hormone in poly(lactic-co-glycolic acid) (PLGA) microspheres demonstrated the appropriate use of various animal models to evaluate biological responses and the potential for immunotoxicity. Using biodegradable PLGA microspheres containing recombinant human growth hormone (rhGH), Cleland et al. used rhesus monkeys, transgenic mice expression rhGH, and normal control (Balb/C) mice in their in vivo studies [112]. Rhesus monkeys were used for serum assays in the pharmacokinetic study of rhGH release as well as tissue responses to the injected microcapsule formulation. Placebo injection sites were also employed, and a comparison of the injection sites from rhGH PLGA microspheres and placebo PLGA microspheres demonstrated a normal inflammatory and wound healing response with a normal focal foreign body reaction. To examine the tissue response further, transgenic mice were used to assess the immunogenicity of the rhGH PLGA formulation. Transgenic mice expressing a heterologous protein were previously used to assess the immunogenicity of sequence or structural mutant proteins [121,122]. With the transgenic animals, no detectable antibody response to rhGH was found. In contrast, the Balb/C control mice had a rapid onset of high titer antibody response to the rhGH PLGA formulation. This study pointed out the appropriate use of animal models not only to evaluate biological responses but also for one type of immunotoxicity (immunogenicity) of controlled-release systems.

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The focus in tissue engineering traditionally has been on modulating the fate of transplanted host cell populations that participate directly in reconstructing tissues. However, new materials for tissue engineering are being considered that give greater control over the inflammatory and immune responses [123]. Biomimetic strategies based on viruses and bacteria are being used to develop immune evasive biomaterials [124]. Materials are being investigated that can promote tolerance to specific antigens and cells by directly signaling APCs such as dendritic cells, or by releasing growth factors or cytokines that promote tolerance. On the other hand, materials might promote a destructive immune response by directly providing immunity-promoting signals or releasing insoluble factors. This approach could be used to combat infections and cancer [125].

CONCLUSION Tissue-engineered devices are biologicalebiomaterial combinations in which some component of tissue has been combined with a biomaterial to create a device for the restoration or modification of tissue or organ function. The biocompatibility and bioresponse require the ultimate achievement of four significant goals if these devices are to function adequately and appropriately in the host environment. These goals are (1) restoration of the target tissue with its appropriate function and cellular phenotypic expression; (2) inhibition of the macrophage and FBGC foreign body response that may degrade or adversely modify device function; (3) inhibition of scar and fibrous capsule formation that may be deleterious to the function of the device; and (4) inhibition of immune responses that may inhibit the proposed function of the device and ultimately lead to the destruction of the tissue component of the tissue-engineered device. This chapter has presented a brief and limited overview of mechanisms and biological responses that determine biocompatibility: inflammation, wound healing, and immunotoxicity. Given the unique nature of the combination of tissue components and biomaterials in tissue-engineered devices, coupled with the species differences in biological responses, a significant future challenge in developing tissue-engineered devices is to construct and use a unique set of tests that will ensure that these four goals are achieved for the lifetime of the device in its in vivo environment in humans.

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Further Reading Cleland JL, Duenas E, Daugherty A, Marian M, Yang J, Wilson M, et al. Recombinant human growth hormone poly(lactic-co-glycolic acid) (PLGA) microspheres provide a long lasting effect. J Control Release 1997;49:193e205.

C H A P T E R

40 Hybrid Composite Biomaterials Nirmalya Tripathy1, Elumalai Perumal2, Rafiq Ahmad1, Jeong Eun Song1, Gilson Khang1 1

Chonbuk National University, Jeonju-si, Republic of Korea; 2The Catholic University of Korea, Seocho-gu, Republic of Korea

INTRODUCTION Damage and degeneration of tissues caused by disease, injury, and trauma to the human body necessitate treatments including transplanting tissue from one site to another in the same patient (an autograft) or from one individual to another (a transplant or allograft). All techniques have the serious constraints of being expensive and painful and having the risks for rejection by the patient’s immune system and the possibility of introducing infection or disease from the donor to the patient. Therefore, the approach of tissue engineering to repair defects in tissues is perceived as a smart strategies because the repair process proceeds with the patient’s own tissue. The field of tissue engineering is highly multidisciplinary; it has the aim of developing biological substitutes to regenerate damaged tissues and to restore, maintain, or improve tissue function [1,2]. The tissue engineering triad is mainly composed of three main key components of engineered tissues: cells, scaffolds, and the growth-stimulating signal. Typically, the three-dimensional (3D) scaffolds are made from natural or synthetic biomaterials in various formats, serving as a template to provide an appropriate environment for tissue regeneration. These tissue engineered scaffolds are mainly seeded with cells and growth factors or subjected to biophysical stimuli in a bioreactor (a device that applies different types of mechanical or chemical stimuli to cells) [3]. These cell-seeded scaffolds are then either cultured in vitro to produce tissues followed by implantation into an injured site or directly implanted into the injured target site using the body’s own systems in which tissues regeneration is induced in vivo [4]. In this chapter, we outline various materials used for scaffolding, the functions of scaffolds, approaches to scaffolding, and tissuespecific considerations for scaffolding, using bone as an example. Materials that are implanted to repair, replace, or augment existing tissues in the body are generally known as biomaterials. In the wider context covered in this chapter, biomaterials include all materials formed in biological systems, e.g., the specific products of biomineralization. The development of biomaterials, both as products and in understanding their in vivo behavior, has been driven largely by the desire to assist in caring for patients. The material-forming processes occurring in living organisms require much milder reaction conditions than are currently used in the laboratory, such that a new area of materials chemistry, “biomimetics,” has been established in which scientists are taking ideas from biology to generate “softer” routes to useful materials. Biomaterials present in nature provide the necessary structure and architectures of all animal and plant species on earth and function to maintain the structure of organs as well as the organism itself. In nature, materials that are used are polymers such as polysaccharides and proteins and a relatively small number of simple insoluble oxides and salts. These can be put together in a wide range of combinations to produce materials that are soft, hard, flexible, elastic, etc. In contrast, the range of available materials for biomedical applications is vast and includes metals, polymers, ceramics, and their composites. In designing medical devices, materials are chosen to suit their intended use and the proposed implantation area. The materials that are used must have properties that are compatible with the location in which they are placed. Properties such as the tensile strength, toughness, elasticity, and hardness have to be considered, and other factors such as material transparency may have to be thought about if, for instance, the device is to be used within the eye.

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Although technologists have a wider array of materials at their disposal, it is not a simple matter to come up with a material or series of materials that fulfil all of the criteria required for successful implantation or biomedical use. In the wider context covered in this chapter, biomaterials include all materials formed in biological systems, e.g., the specific products of biomineralization. The development of biomaterials, both as products and in understanding their in vivo behavior, has been driven largely by the desire to assist in caring for patients. The material-forming processes occurring in living organisms require much milder reaction conditions than are currently used in the laboratory, such that a new area of materials chemistry, “biomimetics,” has been established in which scientists are taking ideas from biology to generate “softer” routes to useful materials.

FUNDAMENTALS OF BONE DEVELOPMENT AND DEFECTS Bone is a calcium reservoir in the body in the form of hydroxyapatite (HA) (85%), calcium carbonate (10%), calcium fluoride (w3%), and magnesium fluoride (w3%). There are two main structural types in the bone: cortical bone forms a dense outer shell and cancellous or trabecular bone forms the porous inner core. Cortical bone is highly dense structure consisting of hierarchical structures, each within a different size scale; it provides torsion, bending resistance, and compressive strength. On the other hand, cancellous bone is highly porous and possesses an interconnected trabeculae network filled with marrow, a hierarchical structure ranging from solid material to trabeculae, lamellae, and a collagenemineral composite. The trabeculae have a large surface area for the diffusion of nutrients and the circulation of growth factors, which allows cancellous bone tissue to be metabolically active and which can be remodeled more frequently than that of cortical bone. Bone formation occurs via two distinctive pathways: intramembraneous and endochondral. First, mesenchymal cellular condensation occurs and acts as a template for subsequent bone formation. Intramembraneous bone formation involves the differentiation of mesenchymal progenitor cells directly into osteoblasts; it further develops into parts of the mandible, clavicle, and many cranial bones. Most of the bones in the human body, including all of the long bones and vertebrae, were formed through endochondral bone formation, in which mesenchymal progenitor cells differentiate into chondrocytes, which are responsible for cartilaginous template formation, and then are further mineralized and reinstated by bone. Upon fracture, the bone becomes repaired by recapitulating several events of endochondral and intramembraneous bone formation and heals with no scar tissue formation. Usually, the formation of hematoma involves an inflammatory response and many signaling molecules specific for the regulation of new bone formation (i.e., tumor necrosis factor-a, interleukins, fibroblast growth factors, platelet-derived growth factor, vascular endothelial growth factor [VEGF], bone morphogenetic proteins [BMPs], etc.). The formation of intramembranous bone occurs immediately at the cortex and periosteum. Then the fracture becomes stabilized by the external soft tissues via callus formation; subsequently, chondrocyte proliferation takes place. Then the ingrowing blood vessels carry chondroclasts, which reabsorb calcified cartilage and osteoblastic progenitors and initiate new bone formation. The mechanical continuity of the cortex is achieved via subsequent remodeling of the newly formed bone. In case of damage or disease requiring bone tissue regeneration, the formation of hematoma and an early inflammatory response occurs, which facilitates host cell recruitment and the release of critical signaling molecules. Thus, smart scaffolds mimicking the properties of normal bone tissue development are important to efficient bone tissue engineering. Fundamental developmental biology principles that should be considered for bone tissue engineering are: (1) the use of pluripotent or multipotent stem cells; (2) the establishment of required growth factors, involved genes, and signal transduction cascades facilitating bone formation; (3) an understanding of the bone formation process and normal tissue healing; (4) an understanding of complex interactions between mesenchyme and epithelium, and mesenchyme-encoding, tissue-specific patterns; (5) an understanding of the criticality of the tissue microenvironment’s physical characteristics (mechanotherapy, etc.); and (6) neovascularization and angiogenesis of the newly formed and developed bone tissue.

FUNCTIONS OF SCAFFOLDING AND EXTRACELLULAR MATRIX Various tissue engineered constructs or scaffolds have been extensively fabricated from numerous materials and a plethora of manufacturing approaches for the neoregeneration of different tissues and organs in the body. When designing an ideal scaffold for tissue engineering, a number of important features and functions must be considered regardless of the type of tissue:

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Architecture: Scaffolds should possess an interconnected pore structure and high porosity and void volume for proper vascularization, the formation of neotissue, and remodeling, which facilitates cell penetration and host tissue integration upon implantation. The biomaterial must provide enough interconnected porous structure for efficient nutrient and metabolite diffusion to cells within the scaffolds and to the extracellular matrix (ECM) formed by these cells without significantly compromising mechanical scaffold stability. Moreover, the scaffold should have enough mean pore size for cell migration within the structure where they eventually bind to the ligands present in scaffold materials, but small enough to establish a sufficiently high specific surface leading to a minimal ligand density to allow efficient binding of a critical number of cells to the scaffold [5]. Thus, for any scaffold, a critical range of pore sizes exists that may vary depending on the cell type used and the tissue being engineered [6]. Furthermore, the materials must be degradable upon implantation at a rate corresponding to the newly formed matrix by the developing tissue. The issue of core degradation, which arises from the lack of vascularization and waste removal from scaffolds, is one of the major concerns in the tissue engineering field [7,8]. Cytocompatibility: An important criterion for any tissue engineered scaffold is that it must be biocompatible: seeded cells must adhere, function normally, migrate onto or through the scaffold surface, and start proliferating during both in vitro culture and in vivo implantation before laying down new matrix. After implantation, the scaffold must be compatible with the cellular components and endogenous cells in host tissue and thus must elicit a negligible immune reaction or inflammatory response to avoid reduced healing or rejection by the body. Bioactivity: Scaffolds must interact actively with the cellular components of the engineered tissues to regulate their activities. The construct material must be composed of biological cues such as cell-adhesive ligands to enhance attachment or physical cues such as topography to influence cell morphology and alignment. It can also act as a delivery vector for exogenous growth-stimulating signals to speed up regeneration such as growth factors. For example, hydrogels synthesized by covalent or ionic cross-linking can encapsulate proteins and perform a responsive release by hydrogels swelling [9]. Biodegradability: Scaffolds implanted for tissue engineering purposes were intended to be eventually replaced by the body’s own cells. Thus, the scaffold material must be biodegradable, facilitating cells to generate their own ECM and produce nontoxic by-products followed by their removal from the body with no interference from other organs. Moreover, an inflammatory response with the infusion of controlled cells such as macrophages is essential for the execution of materials degradation with the formation of tissues or cells. Mechanical characteristics: From a practical point of view, a scaffold should be strong enough to allow surgical handling during implantation. Moreover, its mechanical characteristics must be similar to the implantation anatomical site, providing mechanical and shape firmness to the tissue defect. Especially for orthopedic and cardiovascular applications, it is critical to design a scaffold with sufficient mechanical integrity mimicking host bone or cardiac tissue integrity. This becomes more challenging with age-dependent defects or fractures; for example, generally, fractures heal to an acceptable weight-bearing point at about 6 weeks in young individuals, but the repair rate slows down in elderly people. With the evolution of tissue engineering and regenerative medicine, the focus has been mostly diverted to manufacturing scaffolds with good mechanical properties and reduced porosity, which have been shown to have potential in vitro; however, these were unsuccessful when implanted in vivo owing to their insufficient capacity for vascularization. Thus, a well-established balance between the scaffold mechanical features and porous architecture allowing cell infiltration and vascularization is a vital factor for the success of any tissue engineered scaffold [4]. Except for blood cells, normal cells in human tissues are anchorage-dependent and located in a solid matrix called the ECM. There are various types of ECM in human tissues consisting of multiple components and a tissue-specific composition. Generally, functions of the ECM in tissues can be divided into five categories: (1) it provides a physical environment and structural support to cells in specific tissues to attach, grow, migrate, and respond to signals; (2) it provides mechanical and structural strength to tissues including elasticity and rigidity; (3) it actively provides bioactive cues to native cells to regulate their activities; (4) it behaves as a reservoir for growth factors and potentiates their bioactivities; and (5) it provides a degradable physical platform, allowing vascularization and remodeling in response to developmental, physiological, and pathological challenges during tissue dynamic processes such as morphogenesis, homeostasis, and wound healing, respectively. Thus, it is important for any ideal scaffold material to mimic the dynamic nature of the ECM in native tissues, at least partially [10].

SCAFFOLDING APPROACHES IN BONE TISSUE ENGINEERING In this section, we discuss various approaches employed in bone tissue engineering from a materials point of view; further details are elaborated on in later sections.

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Osteoinductive Materials Osteoinductive refers to materials having the ability to induce the formation of ectopic bone by summoning the surrounding environment to form neobone [11e13]; thus, it holds immense potential for bone tissue regeneration. Materials including natural and synthetic ceramics (for example, HA, alumina ceramic, and porous Bioglass) [14] and several calcium phosphate compositions and their composites (for example, HAepoly[lactic-co-glycolic acid] [PLGA]) have been reported to be efficient osteoinductive materials in various bone tissue engineering studies. In particular, calcium phosphateebased designs (i.e., cements [15,16], coatings [17e19], sintered ceramics, and coralderived ceramic [20,21]) have illustrated osteoinduction in several animal models. Polymer-ceramic blends (i.e., HAePLGA) have displayed osteoinductive properties and induced ectopic bone formation [22,23]. Other than the material’s chemical composition, several other factors have a critical role in determining osteoinduction, including porosity, surface properties, and nanotopography and microtopography. To explain osteoinductivity, two main hypotheses were proposed. The first is based on biomaterial surface properties that absorb and present osteoinductive factors to the surrounding cells. The second hypothesis is that calcium phosphateebased materials release calcium and phosphate ions, which influence stem cell differentiation into bone cells [13].

Immunomodulatory Materials These systems have the ability to manipulate or modulate the immune system in favorably so as to enhance tissue regeneration. In general, the host’s immune system reacts to an external implant by initiating an acute reaction to the surgical injury and recognize the alien material innately; this followed by adaptive immunity-mediated chronic inflammation in response to the recognition of specific antigens. Therefore, myriads of approaches were proposed in immune-bioengineering, highlighting the relevance of rational control over host inflammation that suggest cell-specific responses, accelerated wound healing, and bone tissue regeneration (enhanced activity of boneforming cells and decreased activity of natural killer (NK) cells and T and B cellemediated immunity). This includes the appropriate selection of materials, material surface functionalization and modulating agents (i.e., surface treatments and topography) and the inclusion of artificial ECM and bioactive agents. One approach is to develop biomaterials mimicking ECM composition and structure. Smith et al. showed polydioxanone and collageneelastin composites with immune-modulating features by decreasing NK cell activity and T- and B-cell proliferation [24]. Material microscale and nanoscale architecture and its surface topography have a key role in modulating the immune system and host acute immune response upon scaffold implantation. Thus, the biomaterial surface should be functionalized to limit macrophage adhesion and activation, shielding them from protein absorption via polymer coating, delivering bioactive molecules (i.e., growth factors, antiinflammatory drugs), and thereby facilitating their fusion into foreign body giant cells. Some studies evaluated the pharmacologic modification effects of inflammatory response upon in vivo bone regeneration including cytokine-specific agents, corticosteroids, prostaglandins, nonsteroidal antiinflammatory drugs, and selective prostaglandin agonists [25]. A few studies have shown the enhanced strength of healed bone, vascularization at the fracture site, or accelerated fracture repair and regeneration with the controlled release of TP508 from various biodegradable scaffolds (PLGA microspheres and poly[propylene fumarate]) [26,27]. Although these results are interesting and promising, further detailed and advanced studies are essential that integrate inflammatory modulation strategies into tissue engineering to enhance tissue regeneration.

Hybrid Materials Any biomaterial chosen for tissue engineering scaffolding needs to possess certain biological and physiochemical characteristics specific for the target tissue. However, it is difficult for any material to meet all of these expectations by itself; thus, hybrid materials have been made by combining two or more materials complementing each other in terms of the required features. Various optimal and advanced hybrids were developed (i.e., copolymers, polymere polymer blends, polymereceramic composites) These are described in detail in later sections.

Hydrogels Hydrogels, which have the ability to mimic ECM topography and deliver bioactive substances to promote tissue regeneration owing to their physical characteristics and biocompatibility, have become popular in tissue engineering and regenerative medicine. Examples includes both nature-derived (collagen, silk, and gelatin) and synthetic biomaterials (poly[vinyl alcohol] and poly[ethylene glycol]). Self-assembling peptides have attracted the interest of

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researchers for scaffolding applications because these systems are nonimmunogenic, biocompatible, and biodegradable, and can be readily manufactured chemically and biologically to serve as a starting scaffolding matrix [28,29]. For instance, self-assembling RAD16-I (i.e., PuraMatrix, Cambridge, MA) can form an injectable nanofiber network or hydrogel upon implantation [30e32].

SCAFFOLDING MATERIALS Natural Polymers The major goal of bone tissue engineering is to develop bioconstructs that substitute for the functionality of damaged natural bone structures as much as possible if critical-sized defects occur. Scaffolds that mimic the structure and composition of bone tissue and cells have a pivotal role in bone tissue engineering. Natural polymers have attractive properties for the construction of 3D scaffolds, such as biocompatibility and biodegradability. The porosity, charge, and mechanical strength can be controlled by changing polymer concentrations or polymerization conditions, or by introducing various functional groups. Bioactivity can also be tuned by adding chemicals, proteins, peptides, and cells. The most commonly studied natural polymers for the purpose of bone engineering are collagen/gelatin, chitosan (CHI), silk, alginate, hyaluronic acid (HAc), and peptides [33]. Here we discuss studies in which these natural polymers were studied as 3D scaffolds for bone regeneration and were modified in different manners to improve their osteogenic capabilities. Silk Silk fibroin (SF) is a fibrous protein that is produced mainly by silkworms and spiders. Its unique mechanical properties and fine-tunable biodegradation rate, and the ability to support the differentiation of mesenchymal stem cells along the osteogenic lineage have made SF a favorable scaffold material for bone tissue engineering. SF can be processed into various scaffold forms that can be combined synergistically with other biomaterials to form composites and chemically modified, which provides an impressive toolbox and allows SF scaffolds to be tailored to specific applications [34]. Silk is composed of two major proteins: SF (fibrous protein) and sericin (globular protein). SF is a protein isolated from different animals in the form of an aqueous protein solution. The ability to produce silk has evolved multiple times among insects such as Bombyx mori, spiders, mites, and beetles with diverse functions [35]. In one study, calcium phosphate (CaP)/silk powders were incorporated into silk scaffolds to improve the porous structure and distribution of CaP powders in the composite scaffolds [36,37]. The scaffolds tested contained pure silk, or silk with 5% or 10% CaP. The pure silk and silk composite scaffolds were prepared using a freeze-drying method. The addition of CaP did not affect the compressive strength of the material (all tested were w70 KPa) or the compressive modulus (w250 MPa for all materials tested). All scaffolds tested supported bone marrow stromal cell (BMSC) proliferation to comparable levels. However, alkaline phosphatase (ALP) activity was significantly higher in CaPesilk scaffolds at days 7 and 14. The expression of the osteogenic markers ALP, type 1 collagen, and osteocalcin (OCN) was also significantly increased in CaPescaffolds at day 7 and 14 compared with pure silk scaffolds. In vivo bone formation was assessed in a calvarial defect model by microcomputed tomography (mCT). Groups studied were silk scaffold, CaPescaffold, silk scaffold plus mesenchymal stem cells (MSCs), and CaPescaffold plus MSCs. Defect areas were collected after 4 weeks. When bone formation was measured, it was clear that the addition of CaP into the scaffold significantly increased bone volume in the area. These results suggest that the presence of CaP in the scaffold was sufficient to enhance osteogenesis, because it did had no effect on the scaffold’s mechanical properties. Decellularized bovine trabecular bone was used for comparison. It was found that the values of these parameters approached those for bovine trabecular bone in scaffolds with 3.1% and 4.6% HA. Calcium content was studied in all scaffolds at up to 10 weeks; there was increase over time but no differences between groups. These studies demonstrate a system in which, although there was no improvement in other work, the relationship between scaffold degradability and osteogenesis was examined [38]. 3D porous SF scaffolds were prepared with two different degradation rates. The water-based scaffold was synthesized as the rapidly degrading scaffold (control) and a slower-degrading scaffold was obtained by adding hexafluoroisopropanol (HFIP) to induce insolubility in aqueous media (HFIP scaffold). Degradation studies were performed; the control scaffold lost more than 90% of its mass at day 7 and the HFIP scaffold’s mass remained nearly the same at 7 days. Various biochemical assays were performed on both scaffolds after seeding with MSCs and culturing for 16 or 56 days. At 56 days, the DNA content in the control scaffolds was around sixfold lower than for HFIP scaffolds.

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There was no significant difference in ALP activity between the two scaffolds. However, calcium content/DNA and total collagen/DNA showed a dramatic increase in the rapidly degrading scaffold compared with the slowly degrading scaffold. There was no significant difference in the expression of osteogenesis-related genes ALP, bone sialoprotein (BSP), Col1a1, or osteopontin (OPN) between control and HFIP scaffolds. However, at day 56, hypoxia-inducible factor-1 (HIF-1) expression was significantly higher in the control than the HFIP scaffold. Studies in transgenic mice have shown that HIF-1 is a critical component of bone regeneration. Another study was conducted to evaluate the ability of fibroin scaffolds combined with human stem cells, such as human dental pulp stem cells (hDPSCs) and human amniotic fluid stem cells (hAFSCs), to repair critical-size cranial bone defects [39]. These scaffolds had 85% porosity and pore diameters ranging from 10 to 250 mM, and a compressive modulus of 25.69  1.61 kPa. hDPSCs and hAFSCs were seeded onto the scaffolds and grown in osteogenic media for 10 days before implantation into the cranial critical-size defect. Groups tested consisted of scaffolds without cells, scaffold plus hDPSCs, scaffold plus hAFSCs, and empty defect. Thirty days after surgery, radiograph images showed that scaffolds without cells were able to repair the area slightly, but this effect was more pronounced when either kind of stem cell was present. Hematoxylineeosin staining showed that vascularization was present in all scaffolds tested; however, hASFCs seemed to have a greater potential for bone regeneration because it was the only scaffold tested that showed bone in scaffold areas distant to the dura matter. Another study reported the effects of different concentrations of SF protein on a 3D scaffold pore microstructure and its effects on bone formation when cultured with BMSCs transfected with BMP-7 [36,37]. At 1 wt% silk protein, scaffolds had a porosity of 94% and a pore size ranging from 250 to 300 mM. At 2 wt% protein, porosity was 87% and the pore size was 200e250 mM. At 3.5 wt% silk protein, scaffolds had a porosity of 80% and a pore size ranging from 150 to 200 mM. At 2 wt% protein, porosity was 71% and pore size was from 80 to 150 mM. MSCs proliferated on all scaffolds, but at day 14 there was a significant decrease in the 5% scaffold. ALP activity was shown to increase in all groups, but there was significantly higher expression in the 3.5% scaffold. After 2 weeks in osteogenic medium, expression of osteogenic markers in transfected and untransfected MSCs in the various scaffolds was analyzed by reverse transcriptaseepolymerase chain reaction (RT-PCR). ALP, Col1, and OCN had similar levels of expression in all scaffolds tested when they were transfected with BMP-7, except in the 5% scaffold, where there was a significant decrease in marker expression. This study indicated that decreased scaffold porosity is detrimental to the promotion of bone formation and that the presence of BMP-7 greatly enhances cells’ ability to express osteogenic phenotypes. Collagen The ECM provides physical support to tissues by occupying the intercellular space, acting not only as benign native scaffolding for arranging cells within connective tissues but also as a dynamic, mobile, and flexible substance defining cellular behaviors and tissue function [40]. For most soft and hard connective tissues (bone, cartilage, tendon, cornea, blood vessels, and skin), collagen fibrils and their networks function as an ECM, the highly organized 3D architecture surrounding various cells. Collagen has a dominant role in maintaining the biologic structural integrity of ECM and is highly dynamic, undergoing constant remodeling for proper physiologic functions [40]. As a primary component of bone, collagen (and gelatin) is an idyllic candidate for the design of 3D scaffolds [41]. It is inherently biocompatible and biodegradable and stimulates the proliferation and differentiation of cells as an ECM. However, it has poor mechanical properties. In many studies, collagen was used as a base for 3D scaffolds and modified by adding polymers and other biomolecules to improve osteoinductivity. In one [42], the authors combined the mechanical properties of a 3D macrochanneled poly-ε-caprolactone (PCL) scaffold, fabricated by robotic dispensing, with the bioactive properties of collagen. An MSC-loaded collagen hydrogel collagen was inserted onto the macrochanneled scaffold. In addition, they studied the effect of growing these cell-seeded scaffolds in a perfusion bioreactor to test how osteogenesis was affected by the continuous supply of fresh media and shear stress. The collagenePCL scaffolds increased cell proliferation when grown in the perfusion chamber compared with cells grown under static conditions. The activity of ALP, an early osteogenic marker, and osteogenic genes OPN, OCN, and BSP were significantly upregulated at 14 days in cells grown in the perfusion chamber. Bone marrow MSCs (BM-MSC) have shown great potential for tissue engineering purposes as they are relatively easy to isolate, retain their multipotency even after several passages, and can be induced to differentiate into bone. However, there is a low amount that can be harvested from bone marrow and they have limited proliferation and high senescence [43]. It has been reported that MSCs can be isolated from the Wharton jelly found in the umbilical cord (UC-MSC), and they are more abundant, more easily expanded, and more resistant to cryogenic storage [44]. The adhesion, migration into scaffold, growth, spreading, osteogenic differentiation, ECM degradation, and synthesis of BM-MSCs versus UC-MSCs were compared. The tested collagen scaffold (CS) showed excellent cytocompatibility

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with both cell types and could maintain high proliferation and viability. When stimulated with osteogenic induction medium, both cell types showed comparable osteogenic gene expression, migration and scaffold colonization, and the ability to contract the CS. UC-MSCs were much more capable of producing an ECM. They demonstrated at least 10-fold higher expression of the ECM marker genes collagen I, collagen III, collagen IV, and laminin than did BMMSCs. Because the production of ECM is an integral component of bone regeneration, UC-MSCs could have a significant impact on bone tissue engineering; its osteogenic potential should be investigated in additional scaffold compositions. In another study, the hydraulic permeability (k) of a collagen-based scaffold was manipulated to improve mechanical properties, cellescaffold interactions, oxygen flow, and nutrient diffusion [45]. Permeability is the ability of porous structures to transfer fluids through their interstices under applied pressure. This can be controlled through pore size, number, orientation, distribution, and interconnectivity [46]. Collagen scaffolds were prepared and exposed to plastic compression using different static stresses to control permeability. Results showed that increasing compression reduced k. It was also found that a decrease in k correlated with an increase in the modulus and permeability of collagen gels. The authors then tested the effect of k on MSC proliferation, differentiation, and mineralization. Compared with noncompressed gels, compressed ones showed higher proliferation, ALP staining, and mineralization, but no significant difference was found between the different compressed gels. These findings suggest that decreasing k provides a good matrix for cell proliferation and osteodifferentiation, but the influence of k on osteoinduction and osteoconduction has not been fully defined. BMP-4 was spatially immobilized in a collagenePLGA hybrid scaffold and has been shown to induce the osteogenic differentiation of osteoblasts to promote bone formation. MSCs were loaded into collagenePLGA scaffolds with or without BMP-4 and cultured in osteoinductive media before implantation into the dorsa of athymic nude mice. Type 1 collagen, OPN, and OCN showed a significant increase in expression compared with the control. However, there was no difference in ALP expression. This may have been because ALP expression increased before the analysis of the removed implants. In another study, a collagen-based silicified matrix was loaded with stromal celle derived factor-1 (SDF-1) [47], which is a chemokine-receptor ligand that is involved in immobilizing and homing stem cells to injured tissues. The silicified collagen scaffold (SCS) was first compared with CS in terms of its mechanical properties. The tangent modulus values (KPa) from 0%e5% strain were 0.80  0.21 for CS and 599.8  166.0 for SCS. The modulus of resilience values (KPa) were 0.18  0.06 for CS and 165.3  4.0 for SCS. In vitro analyses showed that there was no difference in cell viability when MSCs or endothelial progenitor cells (EPCs) were in contact with CS or SCS. The formation of extracellular bone nodules in differentiated MSCs was significantly higher in SCS than in CS, and the formation of capillary-like tubes by EPCs was much higher in SCS than in CS. These results indicate that the presence of silica in collagen hydrogels increases its osteogenic and angiogenic potential. The release kinetics of the SDF-1 from SCS hydrogels was analyzed and it was found that at high concentrations, up to 80% release could be obtained at 30 days. In vitro cell homing experiments were conducted with MSCs and EPCs using a transwell migration assay, employing SDF-1 concentrations that corresponded to those used in the release assay. Both MSCs and EPCs showed higher migration with higher SDF-1 concentrations. For in vivo experiments, the ability to promote osteogenesis was compared between SCS loaded with MSCs (cell-seeding approach) and cell-free SDF-1 containing SCS (cell-homing approach). Scaffolds were implanted into subcutaneous pockets of Balb/c mice. Implants were studied after removal and it was found that ectopic bone formation was similar in both scaffolds. However, ectopic bone formation in scaffolds loaded with SCF-1 showed a statistically significant increase in capillary formation. The results indicate that cell-homing strategies such as this must be further explored, because it reduces the complications of cell seeding and increases angiogenesis, a requirement for the formation of healthy bone. Hyaluronic Acid HAc has also demonstrated potential as a bone scaffold material. It is naturally occurring, hydrophilic, and nonimmunogenic, and has been found in the cytoplasm of osteoprogenitor cells [48]. This natural polymer has been used in combination with other materials, factors, and drugs to enhance its osteogenic potential. HAc, a nonsulfated glycosaminoglycan present in all connective tissue as a major constituent of the ECM and particularly prevalent during wound healing, has been proposed for the preparation of biodegradable ECM-like constructs for tissue engineering applications [49]. The ECM is a hydrophilic matrix with a gel-like uniformity consisting primarily of collagen fibers, proteoglycan filaments, and entrapped interstitial fluid. Similarly, hydrogels are also usually composed of covalently cross-linked hydrophilic polymers that enable them to swell while retaining their 3D structure without dissolving. This results in scaffolds structurally similar to the ECM, with a high water content and excellent biocompatibility characteristics. Cross-linking of HA forms experimentally controllable hydrogels that

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provide a microstructure similar to native ECM. Subsequent encapsulation of cells into permeable HA hydrogels thus provides both structural support and protection, as well as the ability for cells to interact in 3D, which has been shown to increase the viability of transplanted cells significantly [50]. Furthermore, cell encapsulation with biocompatible materials such as HA has been shown to reduce the immunogenicity of the tissue engineered construct, because biocompatible materials reduce the adsorption of proteins [51], which would normally stimulate the recruitment of immune cells such as macrophages. In one study, a photocured HAc hydrogel containing an osteogenesis-inducing drug, simvastatin (SIM), was designed [52]. SIM was found to induce osteodifferentiation of human adipose-derived stromal cells. Hydrogel viscoelastic properties were fine-tuned through 2-aminoethyl methacrylate (AEMA) substitution (HAc-AEMA). Three different HAc-AEMA scaffolds were studied: HAc with 20% (wt/wt) AEMA (HAc-AEMA-20), 30% AEMA (HAc-AEMA-30), and 40% AEMA (HAc-AEMA-40). Rheological measurements showed that elastic storage of the hydrogel increased with increasing AEMA concentrations (from 40 to 80 Pa). Pore size increased with increasing AEMA concentration, but sizes were not reported. MC3T3 fibroblasts were seeded into the hydrogels to test for cytocompatibility; it was found that the material was cytocompatible and that there was no difference in viability between groups. After SIM was loaded into the hydrogel, release kinetic experiments showed that it could sustain release for up to 14 days. HAc-AEMA-40 hydrogel was chosen for the remaining experiments. Loading of 1 mg SIM into HAc-AEMA-40 significantly increased fibroblast proliferation and mineralization at all times tested. The presence of SIM also upregulated the expression of OCN and OPN at all times tested. The effectiveness of this system was then tested in vivo with and without SIM. Hydrogels were implanted in parietal bone defects in rabbits. Cone beam computed tomography was used to assess healing of the bone defect for up to 9 weeks. Healing was only slightly superior at 9 weeks in the SIM-containing hydrogel. These results showed that substitution with AEMA can improve viscoelastic properties and the addition of SIM into the hydrogel improved osteogenesis in vitro, although the results were not as notable in vivo. Alginate AlginateeHA composite scaffolds were prepared by internal gelation followed by freeze-drying to obtain a porous structure. The nanoparticles were prepared in presence of a lactose-modified CHI; this colloidal solution was adsorbed on the scaffolds by exploiting electrostatic interactions and was used as temporary resorbable bone implants [53]. A CHIepolypyrroleealginate composite scaffold can act as a substrate for tissue regeneration and can be employed for bone tissue engineering using osteogenic cells by employing electrical stimulation with a bioreactor system, and thus evaluating the role of conducting substrate in bone regeneration [54]. CHIealginate hybrid scaffolds displayed improved mechanical strength and structural stability and were shown to stimulate new bone formation and rapid vascularization [55]. CHIealginate geleMSCeBMP-2 composites should have become clinically useful injectable materials to generate new bone [56]. Porous HAeCHIealginate composite scaffolds were prepared through in situ coprecipitation and freeze-drying for bone tissue engineering [57]. Two different types of polymer scaffolds, such as CHIealginate and CHIealginate with fucoidan, were developed by freeze-drying; each was characterized as a bone graft substitute by Venkatesan et al. [58]. Alginate microparticle and microfiber aggregated scaffolds were produced with alginate through the aggregation method. Such a porous structure will allow vascularization, oxygenation and cell migration, adhesion, and proliferation, which are biological events that are fundamental for bone tissue regeneration [59]. In another study, macroporous alginate scaffold was fabricated and mineral-coated using a biomimetic approach [60]. The nucleation of HA was achieved by incubating the scaffold in modified simulated body fluids for up to 4 weeks. Mineralization of the scaffold was determined by a change in mass, which increased steadily from weeks 1 to 4. Energy-dispersive X-ray spectroscopy revealed a Ca/P ratio of 1.61 when the Ca/P ratio of pure HA (HAP) was 1.67. Viability experiments demonstrated that the HAP coatings supported the attachment and proliferation of hMSCs. The cell number was significantly higher in the coated scaffolds at all times tested. This work showed that a simple scaffold modification, such as immersion in simulated biological fluids, can change the topography and environment of a scaffold to improve the osteogenic outcome of the scaffold. The in vivo application of a facile polyelectrolyte complexation (PEC) process was employed to condense heparin onto the surfaces of poly-L-ornithine (PLO), poly-L-arginine (PLA), and diethylaminoethyledextran-coated alginate microbead templates that entrap bioactive recombinant human BMP-2 (rhBMP-2). In vivo implantation of PEC shells loaded with rhBMP-2 resulted in new bone formation that could stimulate a mechanically stable posterolateral spinal fusion in rats (PLO- and PLA-based PEC shells). This ability to retain or regulate rhBMP-2 delivery and stimulate new bone formation from heparin-incorporated PEC shells could provide a powerful tool for entrapping and controlling the delivery of several growth factors in a variety of bone tissue engineering [61].

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Chitosan CHI, the deacetylated form of chitin, is the structural component found in the exoskeleton of crustaceans such as shrimp, crabs, and lobsters. It is a natural polymer, with a linear structure consisting of b(1e4)glycosidic bonde linked D-glucosamine residues with a variable number of randomly located N-acetyl-D-glucosamine groups [62]. CHI is bioactive, biodegradable, antibacterial, and biocompatible and possesses a hydrophilic surface, which is absent in many synthetic polymers. Here we describe several studies in which the osteogenic properties of CHI 3D scaffolds were modified by adding other polymers, cells, and bone-inducing factors. CHI and poly(butylene succinate) scaffolds were constructed and seeded with human MSCs to test their ability to induce osteogenesis [63]. The material was found to have 59% porosity, a 144.9-mm pore size, and 60.9% interconnectivity. MSC viability increased for 21 days in the scaffold and ALP activity also increased for 21 days after MSCs were exposed to osteogenic induction. The ability of the MSC-seeded scaffolds to repair a cranial critical-sized bone defect in mice was also examined. The cellescaffold constructs were cultured in osteogenic-induced medium for 2 weeks before implantation. Cell-free scaffolds were used as a control. The crania were observed 8 weeks after implantation. Bone formation was analyzed using mCT. The results showed an elevated rate of bone formation in the cell-seeded scaffold, evidencing the importance of the existence of stem cells near the area where bone formation is needed. In another study, human bone marrowederived stem cells (hBMSC) were encapsulated in hydrogels at CHI-collagen ratios of 100/0, 65/35, 25/75, and 0/100 wt%. b-Glycerophosphate was added to hydrogels, because it has been shown to be an osteogenic supplement when added to cultures of hBMSC, and it has also been used as a catalyst to sol-gel transitions in CHI hydrogels. The effect of adding collagen to CHI on matrix mechanical properties was assessed. Stressestrain profiles (0%e8% strain) showed that all collagen materials were approximately three times stiffer than pure CHI, which has a modulus of 6.3 kPa. When evaluating cell proliferation, the DNA content dropped by about half over 12 days in pure CHI materials whereas it increased twofold in materials containing collagen. For these reasons, only collagen-containing materials were examined for their effect on osteogenic gene expression. hBMSCs were encapsulated in CHI-collagen and collagen hydrogels and exposed to osteogenic medium. Hydrogels with a CHI-collagen ratio of 65/35 had the highest levels of osterix expression, bone sialoprotein expression, and ALP activity. These osteogenic markers started to decrease at lower CHI concentrations. These results suggest that the presence of collagen was highly beneficial for the osteogenic capabilities of the 3D scaffold, although it remains to be determined whether this was caused by the change in mechanical properties or the intrinsic biological properties of collagen. Another research group prepared porous 3D scaffolds based on CHI, CHI/SF, and CHI/SF/HA. SF/HA scaffolds were previously reported to be unsuitable for bone tissue engineering owing to insufficient formability and inflexibility [64]. The CHI scaffolds had a porosity of 94.2%  0.9%, which was statistically higher than the one presented by CHI/SF/HA scaffolds, which had a porosity of 89.7%  2.6%. The CHI/SF scaffold had a porosity of 91.6%  1.2%, which was not significantly different from other materials. SaOs-2 cells were used to measure viability and differentiation. At day 21, there was a statistically significant increase in cell proliferation and ALP activity in the CHI/SF/HA scaffolds compared with CHI and CHI/SF. It is unclear whether it was the presence of HA or the changes in porosity that promoted osteogenesis. CHI has also been used as an injectable biomaterial. Bi et al. [65] produced an injectable composite of tricalcium phosphate (TCP), CHI, and platelet-rich plasma (PRP). A TCP-CHI composite (TC) was used as a control. PRP contains a number of growth factors (i.e., PGDF, transforming growth factor-b, insulin-like growth factor, basic fibroblast growth factor [bFGF], and VEGF), which have often been shown to have an important role in bone tissue engineering applications. The composites were fabricated into cylinders for mechanical testing. It was found that the compressive strength (MPa) of both composites increased over time (w9e16 over 7 days) but there was no difference between them. The MSCs were seeded onto TCP and TC scaffolds and cell proliferation was measured over 7 days; the number of cells on TCP was found to be significantly higher than on TC at every time point tested. MSCs grown in plates in osteogenic media (OM) were used as a positive control for osteodifferentiation analysis of the materials being tested. ALP activity was tested at 7 and 14 days. At 7 days, ALP activity was higher in PTC than TC, and PTC and OM had comparable levels. At day 14, ALP activity was significantly higher in PTC than TC and even OM. RTPCR analysis was used to study the expression of osteogenesis-related genes. Runt-related transcription factor 2 (Runx-2), type-1 collagen and osteonectin expression in TCP were comparable to those in OM. Peptide Hydrogels Hydrogels are 3D networks composed of hydrophilic polymers cross-linked through covalent bonds or held together via physical intramolecular and intermolecular attractions. Hydrogels can absorb huge amounts of water

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or biological fluids, up to several thousand percent, and readily swell without dissolving. The high hydrophilicity of hydrogels results from the presence of hydrophilic moieties such as carboxyl, amide, amino, and hydroxyl groups distributed along the backbone of polymeric chains. In the swollen state, hydrogels are soft and rubbery, resembling living tissues to a great extent. In addition, many hydrogels such as CHI and alginate-based hydrogels have desirable biocompatibility [66], can easily be modified to contain bioactive motifs and composed of self-complementary amphiphilic peptides, and when gelled they provide a 3D structure that has many similarities to the ECM. In one study, self-assembling peptide nanostructured gels were constructed using peptide amphiphile (PA) materials with or without phosphoserine residues near their surfaces and with or without Arg-Gly-Asp-Ser (RGDS) peptide [67]. The phosphoresidues contribute to mineralization and the RGDS peptide contributes to cell adhesion. The authors proposed that the architecture of these PA materials is highly biomimetic of the fibrous elements commonly found in ECM, such as collagen fibrils. The results of this study suggested that a 3D matrix that supports mineralization and cell adhesion is favorable for bone formation, but the outcome was not superior to implantation with demineralized bone tissue. In another study, the commercially available peptidic hydrogel PuraMatrix was evaluated as a candidate to assess its potential for the osteogenic differentiation of dedifferentiated fat cells (DFATs) [68]. PuraMatrix is composed of 8e16 L-amino acid residues and forms a 3D scaffold that is biocompatible and biodegradable. The cell suspension was mixed at a 1:1 ratio with the hydrogel. These results indicate that commercially available PuraMatrix can support the osteogenic differentiation route and that DFATs are a viable alternative to MSCs that should further be explored with regard to bone regeneration.

Synthetic Polymers Copolymers A range of polyethylene oxide (PEO)epolybutylene terephthalate (PBT) copolymers (70%e30% PEO) was investigated for noneload bearing bone replacement. In general, copolymers are attractive for tissue engineering applications because their physicochemical properties are highly controllable. Gel formation dynamics, crosslinking density, and material mechanical and degradation properties can be controlled by regulating molecular weights, block structures, degradable linkages, and cross-linking modes [69]. b-Keto nitrile tautomeric copolymers have demonstrated fine-tunable hydrophilicity and hydrophobicity properties according to their surrounding environment and mechanical properties similar to those of human bone tissue. These characteristic properties make them promising candidates as biomaterials for bone tissue engineering. Based on this knowledge, we designed two scaffolds based on b-keto nitrile tautomeric copolymers that differ in chemical composition and surface morphology. Two of them were nanostructured, using an anodized aluminum oxide template and the other two were obtained by solvent casting. They were used to evaluate the effect of the composition and their structural modifications on the biocompatibility, cytotoxicity, and degradation properties. The results showed that the nanostructured scaffolds exhibited a higher degradation rate by macrophages than the casted scaffolds (6% and 2.5% of degradation for the nanostructured and casted scaffolds, respectively), a degradation rate compatible with bone regeneration times. We also demonstrated that the b-keto nitrile tautomericebased scaffolds supported osteoblastic cell proliferation and differentiation without cytotoxic effects, which suggested that these biomaterials could be useful in bone tissue engineering [70]. In one study, the authors had developed a novel BMP-2erelated peptide (P24) that was shown to enhance the osteoblastic differentiation of BMSCs. The purpose of that study was to incorporate P24 into a modified PLGA-(PEGeASP)n copolymer to promote bone formation [71]. PLGA and PLGA-(PEGeASP)n membranes were fabricated by solvent coating and evaporating. In in vitro release studies, about 70% of peptide was released after 14 days. Adhesion and proliferation studies were performed with BMSCs and the PLGA-(PEGeASP)n scaffolds, with and without p24, promoted cell adhesion and proliferation compared with the PLGA control. There was no difference between PLGA-(PEGeASP)n and PLGA-(PEGeASP) n-p24 scaffolds in either cell adhesion or proliferation. However, in cell differentiation studies, the PLGA-(PEGe ASP)n-p24 scaffold was shown to enhance differentiation at each time point tested for up to 20 days, as determined by ALP activity. A novel three-component biomimetic hydrogel composed of triblock PEGePCLePEG copolymer (PECE), collagen, and nano-HA (n-HA) was successfully prepared [72]. The in vivo biocompatibility of the PECEecollagene n-HA hydrogel composite was tested by implanting the composite into the dorsal muscle pouches of Wistar rats for 3, 7, and 14 days. Some inflammation occurred during the degradation process, but by day 14, the inflammatory response disappeared completely. The activity of bone regeneration was evaluated by reconstructing two rectangular defects in rabbit craniums. The untreated left defects were used as control and the right defects were packed

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with PECEecollagenen-HA hydrogel composite. At 20 weeks, as determined by histological staining, the right defect was filled completely with new bone and the amount of high-density tissue in the control defect was visibly less than the treatment defect. In another study, biodegradable and electroactive poly(ester amide)s containing conjugated segments of phenyl amino end-capped tetraaniline (PEA-g-TA) were prepared [73]. Composites were prepared under different feed weight percentages of tetraaniline (TA) (3.5%, 8.4%, and 15.5%). The copolymer solutions were cast onto a superflat polytetrafluoroethylene plate and placed for 5 h under room temperature to form thin films. Conductivity increased with an increasing percentage of TA (7.11  107, 8.01  106, and 2.45  106 S/cm for 3.5%, 8.4%, and 15.5%, respectively) but all scaffolds had conductivity values adequate to transfer bioelectrical signals in vivo. The PEA-g-TA number 2 copolymer (8.4% feeding ratio) was highly distensible with a breaking elongation rate of 105%  10%, and the tensile modulus of copolymer was 20  2.5 MPa. These values were not reported for the other scaffolds. Mouse MC3T3-E1 cells showed more than 90% viability in all copolymers tested. Cells were stimulated by pulsed electrical signal for 2 hours every day for 14 days. Osteogenic differentiation of MC3T3-E1 cells was assessed by the intracellular free calcium concentration and ALP enzyme activity. After 14 days, there were no significant differences between the materials tested; however, cells that were stimulated by the electric signal had higher intracellular calcium concentration and ALP enzyme activity. This work illustrates that combining appropriate materials with physical stimulation such as electrical impulses can produce superior responses. Polyesters Polyphosphazeneepolyester blends are attractive materials for bone tissue engineering because of their controllable degradation pattern with nontoxic and neutral pH degradation products. Aliphatic polyesters such as polyglycolic acid, polylactic acid, and polycaprolactone are the most commonly used polyesters for tissue engineering. Their degradation products are present in the human body and can be removed by natural metabolic pathways. 3D scaffolds from these materials can be fabricated through various techniques and fine-tuning the molar ratios of these polymers can influence mechanical properties and degradation rates [74,75]. In one study, PCL scaffolds incorporating HA particles were fabricated by combined solvent casting and particulate leaching [76]. The presence of HA slightly increased the density of the scaffold but had no significant effect on the porosity values of the scaffolds. With increasing HA concentrations, there was a significant correlation with an increase in the compressive modulus. Viability assays showed that the number of primary human bone cells that had been cultured on the PCLeHA scaffolds was greater than that of cells that had been cultured on the neat PCL scaffolds at both 24 and 48 h after cell culturing. Expression levels of type I collagen and osteocalcin were evaluated by RT-PCR and were significantly greater in cells cultured on the PCLeHA scaffolds than those in cells cultured on the neat PCL scaffolds on day 10. The formation of mineralized nodules of cells cultured on the PCLeHA scaffolds was significantly greater than that of cells cultured on the neat PCL scaffolds. For in vivo experiments, a circular calvarial defect in mouse was used as a model. Six weeks after implantation, histomorphometric analysis indicated a statistically significant increase in the amount of new bone formation in the PCLeHA scaffolding implants compared with neat PCL counterparts.

Ceramic Scaffolds Bioactive ceramics have received great attention owing to their success in stimulating cell proliferation, differentiation, and bone tissue regeneration. They can react and form chemical bonds with cells and tissues in human body. These ceramics are recognized as bioceramics and are classified into two groups: bioinert or bioactive. Bioactive ceramics are categorized as resorbable or nonresorbable [77]. Ceramics are used because of their chemical properties and crystallinity, which is similar to bone mineral components. These materials exhibit excellent biocompatibility and bioactivity. The inorganic fraction of bone is composed of HA and CaPs, which allow the formation of bone tissue on its surface. This type of material is excellent as an implant, but it has particular problems with its mechanical properties in terms of fracture and fatigue. Common ceramic materials used for bone repair or regeneration are Bioglass, CaPs, and ceramic scaffold derived from corals. Here, we report some studies in which these materials were studied both in vitro and in vivo to assess their osteogenic potential. Bioactive ceramics are known to enhance osteoblast differentiation as well as osteoblast growth. However, their clinical applications have been limited because of their brittleness, difficulty in shaping, and extremely slow degradation rate in the case of HAP. Also, they have poor fidelity and reliability, and new bone formed in a porous ceramic scaffold cannot sustain the mechanical loading needed for weight-bearing bone [78]. Wang and Shaw [79] reported that dense HAP ceramics with a fracture toughness of 0.61e1.06 MPa$m1/2 were fabricated via conventional sintering. Fielding et al. [80] fabricated TCP scaffolds with a compressive strength of 1.75e5.48 MPa using commercial 3D

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printing technology. However, the mechanical properties of fabricated scaffolds were far below those of weightbearing bone (fracture toughness ¼ 2e12 MPa$m1/2; compressive strength ¼ 130e180 MPa) in the human body [81]. Therefore, obtaining an effective method to overcome these limitations has become the focus of current and future research in bone tissue engineering. Calcium Phosphate Calcium phosphate ceramics (CPCs) are a class of fine-tunable bioactive materials that have been widely used for bone tissue repair and augmentation. They possess surface properties that support osteoblast adhesion and proliferation (i.e., osteoconduction) and stimulate new bone formation (i.e., osteoinduction) [82]. TCP was first used in 1920 by Albee [83]. The study reported cases of fracture with bone loss in which more rapid bone growth and union were observed when TCP was injected into the gap between the bone ends than did the controls without its use. Also, it was demonstrated that osteogenesis was stimulated by this material in conjunction with the fracture microenvironment. In 1951, Ray and Ward [84] reported the use of granular synthetic HA to repair defects in long bone and iliac wings of dogs and bur holes in the crania of cats; the study demonstrated that this material could be replaced by new bone, but it was not as effective as autologous bone grafts in repairing the defects. Calcium phosphateebased scaffolds exhibit osteoconductivity, bioactivity, and resorbability in vivo owing to their complex chemical composition (Ca/P ratio) and physical properties such as their crystallographic structure and porosity [85]. In studies combining CPC paste and hydrogel microbeads to encapsulate human umbilical cord MSCs (hUCMSCs) with CHI fibers [86], hUCMSC viability and differentiation capacity were observed. Nanostructured CaP biomaterials and scaffolds mimic natural bone, and have high surface-to-volume ratios, improved wettability and mechanical properties, and increased protein adsorption and other desirable properties, compared to conventional counterparts. Nano-CaP biomaterials have emerged as a promising class of biomimetic and bioactive scaffolds capable of directing cell behavior and cell fate and enhancing tissue formation in vivo. In general, nano-CaP scaffolds can support stem cell attachment and proliferation and induce osteogenic differentiation, in some cases without osteogenic supplements. The influence of nano-CaP on cell alignment is less prominent than that of polymers and metals owing to the no-uniform distribution of the nano-CaP crystals. Nano-CaP biomaterials can achieve significantly better bone regeneration in vivo than conventional CaP biomaterials. The combination of various types of stem cells with nano-CaP scaffolds can further accelerate bone regeneration, the effect of which can be even further promoted by growth factor incorporation. Cell microencapsulation combined with nano-CaP scaffolds is a promising tool for bone tissue engineering applications to distribute cells throughout the interior of the scaffold [87]. More studies are needed to compare various types of nano-CaP compositions and nanostructures side by side in vivo and to compare the efficacy of various types of stem cells in bone regeneration. Bioglass Since the discovery of 45S5 bioactive glasses by Hench, they have been frequently considered scaffold materials for bone repair [88]. The need to find a material that forms a living bond with tissues led Hench to develop Bioglass repair tissues during the Vietnam War [89]. Bioglass offers advantages such as a controlled rate of degradation, excellent osteoconductivity, bioactivity, and the capacity to deliver cells, but they have limitations in certain mechanical properties such as low strength, toughness, and reliability [90]. Advantages of the glasses are ease in controlling the chemical composition and thus, the rate of degradation, which make them attractive as scaffold materials. The structure and chemistry of glasses can be tailored over a wide range by changing the composition or the thermal or environmental processing history. Therefore, it is possible to design glass scaffolds with variable degradation rates to match those of bone ingrowth and remodeling. A limiting factor in using bioactive glass scaffolds to repair defects in load-bearing bones has been their low strength [91]. Work has shown that by optimizing composition, processing, and sintering conditions, bioactive glass scaffolds can be created with predesigned pore architectures and with strength comparable to that of human trabecular and cortical bones [92]. Another limiting factor of bioactive glass scaffolds has been the brittleness. This limitation has received little interest in the scientific community, judging from the paucity of publications that report on properties such as fracture toughness, reliability (i.e., Weibull modulus), or work on the fracture of glass scaffolds. Wu et al. [93] prepared 45S5 Bioglass by foaming with rice husks and sintering at a high temperature; they obtained favorable results in compressive testing and degradability in simulated body fluid. That study reported compressive strength values in the range of trabecular bone. Porous bioactive glass-ceramic (45S5) was tested with human umbilical vein endothelial cells (HUVECs) and human osteoblast-like cells (HOBs). The results of the study demonstrated that the proliferation of HOB and HUVEC cocultures seeded on scaffolds was higher than that of commercial HA scaffolds. Wu et al. [93] studied HOB attachment, proliferation, and differentiation

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on a bioactive diopside scaffold (CaMgSi2O6). Cell proliferation increased with incubation time. In addition, porous CaSiO3 and kermanite (Ca2MgSi2O7) ceramic scaffolds were compared with b-TCP ceramic scaffolds, demonstrating in vivo bone formation and potential applications in bone tissue regeneration [94]. Other studies related to the formation of porous Bioglass were developed by Moawad and Jain [95]. They fabricated nanomacroporous soda lime phosphosilicate glass scaffolds using sucrose as a macropore former, and established process parameters such as the weight ratio of glass/sucrose, the particle size of glass/sucrose powders, and the time and temperature of sucrose dissolution. Coral Coral exoskeleton (CaCO3), which has an interconnected pore structure that resembles that of natural human bone, has been used as a scaffold material to fill bone defects in both animal models and humans since the early 1970s. This natural material is biocompatible, osteoconductive, and biodegradable. Most important, the possibility of seeding coral scaffolds with stem cells or loading them with growth factors has provided a novel alternative for bone tissue engineering. Corals are attractive materials for scaffolds because they have microstructures with highly controlled pore sizes and an interconnected porous architecture similar to trabecular bone [96,97]. After 1991, Ripamonti reported the first study of the morphogenesis of bone in porous bioceramics (coral-derived HA) in the rectus abdomen in the muscle of adult nonhuman primates Papioursinus demonstrated the use of coral line replicas of HA with an average pore size of 600 mm as bone grafts for the controlled formation of bone in humans [98]. Coral scaffolds using adipose-derived MSCs (AMSCs) have been used to repair cranial bone defects in a canine model [99]. They showed adipogenic differentiation and good biocompatibility to support the proliferation of AMSCs in vitro. Subsequent implantation in vivo showed new bone formation. Certain types of proteins such as ALP, OPN, and OCN are employed to identify the activity of osteoblast-specific proteins. Suzina et al. [100] reported an increase in the expression of specific genetic markers from osteoblasts such as Runx-2, OPN, ALP, and OCN in the Porites Goniopora coral, which was implanted in orthotropic calvarial defects of the adult nonhuman primate Papio ursinus. In vivo results showed an increase in ALP activity from 2 to 12 months and induction of bone in the concavities; however, limited conversion of HAecalcium carbonate was observed. A preliminary study in nude mice reported the vascularization of tubular coral scaffold with cell sheets [101]. The results showed that cells promoted new bone formation through an endocrine process. In addition, Zheng et al. [102] evaluated the feasibility of mandibular condyle constructs engineered from hBMSCs. In vitro studies reported that hBMSCs induced differentiation into osteoblasts and chondroblasts, and seeded scaffolds implanted into nude mice showed neovascularization in the temporomandibular joint detected by bFGF expression.

Metallic Scaffolds Several biocompatible metallic materials are frequently used as implanting materials in dental and orthopedic surgery to replace damaged bone or provide support for healing bones or bone defects. Standard surgical implant materials include stainless steel 316 L (ASTM F138), Co-based alloys (mainly ASTM F75, and ASTM F799), and titanium alloys; Tie6Ale4V (ASTM F67 and F136) is the most commonly employed. However, the main disadvantage of metallic biomaterials is their lack of biological recognition on the material surface. To overcome this restraint, surface coating or surface modification offers a way to preserve the mechanical properties of established biocompatible metals improving the surface biocompatibility. In 1909, the first patent of a metallic framework for an artificial tooth root for fixation by bone in growth was accredited to Greenfield [103,104]. He recognized the limitations of natural tooth implantation and started experimenting with implanting artificial hollow cylinders made of iridoplatinum wire soldered with 24 kt gold. In 1971, Galante et al. [105] were pioneers in developing open-pore fiber metals for clinical use as porous coatings in hip and knee arthroplasty. Because metals are materials with high mechanical strength and fracture toughness, they are frequently used as metallic biomaterials in the dental and orthopedic fields to replace and offer support for damaged and healing bone [106,107]. The commonly metals used as standard surgical implants were stainless steel 316 L (ASTM F138), Co-based alloys (mainly ASTM F75, and ASTM F799), and titanium alloys. where Tie6Ale4V (ASTM F67 and F136). However, these metallic biomaterials have disadvantages such as the possible release of toxic metallic ions and/or particles through corrosion or wear processes that cause inflammation and allergic reactions, which affect biocompatibility and tissue loss. Also, they produce poor stimulation of new bone growth owing to the elastic moduli, which does not correspond with natural bone tissue. Despite this, it has been reported that Ti-based metals can be used as bone substitute because of its elasticity, mechanical properties, shape memory effect, porous structure, and biocompatibility [108]. Also, 3D microporous NiTi and Ti scaffolds have been produced by powder metallurgy that obtains a hydrophilic surface to facilitate the deposition of HA and stimulate cell attachment and proliferation.

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Hybrid Materials The material needs to be bone bioactive: that is, it should encourage bone ingrowth and degrade at a rate that allows the newly formed tissue to replace the scaffold gradually, both as a mechanical structure and in terms of the space occupied. Finally, and this is where most current materials fail, the material needs mechanical properties that allow the device to be implanted without failing. This last requirement allows a patient to use the implanted area without mechanical protection, such as a cast, but still enables sufficient loading of the newly formed tissue to stimulate the osteoblasts mechanically. As yet no one has reported a material that fulfils all of these requirements. One group of materials that attempts to fulfill many of these requirements is composites of degradable polymers reinforced with ceramics, glass-ceramics, or bioglasses. If the polymer is biodegradable and the ceramic, glassceramic, or bioglass phase is degradable or metabolized by the body, the degradation requirement is fulfilled. The use of ceramics or glass-ceramics can both stiffen and strengthen a low-modulus, low-strength polymer and increase the bioactivity of the composite [109]. Commonly, composite scaffolds are fabricated using a different type of matrix with a dispersed phase such as polymereceramics, ceramicemetals, and polymeremetals [110]. PolymerePolymer Blends A polymerepolymer blend defines a mixture of two different polymers addressing each other’s drawbacks and giving synergetic advantage to the scaffold design. A miscible composite design with desired features can be fabricated by employing polymers with particular intermolecular or van der Waals interactions. One important example is PLGA and polyphosphazene bends [111]. PLGA is a well-known biomaterial in tissue engineering. Owing to PLGA acidic by-products upon degradation, it has been a critical issue for its further use, because its long-term tissue exposure to acidic products may lead to tissue necrosis and implant failure. On the other hand, polyphosphazene releases neutral or basic products upon degradation. Thus, PLGA has been blended with a wide arena of polyphosphazenes to achieve near-neutral degradation products as an efficient construct for tissue engineering [112e114]. PolymereCeramics Blends Many researchers have studied polymereceramic-based scaffolds. Combinations of CaPepolyesters for scaffold production were investigated by Kaplan and coworkers [115], who developed a scaffold from SFepoly aspartic acid coated with CaP. In vitro studies showed cell viability, proliferation, and osteogenic differentiation; however, nonuniform distribution of cells resulting from mineral deposition was observed. Li et al. [115] used two biocompatible and biodegradable polymers, PLGA and PCL, with a layer of CaP and a gelatin coating. MC3T3 cells adhered to regions with higher CaP content along the scaffold, which indicated that the mineralization gradient affects the adhesion and proliferation of cells and the physical properties of the scaffold. Relative to the mechanical properties, the local strain varied along the long axis of the scaffold and the Young’s modulus increased with increasing levels of CaP. TCP has been used to generate biocompatible and biomechanical scaffolds [116]. In vivo studies using this type of scaffold reported significant biocompatibility, sufficient mechanical strength, osteogenic differentiation, and bone growth. It was also demonstrated that the incorporation of collagen into this type of system improves hydrophobicity and differentiation [117]. Wang et al. [118] reported a novel scaffold composed of the PLGAeb-TCP skeleton wrapped with type I collagen. The aim of that study was to analyze the physical properties and biocompatibility of the coreesheath structure composite scaffold in vitro compared with the PLGAeb-TCP skeleton. No differences were observed in the porosity ratio, compressive strength, or Young’s modulus of the scaffolds. MTT assay indicated that BMSCs showed better adhesion and proliferation activity on the surface of the coreesheath structure composite scaffold than on the PLGAeb-TCP skeleton. In a previous study [119], this scaffold was fabricated via low-temperature deposition manufacturing but its hydrophobicity did not promote adequate cell adhesion, proliferation, or osteoblastic differentiation. Thein-Han and Xu [120] were the first to develop a novel scaffold from collageneCPC using hUCMSCs for in vitro analysis. Good mechanical properties, hUCMSC attachment, viability, and osteogenic differentiation were observed. Previously, studies were conducted using collageneCaP scaffolds, but they were not moldable or injectable [121]. HA was also employed to fabricate ceramicepolymer scaffolds [122e124]. Kim et al. [125] developed this type of scaffold and demonstrated that the combination of PGAePLGAe HA improves their osteogenic capacity. Hollinger et al. [125] used a combination of poly(ε-caprolactone fumarate) ePVAeHA with rhBMP-2 and the preosteoblast cell line MC3T3-E1 obtained good cell viability and proliferation, bioactivity, and bone regeneration and an increase in compressive modulus with an increase in HA; however, mCT showed a nonuniform distribution of the porogen indicating that the degradation process, porosity, and pore

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interconnectivity needed to be improved. Geng et al. [126] fabricated a scaffold from a magnesium phosphate (MP)epoly(ε-caprolactone) composite in which the mechanical analysis showed low values of the compressive modulus compared with trabecular bone; thus, this scaffold was found not to be appropriate for higheload bearing applications. It was also reported that the degradation rate of the composite scaffolds could be modulated by varying the amount of MP particles introduced into the polymer matrix. MetalePolymer Blends The development of successful scaffolds for bone tissue engineering requires a concurrent engineering approach that combines different research fields. Researchers have tailored metallic scaffolds that are useful for a wide variety of medical and dental applications. Surface modification of already proved biocompatible metals is an essential requisite for their use in tissue engineering because the metal surface must be controlled to induce the adhesion and proliferation of cells and the adsorption of essential biomolecules. Chemical and physical properties have a crucial role in the osteointegration of implant surface; they allow protein adsorption between implanted biomaterials and the biological environment. Thus, numerous strategies have been developed to create a bond between the implant and the living host tissue. The use of metallic implants with a polymer coating has been reported by many for these purposes. Ti, TiO2, and Ti-alloy combined with polyester coatings have been intensively studied to fabricate scaffolds. Lagoa et al. [127] developed a partially biodegradable implant from titanium, polylactide, HA, and calcium carbonate. The implant had mechanical stability, biocompatibility, and partial biodegradability. Helary and coworkers used poly(sodium styrene sulfonate) (polyNaSS) on oxidized or grafted Ti samples [128] and polyNaSSe(methylacrylic acid) MA grafted onto Ti6Al4V alloy surfaces [129] They reported that cell adhesion and differentiation on Ti (grafted) were higher than on oxidized titanium and titanium because of the presence of active sites that interact with extracellular proteins. In addition, polyNaSSeMA grafted onto Ti6Al4V alloy in femoral rabbit model showed lamellar trabecular bone with wide haversian canals lined by osteoblasts; however, the high levels of Na, P, Ca, Zn, etc., indicated that the piranha treatment used to oxidize the alloy surface was intense. Oughlis et al. [130] developed a scaffold from polyNaSS polymer and titanium. The results showed cell viability, proliferation, and osteoblastic differentiation of hMSCs on this scaffold. MetaleCeramic Blends Many researchers have reported the development of metaleceramic scaffolds that have been shown to possess favorable characteristics regarding their mechanical properties and bioactivity (cell attachment, proliferation, and differentiation). Yang et al. [131] developed a biodegradable and bioactive scaffold composed of magnesium with a coating of b-TCP. The in vitro results showed good cell adhesion and bioactivity. In similar studies, when tantalum and titanium were used as metals, osteoconductivity and osteoinductivity were improved in vivo. Wu et al. [132,133] reported the fabrication of a hypoxia-mimicking mesoporous bioactive glass (MBG) scaffold by incorporating Co2þ ions into the MBG scaffold. Results showed that MBG scaffolds had no significant cytotoxicity and that the incorporation of ionic Co2þ ions enhanced VEGF secretion, HIF-1a, expression, and bone-related gene expression in BMSCs. Also, the CoMBG scaffolds supported BMSC attachment and proliferation. It was reported that ZrO2 itself does not have good cellular and tissue affinity [134] although in vitro and in vivo studies demonstrated that ZrO2 is not toxic [135]. Lee et al. [136] fabricated a scaffold composed of (biphasic CaP) ZrO2 and PCL layers. The scaffold had a compressive strength of 12.7 MPa and a porosity of 78 vol% and showed excellent MG63 cell attachment, and OCN and collagen expression. It was also shown that PCL incorporated into biphasic CaP gave scaffolds high biodegradability, cell attachment, and proliferation. Li et al. [137] reported a novel ion doping method applied in calcium polyphosphate (CPP)-based bioceramic scaffolds to be substituted by potassium and strontium ions (K/Sr). The K/Sr CPP scaffolds had a higher compressive strength, cell biocompatibility, biodegradability, osteoinductivity, and osteoconductivity and better degradation properties than the pure CPP scaffolds; however, the mechanical strength of K/Sr-CPP was not good after degradation. Studies suggested that directly mixing Cu2þ ions with bioactive materials improves angiogenesis [138]. Xiao et al. (2013) prepared copper (Cu)-containing mesoporous bioactive glass (Cu-MBG) scaffolds with interconnected large pores. Attachment, proliferation, and ALP activity of hBMSCs on CueMBG scaffolds were observed, and ionic products of CueMBG extracts enhanced the osteogenic differentiation of hBMSCs. Studies using porous ceramice coated TiO2 as scaffolds were reported by Dimitrievska et al. [139]. They evaluated the cell adhesion, growth, and osteoblastic differentiation of hMSCs in the TiO2eHA nanocomposite; however, applications in bone implants were limited by their bioinertness. Also, Haugen et al. [140] reported in vitro results that showed the cell viability of the mouse osteoblastic cell line MC3T3-E1. No cytotoxic effects from the TiO2eHA scaffolds were found. It was also reported that healing occurred when this scaffold was implanted into rabbit defects.

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CONCLUSIONS AND FUTURE PROSPECTS We have summarized the current state of the art of materials technology and the development requisites for successful bone graft substitute. Clearly, the smart selection of materials is the key to successful tissue engineering. Nevertheless, various other factors also influence the overall outcome. Scaffold composed of natural or synthetic biomaterials or their combination should closely mimic the ECM in native tissues, provide the required support, promote neovascularization, and allow access to nutrients to support the entire process of tissue regeneration. All of these areas of advanced, fast-growing interest and expanding research demonstrate the multidisciplinary nature of tissue engineering and the field of regeneration medicine. There are vast challenges as well as wide opportunities to improve human health immensely in a variety of areas. However, much study is needed to explore various novel materials and the dynamics of the bone tissue microenvironment for proper simulation.

Acknowledgments This research was supported by Basic Science Research Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Science, ICT & Future Planning (NRF-2017R1A2B3010270) and the Korea Health Technology R&D Project through the Korea Health Industry Development Institute (KHIDI), funded by the Ministry of Health &Welfare, Republic of Korea (grant number : HI15C2996).

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C H A P T E R

41 Materials-Based Cancer Immunotherapies Jared M. Newton1, Andrew G. Sikora1, Simon Young2 1

Baylor College of Medicine, Houston, TX, United States; 2The University of Texas Health Science Center at Houston, School of Dentistry, Houston, TX, United States

INTRODUCTION AND OVERVIEW OF CANCER IMMUNOTHERAPY The immune system is composed of two primary cellular subsets; innate and adaptive. Cells of the innate immune system are often the first responders to reach a site of injury. One of their primary functions is to contain locally and eliminate, if possible, any potentially dangerous substances, which are commonly termed pathogens. In addition, a number of innate immune cells function to collect small pieces of the dangerous substance, termed antigens, and present them to cells of the adaptive immune system. Macrophages and dendritic cells (DCs) are two key cells that perform this function, and thus are appropriately known as antigen presenting cells (APCs). After successful presentation and stimulation by APCs, adaptive immune cells, namely antigen-specific cytotoxic T cells, can mount a more pathogen-specific immune response with long-term components that can protect the host more rapidly if reexposed to that particular antigen. In response to identifying features of pathogens, APCs can release specific combinations of cytokines, chemokines, and other cues to promote T-helper 1 (Th1) or T-helper 2 (Th2) adaptive immune responses. Th1 immune responses induce cell-mediated responses, typically provoking T-cell proliferation and investigation of intracellular pathogens. Alternatively, Th2 immune response induce humoral immunity, promoting B-cell antibody production targeting primarily extracellular pathogens. Although greatly simplified, this is one of the primary processes that the immune system uses any time it encounters potentially harmful substances, whether that be a virus, bacteria, or cancer cell. Cancer presents a variety of challenges for the immune system as evidenced by the fact that avoidance of immune destruction has been recognized as a primary hallmark of cancer [1]. One of the most difficult challenges is the fact that cancer cells are not foreign entities, and thus look similar to normal healthy tissue, which makes it difficult for the immune system to identify and eliminate them. As a result, the adaptive immune system often struggles to present tumor-associated antigens in an immunogenic context that is critical for generating a productive immune response. However, cancer has long been recognized as a disease mediated by genomic mutations, whether naturally or virally induced, and if these genomic mutations result in mutated proteins, the immune system theoretically has the potential to recognize that. As proof of this, tumors with a higher mutational burden are typically more immunologically active, with higher immune cell infiltrate and a greater potential for immunemediated attack [2]. In addition to mutated antigens, differential expression of self-antigens provides another mechanism of immunologic recognition of cancer (i.e., human estrogen receptor 2 in breast cancer, prostatespecific antigen in prostate cancer, cancer-testes antigens). Another challenge that hinders immune recognition and destruction of cancer is the highly immunosuppressive protumor microenvironment observed in solid tumors. This occurs largely through a wound healingelike intratumoral inflammation and an increase in suppressive immune cell populations such as myeloid-derived suppressor cells (MDSCs), tumor-associated macrophages (TAMs), and T-regulatory cells (Tregs) [3]. In light of these obstacles, cancer immunotherapeutic strategies are warranted to direct and stimulate an effective immune response against cancer. The idea of using one’s own immune system to fight cancer is by no means a novel concept; in fact, this concept has existed for over 100 years. Perhaps the oldest recorded evidence came from an American surgeon, William B. Coley, when he demonstrated successful remission in numerous inoperable cancer patients after intratumoral

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injection of a mixture of bacterial lysates known as Coley’s toxin [4]. In the 90 years after Coley’s findings, the fields of immunotherapy and immunology matured in concert with one another. Significant immunotherapy achievements include the development of the first immune adjuvants [5], proposal of the immune surveillance theory or the idea that cancer cells could be recognized and killed by components of the immune system [6], proof that tumors have tumor-specific antigens that can be targeted by the immune system [7e10], and proof of tumoral immune escape through the loss of tumor-specific antigens [11]. During this same time, we note the discovery of numerous important immune mediators such as the Treg [12], DCs [13], natural killer cells [14], and interferons (IFNs), which are a family of critical protein mediators that act as signals for the immune system to regulate various aspects of an immune response [15]. Each of these discoveries and countless more were monumental in the development of the field of cancer immunotherapy. These discoveries led to some of the first organized cancer immunotherapy clinical trials. Pioneering trials by Steve Rosenberg and colleagues, and others, aimed to use systemic administration of interleukin-2 (IL-2), a potent T celleactivating cytokine, in treating melanoma and renal cancer [16,17]. Despite promising preclinical data, clinical trials with IL-2 appeared to generate little improvement in overall survival and were often associated with concerning side effects. The early 1990s also saw the development of several promising new cancer vaccine strategies. One such strategy was GVAX, a vaccine that used irradiated cancer cells genetically modified to secrete granulocyte macrophageecolony-stimulating factor (GM-CSF), a potent APC activator [18]. A second successful strategy involved the use of autologous DCs pulsed with cancer peptides that could be reintroduced to the patient [19]. Both strategies were found to elicit potent antigen-specific T-cell responses and led to the development of Sipuleucel-T, or Provenge, for the treatment of prostate cancer. Sipuleucel-T treatment requires the isolation of patient-derived DCs and incubation with both an antigenic target overexpressed on most prostate cancers (i.e., antigenic prostatitic acid phosphatase) and GM-CSF for stimulation. Blinded, placebo-controlled phase III clinical trials showed a 4-month improvement in median survival for patients receiving Sipuleucel-T compared with placebo-treated cohorts [20]. This eventually led to the US Food and Drug Administration (FDA) approval of Sipuleucel-T for the treatment of prostate cancer in 2010. Sipuleucel-T represents the first major cancer immunotherapy to receive FDA approval and ultimately paved the way for numerous immunotherapies. A variety of other cancer vaccine strategies are currently under investigation, which vary according to what stimulus they provide and whether they deliver only a single antigen, termed subunit vaccines, or a whole library of tumor antigens (i.e., tumor lysates, irradiated whole tumor cells) [21]. During this period, adoptive T-cell therapies were also showing promise as an effective cancer immunotherapy. Applications include genetically modified T cells made to recognize and kill specific tumor antigens or chimeric antigen receptor (CAR) T-cells and adoptive T-cell therapies in which tumor reactive T-cells isolated from the patient were stimulated ex vivo and then reinfused [22e24]. Despite showing significant success in blood-borne cancers (i.e., lymphoma, leukemia), T-cell therapies continue to show minimal benefits against established solid tumor cancers; however, numerous technical advances and clinical trials are under way investigating new approaches and combinatorial treatment strategies with these therapies. The mid 1990s saw the emergence of an immunotherapy platform termed immune checkpoint inhibitors (ICIs), which continue to show significant potential as a solid tumor therapy. Two ICI targets of extensive investigation are cytotoxic T lymphocyteeassociated antigen-4 (CTLA-4) and programmed cell death protein-1 (PD-1). CTLA-4 is expressed on activated T cells and functions as the “brakes” of a T-cell response by saturating away costimulatory molecules on activated APCs and promoting overall deactivation of the T cell [25,26]. Blockade of CTLA-4 using monoclonal antibodies has been shown to prolong T-cell activation and promote significant antitumor effects in murine models [27,28]. Phase III clinical trials of Ipilimumab, a humanized antibody targeting CTLA4, improved median survival by 4 months for patients with advanced melanoma, with a small subset of patients maintaining cancer-free progression up to 10 years later [29,30]. This led to the FDA approval of Ipilimumab in 2011 to treat advanced melanoma. CTLA-4 blockade is currently being tested in numerous other cancers both as a monotherapy and combined with other treatment modalities. Success of CTLA-4 inhibition led to the investigation of numerous other ICI targets, one of particular success being PD-1. PD-1 functions to deactivate T cells after engagement with either of its known ligands, PD-L1 or PD-L2 [25]. Of interest in cancer, high PD-L1 expression has been observed in numerous solid tumor cancers and has been correlated with a poorer prognosis [31e33]. In addition, cancer-induced exhaustion of T cells appears to upregulate PD-1 drastically on their surfaces, making them more susceptible to deactivation through this pathway [34]. Strong evidence suggests that blockade of PD-1 can restore functionality to these exhausted T cells and promote significant antitumor effects [35,36]. Thus, after clinical investigation, both nivolumab and pembrolizumab, two humanized anti-PD-1 monoclonal antibodies, received FDA approval in 2014 for the treatment of unresectable melanoma. Since then, they gained FDA approval

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for the treatment of lung and kidney cancers in 2015 and Hodgkin lymphoma in 2016 and are actively being investigated in numerous other cancer types, again as both monotherapies and combined with other treatment modalities. Despite preclinical and some clinical success of immunotherapies in treating cancer, it should be noted that although small subsets of patients have full durable responses, most patients do not benefit from these therapies. In addition, many of these therapies result in a number of systemic side effects, many of which could have lifethreatening or long-term consequences for the patient’s health. This leaves much to be desired from the field of immunotherapy in treating cancer.

ADVANTAGES AND DISADVANTAGES OF CANCER IMMUNOTHERAPY Cancer immunotherapy offers numerous advantages compared with other standard of care treatments. Most notably, immunotherapy optimizes the patient’s own natural immune system as the primary therapeutic source. This allows use of the intricate biologic features of the immune system that engineered therapeutic systems cannot currently achieve. Examples of these features include the natural ability of T cells to traffic and extravasate into tumor tissue, the ability to target cells preferentially harboring specific aberrations while leaving surrounding healthy tissue relatively unharmed, the ability to seek and destroy cancer cells throughout the entire body, even micrometastatic lesions, and finally, features of immunologic memory that can protect against recurrence. These advantages are highly unique to immunotherapy and thus provide the rationale for why it remains an exciting and highly investigated therapy in the fight against cancer. Despite these numerous advantages, current immunotherapy applications face a number of challenges. Most notable is their inconsistent efficacy, because the majority of patients do not benefit from current immunotherapeutic strategies. Even more challenging is that currently few correlative biomarkers can predict patient success from these therapies. This has led numerous groups to investigate combinatorial treatment strategies in the hope of achieving synergistic treatment effects. Some examples include combinations of immunotherapies with chemotherapy, radiation, small molecule inhibitors, and even combined with other immunotherapy strategies [37]. In addition to low success rates, a number of immunotherapies have shown concerning side effects. Most often, these side effects are associated with systemic hyperactivation of the immune system, which can promote autoimmune responses against vital organ systems, leading to long-term and sometimes deadly health consequences for patients. The last major challenge of immunotherapy is affordability. Many current immunotherapies require isolation of patient-derived cells (i.e., T cells, DCs, tumors cells), which are then modified and readministered to the patient. This personalized medicine approach requires an extensive manufacturing pipeline and is often associated with vastly higher costs than a typical “off-the shelf” cancer drug such as chemotherapy or small molecule inhibitors. Nonepatient derived immunotherapies are often expensive as well, typically requiring months to years of clinical visits for therapeutic administration. This creates major issues in terms of patient affordability and compliance. Thus, despite significant efforts and advances in the field of immunotherapy, it continues to remains significantly hindered by its limitations. This is where biomaterials-based approaches could have a significant role. Whether macro-, micro-, or nanoscale, biomaterials offer the potential for sustained and/or targeted release of drugs (Fig. 41.1). For immunotherapeutic applications, this simple feature could provide a significant opportunity to overcome many of its current challenges. Biomaterials would provide a platform to contain an immune response locally within target areas. This would greatly benefit immunotherapies because the desired immunomodulation could occur in important areas such as in a subcutaneous zone, within the tumor itself, or even within critical immunologic tissues (i.e., lymph nodes and spleen). This could benefit the safety of immunotherapies by eliminating systemic immune activation and also would likely improve their efficacy, because more of the delivered drug would come into contact with the effector cell types rather than being removed from the systemic circulation or be degraded through normal physiologic routes. In addition, advances in biomaterial technology allow for accurate and highly controllable fine-tuning of drug release, in some cases providing extended release for up to years. Numerous groups have also developed systems for the controlled release of more than one drug with the potential to control the release kinetics of each drug individually. These multidrug release systems could enable the delivery of combinatorial immunotherapies to mimic the true response progression of the immune system more closely (i.e., first priming the innate system and then providing activation markers for the acquired immune system to maintain that response). This has the potential to enhance the poor response outcomes of immunotherapy greatly and could minimize the necessary time that patients need to spend in the clinical

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Cancer vaccines • NPs for subunit vaccines • scaffolds for immune cell programming

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FIGURE 41.1 Schematic of the cancer-immunity cycle illustrating each sequential step with the labeled physiologic location needed to generate an efficient immune response against a tumor. Current immunotherapy approaches are shown at their point of augmentation (bolded) and biomaterial-based strategies used to benefit those immunotherapeutic strategies are listed below (bulleted points). MPs, microparticles; NPs, nanoparticles. Reproduced with permission from Cheung AS, Mooney DJ. Engineered materials for cancer immunotherapy. Nano Today August 1, 2015; 10(4):511e31, Elsevier.

setting, because the biomaterial could be implanted and release the drug continuously for extended times. Finally, although biomaterial approaches in immunotherapy seemingly make it more expensive, there are many opportunities for biomaterials to make these strategies cheaper. Biomaterials approaches allow the more efficient use of immunotherapeutic drugs by targeting them to key areas and cells types while minimizing physiologic clearance and drug degradation, and thus could minimize the necessary time patients need to spend in the clinic receiving the drugs. By providing a platform for immunologic modification of patient-derived cells in situ, biomaterial systems could also cut the costs of adoptive cell therapies. Thus, biomaterials have the potential to remove many of the challenges that immunotherapy currently faces.

NANOPARTICLE BIOMATERIALS FOR CANCER IMMUNOTHERAPY Introduction of Nanomedicine in Cancer Since its original discussion and conception in 1959 by Richard Feynman during his famous lecture, “There’s Plenty of Room at the Bottom” [38], nanotechnology has continued to show exciting potential in the field of cancer therapy. The National Nanotechnology Initiative defines nanotechnology as the manipulation of matter at sizes

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between 1 and 100 nm; however, many consider materials under 1 mm to warrant nanoparticle status. Because nanotechnology applications are so broad, the term “nanomedicine” further defines the application of nanotechnology in medicine. Most nanomedicine approaches rely on their intrinsic ability to accumulate in tumor tissue owing to a phenomenon known as the enhanced permeation and retention effect. Essentially, leaky tumor vasculature and inadequate lymphatic drainage result in the preferential localization of nanoparticles in the intratumoral space [39]. Thus, many groups have attempted to localize various therapeutic agents in tumors using nanoparticles to limit systemic toxicity and maximize drug efficacy [40]. In these attempts, a number of nanoparticle formulations have been widely optimized (i.e., metallic, liposome, dendrimer, and polymeric) (Fig. 41.2). Nanoparticles offer a number of unique properties that make them an attractive approach to improving current immunotherapy applications. The biocompatibility of the formulation material, surface charge, size, and shape of the nanoparticle all have critical roles in how it interacts with the body. In particular, these features greatly influence where it localizes in the body, with what cell types it interacts or by which it is internalized, and its cytotoxicity effects [41,42]. Thus, when designing nanoparticles for a specific application, these features must be taken into careful consideration. A more detailed description of these various nanoparticle design criteria will be provided subsequently.

FIGURE 41.2 Schematic of the four most common nanoparticle platforms (top) and their application as tumor drug delivery vehicles. Optimally designed particles with adequate shape, size, surface charge, hydrophilicity, and surface modifications will promote a longer circulating half-life and accumulation in the tumor through the enhanced permeation and retention (EPR) effect (bottom left). Surface modification of the particle with ligands or other targeting moieties can be used to target receptor on tumor cells and promote the intracellular delivery and release of drug payloads in response to specific intracellular stimuli (bottom right). Nps, nanoparticles. Reproduced with permission from Shao K, Singha S, Clemente-Casares X, Tsai S, Yang Y, Santamaria P. Nanoparticle-based immunotherapy for cancer. ACS Nano January 27, 2015;9(1):16e30, American Chemical Society.

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Effects of Nanoparticle Size and Shape With improvements in nanoparticle manufacturing techniques, it is possible to make a variety of nanoparticle shapes. This has led to a number of studies determining how cellular and physiologic systems interact with different nanoparticle shapes. Cellular particle uptake rates depend on both size and shape. It appears that for particle sizes larger than 100 nm, rod-shaped particles have the highest uptake rates, followed by spheres, cylinders, and cubes. However, with particles smaller than 100 nm, sphere-shaped particles surpass rods in terms of cellular uptake rates [43]. Other more exotic shapes with various surface characteristics are being investigated; however, it remains unclear whether these will provide differential uptake. When introducing nanoparticles to a physiologic system, several other effects such as pharmacokinetics and tissue distribution influence nanoparticle size and shape design criteria. For nanoparticles delivered intravenously, particles smaller than 6 nm are readily removed from the circulation by the kidneys. Nanoparticles larger than 200 nm typically accumulate in the liver and spleen and encounter a variety of immune cells tasked with removing large particulates from the circulation known as the mononuclear phagocytic system (MPS). Cells of the MPS typically phagocytose the nanoparticles, degrade them if possible, and then exocytose them for either filtration and removal or potential distribution in other physiologic tissues. Nanoparticle shape also is important in the physiologic systems, as evidenced by the fact that cells that circulate in the blood are typically not spherical. Thus, unsurprisingly, rod- or disk-shaped nanoparticles can promote longer circulation times compared with spherical nanoparticles [44]. Particle sizes between 10 and 200 nm are optimal in cancer nanomedicine for promoting long-term circulation, which ultimately provides more opportunities for delivery into tumor tissue. Within the tumor interstitium, particles in the upper end of this optimal range (100e200 nm) cannot extravasate far beyond the blood vessel and often remain stuck in the extracellular region between cells. Smaller particles (10 nm) have the potential to penetrate much farther into tumor tissue; however, without some form of targeting agent, they are not well-retained for more than 24 h [45]. Although many of these design optimization studies aimed for tumor drug delivery applications, in which the primary goal was to enhance nanoparticle localization within the tumor, an understanding of how particle size and shape dictate particle fate continues to shape the design characteristics of its newer applications in immunotherapy.

Effects of Nanoparticle Surface Charge and Hydrophilicity Nanoparticle surface charge and hydrophilicity are also features that have major implications in the design of nanoparticles. These features are easily modifiable through a number of technologies such as layer-by-layer, conjugation of surface moieties, and careful selection of particle material. Generally, positively charged particles are taken up by cells at a much higher rate than are neutral or negatively charged particles. Many researchers postulate that this is the result of the slightly negative surface charge of cellular membranes, which provides an electrostatic interaction that draws positively charged particles to the surface of the cell. In addition to having higher uptake, positively charged particles induce higher cytotoxicity effects because of detrimental disturbances that they cause in cellular membranes [46]. Similar to size, however, introduction into a biologic environment creates a number of additional considerations. Most notable is that the nanoparticle surface is rapidly covered by a variety of serum proteins, forming what is known as the corona on the particle surface. The surface charge and hydrophobicity of the particle largely dictate the composition of this corona and thus influence the future fate of the particle. Most highly charged particles, positive or negative, rapidly and strongly bind a number of serum proteins that tag them for removal through the MPS system [47,48]. Particles containing a more hydrophobic surface experience a similar fate, with high serum protein binding and removal. Thus, neutral particles with highly hydrophilic surfaces are the most optimal for naturally promoting long-term circulation; however, depending on the application of the nanoparticle, this may not be the desired surface features. With these understandings, many groups have developed ways to alter the surface characteristics of their nanoparticles to provide beneficial delivery features.

Effects of Nanoparticle Surface Functionalization As previously described, the surface characteristics of nanoparticles are a major parameter in the development and design of nanomedicine applications. Thus, numerous surface modification techniques have been developed to create favorable surface features without having to modify the bulk material of the nanoparticle. Surface attachment of hydrophilic polymers such as polyethylene glycol (i.e., PEGylation) is one of the most widely studied

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nanoparticle functionalization techniques. Many groups have shown significant improvements in the particle circulation time as a result of decreased protein absorption and MPS recognition after PEGylation [49]. Another surface modification commonly used is the attachment of targeting agents. A variety of targeting moieties have been developed, the most prominent of which are variable antibody fragments, peptides, receptor ligands, and aptamers. These are often made specific to a receptor or surface target on the target cell; upon binding, they promote the internalization of the nanoparticle carrier. The small size of nanoparticles provides a vast improvement in the total surface area exposure compared with many micromaterials and macromaterials, which allows for drastic advantages for targeting applications. Numerous groups have further shown that there are a variety of parameters that must be optimized for targeting ligands to make them maximally effective; these typically rely on targeting the agent density on the nanoparticle surface and the nanoparticle size [50]. With beneficial surface characteristics, nanoparticles can be considerably improved in nanomedicine applications. These important findings have promoted many of the design criteria for applying nanoparticles in immunotherapy.

Nanoparticle Targeting Applications in Immunotherapy Nanoparticle targeting often occurs through two primary methods, active or passive. Active targeting requires conjugating targeting moieties to the surface of the nanoparticle to encourage localization in specific areas or uptake by critical cells. Alternatively, passive targeting relies on the natural properties of the nanoparticle (i.e., surface charge, size, shape) to promote preferential uptake. Drug targeting is a unique ability of nanoparticles that has been widely studied in the field of drug delivery, with extensive success including a number of FDA-approved nanoparticle drug targeting formulations [51]. Compared with standard cancer nanomedicine targeting, immunotherapeutic applications require slightly more consideration of the intracellular or extracellular compartment in which the therapeutic payload will be released. For example, proinflammatory cytokines or other agents that target surface-bound cell receptors must be released into the extracellular tumor space. A number of strategies have been used to promote extracellular release, typically aiming to discourage phagocytosis of the particle. Notable examples include PEGylation or coating of the particle’s surface with antiphagocytic signals such as CD47 [52]. In addition, fine-tuning of the particle’s physical properties can be used to diminish cellular uptake, as previously discussed. In opposition to extracellular acting agents, delivery of other therapeutics such as specific Toll-like receptor (TLR) agonists, short interfering RNAs (siRNAs), or tumor-specific antigens for cancer vaccine applications require the delivered cargo to be released in specific intracellular compartments (i.e., endosome, lysosome, or cytosol). Enhancement of the particle surface with ligands targeting endocytosis-associated receptors and optimization of a particle’s physical characteristics can be used to promote more enhanced cellular uptake, especially in phagocytic cells [53]. In addition to targeting key compartments, particles must be designed to release their payload within a target environment. This can be achieved by chemically modifying the particle drug carrier to release its payload upon encountering some form of location-specific stimuli. The most common release cues include low pH and specific intracellular or extracellular proteases [54]. A number of triggered release systems have also been developed in which the release stimuli is delivered exogenously (i.e., light, magnetic field, heat) and promotes particle degradation and payload release [55,56]. Certain nanoparticle systems also provide the ability to fine-tune the rate of drug release by altering various particle features such as porosity, degradability, and the drug incorporation method (basic encapsulation versus stimuli-cleavable chemical conjugation). Thus, optimizing these various features has allowed for highly efficient, selective, and fine-tunable targeting of immunotherapeutic agents. This which will be discussed in more detail subsequently.

Nanoparticle Targeting of the Tumor Microenvironment Most solid tumor cancers harbor a wound healingelike inflammation and a highly immunosuppressive microenvironment. This feature of solid tumors is elicited by an intricate network of numerous cell types including the cancer cells themselves, a highly active stroma (i.e., cancer-associated fibroblasts), and a number of immune cells such as MDSCs, Tregs, and TAMs. Through a variety of mechanisms, this tumor microenvironment can drastically prevent effector T-cell infiltration into the tumor and abrogate the cytotoxic function of effector cells that manage to infiltrate it. The tumor microenvironment has even been shown to promote tumor growth and induce tumoral immune tolerance [3]. Thus, modulating this tumor microenvironment would provide a significant therapeutic opportunity, especially in situations in which a notable population of cancer-specific T cells exists but is unable

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overcome the immunosuppressive state of the tumor. The most common immunotherapeutic applications targeting the tumor microenvironment aim to inhibit or downregulate immunosuppressive features of the tumor, stimulate suppressed effector immune cells within the tumor parenchyma, or combinations of these approaches. Bolus delivery of immunomodulatory agents (i.e., cytokines, ICIs) have rapid saturation and/or degradation, which limits the effective dose delivered to the tumor and generates significant concern for systemic inflammatory toxicities. Thus, leveraging the advantages of nanoparticles to deliver these agents selectively to the tumor is an obvious application of nanomedicine in immunotherapy. Rather than attempt to discuss the robust amount of preclinical application of nanomedicine targeting the tumor microenvironment, we will provide a few examples of the diverse applications that are currently being explored. One interesting application involves targeting dysregulated genetic pathways within tumorinfiltrating immune cells. Signal transducer and activator transcription 3 (STAT3), a transcription factor that has previously been shown to promote Th2-based inflammation and enhance the survival of Tregs, is an intriguing target. Using tumor-targeting liposomes delivering intracellular agents to deplete STAT3 (i.e., small-molecule inhibitors [siRNA]), a few groups have shown impressive tumor immunosuppression reversal effects [57]. This strategy provides a valuable platform for the inhibition of numerous other dysregulated immunologic genetic pathways within the tumor. In addition to inhibiting immunosuppressive mechanisms directly, many groups have aimed to activate naive, suppressed, or tolerogenic effector cells within the tumor microenvironment. A number of strategies have been employed for this, including tumor localized delivery of proinflammatory and T cellestimulating cytokines (i.e., IFN-gamma, IL-1a, IL-2, IL-12, tumor necrosis factor-a) and APC-activating agents (i.e., CD40 ligand, cytosine-guanosine oligodeoxynucleotides [CpG], lipopolysaccharide, monophosphoryl lipid A [MPL-A], polyinosinicepolycytidylic acid [poly I:C]) [58e60]. Nearly all of these nanoparticle delivery systems showed benefits compared with bolus delivery of the same agents because they provided less systemic toxicity and extended release for up to many weeks in some cases, thus promoting more potent immunomodulatory effects. These strategies have shown relatively minimal therapeutic benefits as monotherapies; however, they are readily being investigated as combination modalities with more conventional immunotherapeutic strategies. Toward this combinatorial theme, a few groups have developed dual therapeutic delivery systems encapsulated within a single nanoparticle platform. One particular group used a polymeric liposome system, termed nanolipogels, to codeliver a small molecule inhibitor for transforming growth factor-b (TGF-b) (an immunosuppressive cytokine) and a potent T celleactivating cytokine, IL-2 (Fig. 41.3A). Their system was able to extend the release of these agents for many days and promoted significant survival benefits in preclinical models (Fig. 41.3B and C) [61]. In addition to releasing proinflammatory markers, some nanoparticle systems have been shown to promote the activation of target cell types by surface engagement with effector cells. One group optimized liposomes enhanced with IL-2 fragments and costimulatory 4-1BB, which were retained at the tumor site and promoted marked T-cell activation within the tumor through active engagement [62]. Applications of nanomedicine in targeting the tumor microenvironment are extensive; the example listed earlier represents only a portion of the vast work that has been done in this area. Overall, nanoparticles provide major advantages over the bolus delivery of immunomodulatory drugs. First, they enable the agent to be localized and concentrated at the tumor site, which benefits not only its efficacy but also its systemic toxicity. Second, they allow for extended immunomodulatory drug release, which provides sufficient time for effective immune responses to be generated and mounted against the tumor. Finally, they enable the targeting of drugs not only to important cells but also to necessary spatial and temporal locations. Overall, this renders these applications highly warranted and provides the rationale for the extensive past and current investigation of these systems.

Nanoparticle Targeting of Antigen Presenting Cells In addition to targeting the tumor directly, significant efforts have been made to target immunologic stimuli and antigenic information to key lymphoid cells or tissues using nanoparticles, primarily targeting APCs in the lymph node and spleen. DCs are commonly recognized as the most potent APCs, which makes them a logical cellular target of nanoparticle systems. By providing DCs with the appropriate “danger signals,” they have the ability to generate potent T-cell responses against antigens that they have acquired. This creates two unique applications for nanoparticles: one in which they could deliver only stimulus to DCs, in the hope that they have already acquired tumor antigens, and a second in which they could deliver both antigen and simulation to generate highly antigenspecific immune responses. Using nanoparticles, DCs can be targeted through a number of delivery routes. Because a subpopulation of DCs dwells in peripheral tissues, nanoparticles delivered into the dermal or subcutaneous space

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FIGURE 41.3 Nanolipogels (NLGs) are an effective immunotherapy codelivery platform (A) Synthesis illustration of NLGs. Lyophilized liposomes were loaded with a tumor growth factor-b receptor 1 inhibitor that was solubilized using methacrylate-conjugated b-cyclodextrins (SB), interleukin-2 (IL-2) cytokine, and a biodegradable cross-linking polymer. This allowed the formation of a core-shell structure with entrapped CDSB505 (blue) and IL-2 (green) in a biodegradable polymer matrix (red), all within a PEGylated liposome. (B) Percentage of intravenously delivered rhodamine (a fluorescent pseudodrug) remaining at 1, 24, 48, and 72 h postinjection. Rhodamine was either encapsulated in their NLGs or bolus delivered in saline. (C) Plot of tumor area following treatment of established B16eF10 melanoma in B6 mice. Treatment groups include no treatment or intratumoral injections of soluble SB, soluble SB and IL-12, NLG-encapsulated SB, NLG-encapsulated IL-2, and NLG-encapsulated SB and IL-2. Dates of injection are indicated by red arrows. PE, polyethylene; PEG, poly(ethylene glycol); PLA, poly(lactic acid). Reproduced with permission from Park J, Wrzesinski SH, Stern E, Look M, Criscione J, Ragheb R, et al. Combination delivery of TGF-b inhibitor and IL-2 by nanoscale liposomal polymeric gels enhances tumour immunotherapy. Nat Mater October 2012;11(10):895e905, Nature Publishing Group.

can be phagocytosed by DCs, activate them, and promote their migration to the lymph nodes, where they can induce T cell expansion. Alternatively, nanoparticles can be designed to drain into the lymphatics and encounter lymphoid tissueeresiding DCs. Both of these strategies provide a unique application for nanomedicine, because they essentially allow for DC priming and activation without the need for ex vivo cell isolation and manipulation. Thus, cancer vaccine applications using nanoparticles have promoted great excitement in the field and continue to show promising effects. A number of groups have developed nanoparticle formulations that can be targeted to the lymph node. Lymph nodes are obviously a useful target organ because of their primary function in the priming and generation of clonal T cell expansion. Various groups have demonstrated that 20- to 100-nm particles are optimal for lymph node targeting through lymphatic drainage, with markedly faster lymph node localization and uptake at the smaller sizes [63]. At this size, particles can extravasate from vascular beds and are driven into the lymphatic network, where

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they are transported to lymph nodes via interstitial pressure. Once in the lymph node, particles are readily taken up by DCs, providing an efficient route to promote immunologic priming. Various groups have then used this feature for subunit vaccine strategies, typically optimizing a model antigen ovalbumin (OVA). A study using subcutaneously injected PEGylated lipid nanoparticles with encapsulated OVA and cyclic dinucleotide stimulus promoted significant lymph nodal immune responses. Using their system, they noted a 15-fold improvement in adjuvant localization to the lymph node and a threefold improvement in OVA-specific CD8þ T-cell generation compared with bolus antigen and adjuvant delivery [64]. To improve on poor liposome serum stability, another group developed multilayered liposomes with intralayer cross-linking, termed interbilayer-crosslinked multilamellar vesicles (ICMVs). PEGylated ICMVs with OVA loaded into the particle center and MPL-A, a TLR-4 agonist, embedded within the ICMV wall promoted a 14-fold enhancement in antigen-specific T cells compared with bolus delivery [65]. Another group developed an alternative strategy using the natural trafficking phenomenon of serum albumin to traffic nanoparticles to the lymph node efficiently and selectively. Their most successful lipid formulation was composed of a lipophilic albumin-binding tail, a polar linking chain, and either a potent APC agonist or a target antigen peptide (Fig. 41.4A) [66]. When delivered together, this adjuvant and antigen nanoparticle formulation complexed with albumin and was highly localized to the lymph nodes, where they promoted significant activation and expansion of target antigen-specific CD8þ T cells (Fig. 41.4B and C). Impressively, when delivered in combination with a tumor-targeting antibody, modified IL-2 cytokine, and ICI therapies, the lymph nodeetargeting nanoparticles were also able to promote the rejection of large established tumors (Fig. 41.4D) [67]. Overall, these subunit vaccine nanoparticle strategies provide a promising method for generating antitumor T-cell responses. Despite the relative success of subunit vaccine strategies, they have a number of limitations. Most notable is that targeted vaccine strategies against a single antigen will not benefit cancers with unknown or unpredictable antigenic features. In addition, these strategies do not consider the vast patient-specific tumor antigen repertoire, which could ultimately promote immunoediting (i.e., selection of only the minimally immunogenic cancer cell populations) and eventual immunologic escape by the tumor. This has driven a number of groups to investigate the delivery of adjuvant nanoparticles to the tumor-draining lymph nodes (tdLNs). The tdLNs are an attractive target for simulation because they are continuously exposed to a plethora of cancer-specific antigens throughout the course of tumor progression. However, they also represent a relatively challenging stimulation site, with depleted effector cell populations and high levels of cellular suppression and exhaustion, especially at later stages of tumor progression [68]. Nevertheless, numerous nanomedicine attempts have been made to stimulate tdLNs, most often involving the delivery of APC stimulatory agonists in the hope of promoting clonal T cell expansion targeting numerous cancer antigens. One example includes the use of a 30-nm pyridyl disulfide (PDS) nanoparticle surface conjugated with CpG. The CpG conjugation was performed using pH-sensitive chemistry to promote release in the endolysosomal compartment of DCs for TLR-9 stimulation. PDSeCpG nanoparticles delivered intradermally near the tumor (B16eF10 melanoma) allowed for efficient draining and uptake by DCs in the ipsilateral lymph node, which the researchers verified to be tdLN by intratumoral injection of fluorescent dextran. These stimulatory particles increased the percentages of activated DCs and promoted a marked increase in CD8þ T cells specific for tyrosinase-related protein-2 (TRP2), a tumor-specific antigen known to be associated with this tumor model; they also significantly slowed tumor growth [69]. Although targeting of lymph nodes using an intradermal or subcutaneous injection of nanoparticles has shown promising immunologic effects, this approach raises a number of questions when considering applying them in the clinical setting. The most obvious question is which specific lymph nodes these systems should target, because the tdLN may contain the most tumor antigeneexperienced DCs, however, it also contains the highest levels of immunosuppression and exhaustion, which could hinder effective stimulation. This issue spurred the development of a systemically delivered nanoparticle vaccine strategy that could be delivered intravenously and would naturally accumulate in a variety of lymphoid tissues such as spleen, lymph nodes, and bone marrow. Using systemically delivered RNA liposome complex (RNA-LPX) nanoparticles, one group demonstrated an impressive ability to target lymphoid tissues and resident DCs and macrophages passively simply by altering the net surface charge of their particles (Fig. 41.5A and B). Even more impressive was the fact that systemic delivery of their optimally charged RNA-LPX could promote significant activation of DCs and strong expansion of antigen-specific T cells (Fig. 41.5C). This ultimately promoted considerable rejection and survival benefits in a variety of established murine tumor models (Fig. 41.5D) [70]. The application of nanoparticles as APC targeting agents is one of the most exciting developments in the field of nanomedicine. Numerous groups have shown the potent effects that APC-targeted nanoparticles can induce, especially in terms of generating tumor-specific T-cell expansion. The unique ability of small nanoparticles to traffic to

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FIGURE 41.4 Amphiphilic peptide and adjuvants efficiently target lymph nodes and promote antigen-specific immune responses. (A) Structure of amphiphilic peptide used to complex with serum albumin and target peptide antigens the lymph nodes. For treatment applications, this was codelivered with liposome cytosine-guanosine oligodeoxynucleotides (CpG). (B) IVIS fluorescence imaging showing both axillary and inguinal lymph node accumulation of various targeted molecular adjuvant formulations. CpG was conjugated with fluorescein amidite to make it fluorescent. Higher intensity indicates higher CpG localization to the lymph node. Groups tested include free CpG, CpG in incomplete Freund’s adjuvant (CpG in IFA), CpG encapsulated in a poly(ethylene glycol) (PEG)-coated liposome (CpG in liposome), or various amphiphile CpG conjugates; mono-acyl-conjugated CpG (C18-CpG), cholesterol-conjugated CpG (Cho-CpG), and diacyl lipid-conjugated CpG (lipo-CpG). Of note, lipo-CpG was found to associate strongly with albumin in serum (not shown), a likely mechanism of this delivery. (C) Plot of tumor area after vaccination treatment of established B16eF10 melanoma tumors in C57/BL6 mice. Mice were vaccinated with CpG and Trp2 peptide, an antigen known to be associated with B16eF10 tumors in bolus form or in designed amphiphile (amph) forms at the days indicated by black arrows. (D) Similar tumor area curves in B16eF10 melanoma tumors. AIPV trivalent-treated mice received vaccination with amphiphilic CpG and peptide antigens (Trp1, Trp2, and gp100) known to be associated with this tumor model, systemic administration of tumor targeting antibodies (targeting Trp1), systemic administration of albumin-fused interleukin-2, and administration of a-programmed cell death protein-1 immune checkpoint inhibitor. (AeC) Reproduced with permission from Liu H, Moynihan KD, Zheng Y, Szeto GL, Li AV, Huang B, et al. Structure-based programming of lymph-node targeting in molecular vaccines. Nature. March 27, 2014;507(7493):519e22, Nature Publishing Group. (D) Reproduced with permission from Moynihan KD, Opel CF, Szeto GL, Tzeng A, Zhu EF, Engreitz JM, et al. Eradication of large established tumors in mice by combination immunotherapy that engages innate and adaptive immune responses. Nat Med October 24, 2016;22(12):1402e10, Nature Publishing Group.

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FIGURE 41.5 RNA-lipoplex (RNA-LPX) lipid carriers promote lymphoid tissue targeting potent antigen-specific immune responses. (A) Particle size/polydispersity index (top) and zeta potential (bottom) for RNA-LPX achieved by using different lipid (DOTMA/DOPE) to RNA ratios. (B) Bioluminescent imaging of Balb/c mice after intravenous injection of various luciferase encoding lipideRNA ratios. A representative fluorescent image of a lung, liver, and spleen from each ratio group is shown below whole-mouse images with pie charts summarizing the distribution of signal in each. Fluorescent image to the far right shows additional uptake of RNA-LPX in lymph node and bone marrow at a 1.3:2 charge ratio. (C) Kinetics of ovalbumin (OVA) specific CD8þ T cells in the blood of C57/BL6 mice after multiple intravenous injections of OVA-encoding LPX carriers. Injection schedule is noted above the figure with black arrows indicating dates of OVA-LPX administration. (D) KaplaneMeier survival curves of C57/BL6 mice with established TC-1 tumors harboring E6 and E7 human papillomavirus oncoproteins as target antigens. Treatment groups include untreated (control) or intravenous administration of irrelevant OVA-LPX at day 7 (Irrelevant-LPX d7), E6/E7-LPX at day 7 (E6/E7RNA-LPXd7), or E6/E7-LPX administered at day 10 (E6/E7RNA-LPX d10). The treatment schedule can be seen above the figure, with yellow arrows indicating the administration dates for groups started on day 7 and red arrows for the group started on day 10. Reproduced with permission from Kranz LM, Diken M, Haas H, Kreiter S, Loquai C, Reuter KC, et al. Systemic RNA delivery to dendritic cells exploits antiviral defence for cancer immunotherapy. Nature 2016;534(7607):396e401, Nature Publishing Group.

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lymphoid tissues and readily be internalized by APCs, namely DCs, provides a unique application of nanomedicine in treating cancer. Furthermore, the improved efficacy and safety that nanoparticle targeted systems provide compared with the bolus delivery of immune stimulatory agents and/or antigens is unparalleled. Thus, many of these systems have translated to the clinical setting and likely will continue to do so, where it is hoped that they will provide immunotherapy applications with these pronounced benefits.

Conclusion of Nanomedicine Applications in Immunotherapy Unsurprisingly, nanomedicine applications have demonstrated amazing potential for the enhancement of cancer immunotherapies. This section focused largely on their targeting applications, such as targeting of the tumor microenvironment or targeting of APCs in important lymphoid sites. However, nanomedicine is being applied in many other applications in immunotherapy as well [71e73]. Most of these applications apply the extended release features of nanoparticle systems to the delivery of immunomodulatory agents. One such example involves the ex vivo conjugation or the in vivo targeting of T cells using nanoparticles that provide the extended release of T-cell stimuli or inhibitors of T-cell inactivation. These “T-cell piggyback” nanoparticle systems have shown amazing potential for prolonging the persistence of adoptively transferred T-cell therapies, a major hindrance for those therapies [74,75]. This demonstrates just one example of how nanoimmunosystems could theoretically be applied to target any immune cell population and deliver any number of immunomodulatory agents, which thus providing much future work for the field of nanomedicine immunotherapy. Overall, the applications of nanomedicine in immunotherapy will likely continue to exploit the unique ability of nanoparticles to prolong the release kinetics and selectively deliver immunotherapeutic agents to important cell types or physiologic locations. This simplistic-seeming ability has the potential to benefit immunotherapeutic strategies greatly in cancer, making them safer, more efficacious, and applicable to far more cancer settings than was previously thought achievable.

MACROSCALE BIOMATERIAL SCAFFOLDS FOR CANCER IMMUNOTHERAPY Implantable Biomaterial Scaffolds as Cancer Vaccines As we have discussed, the ability of nanoparticles to target bioactive factors toward important cell populations at specific physiologic locations is an example of how biomaterials have evolved over the past decades to be more than just inert, biologically compatible replacement parts. This paradigm shift toward the design of “smart,” biologically active materials that can manipulate and interact with cells in their environment is a fundamental concept familiar to those in the field of tissue engineering and regenerative medicine. For example, macroscale drug delivery biomaterial platforms (i.e., with at least one dimension large than 1 mm) are widely used to control the spatiotemporal delivery of multiple bioactive molecules and/or cells to direct cell behavior and drive functional tissue formation [76]. The ability of these engineered scaffolds to perform in situ cell programming has made them an attractive technology to improve DC and T cell function in the context of therapeutic cancer vaccines. In contrast to nanoparticulate vaccines, which target antigen and/or adjuvant components to DCs located in tumor-draining lymph nodes, the paradigm for using microscale and macroscale cancer vaccines is to recruit large numbers of immature DCs to the scaffold itself, where in situ programming takes place in a controlled microenvironment. This has the potential to reduce the financial and regulatory burdens associated with conventional methods of DC-based cancer vaccinations such as Sipuleucel-T, which require cell isolation, ex vivo cell culture and manipulation, and multiple patient procedures [77]. With the in situ DC vaccination approach, the implanted biomaterial scaffold is designed for controlled release of a recruitment factor that promotes the trafficking of immune cells to the implantation site. Once there, recruited cells such as immature DCs infiltrate the scaffold and are simultaneously presented with tumor antigen and proinflammatory “danger signals” in the form of a pattern recognition receptor ligand adjuvant. The mature, antigen-loaded DCs that are generated then traffic out of the scaffold toward draining lymph nodes, where they can facilitate anticancer immunity through T cell priming and activation (Fig. 41.6). This approach was first described by Ali et al. [77], who used macroporous scaffolds composed of the copolymer poly(lactide-co-glycolide) (PLG) as subcutaneously implanted DC cancer vaccines (Fig. 41.7). Incorporated into the PLG matrix were GM-CSF as a recruitment factor for immune cells, cationic nanoparticles of CpG as the danger signal, and melanoma tumor lysate as the antigen. Sustained release of bioactive GM-CSF was

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FIGURE 41.6 Schematic of an implantable biomaterial system that can be used for in situ dendritic cell programming. Recruitment factors released from the implant promote the trafficking of immature dendritic cells to the scaffold, where they encounter programming factors such as adjuvant and tumor antigen. The activated, antigen-presenting dendritic cells then leave the scaffold and migrate to draining lymph nodes, where they trigger downstream events such as the activation of antigen-specific cytotoxic lymphocytes. Reproduced with permission from Huebsch N, Mooney DJ. Inspiration and application in the evolution of biomaterials. Nature November 26, 2009;462(7272):426e432, Nature Publishing Group.

achieved over 15 days in vivo, resulting in a significant number of DCs infiltrating the scaffold (on the order of 106 cells, similar to the number of cells administered by ex vivo protocols). Scaffold-infiltrating DCs were shown to be activated by the CpG immobilized within the scaffolds and then subsequently dispersed to draining lymph nodes. After vaccination, cytotoxic lymphocytes targeting the melanoma-associated antigen TRP2 were markedly expanded in the spleen, and a threefold increase in CD8þ T-cell infiltration into tumors was noted for vaccinated mice compared with controls. Further work to characterize the multiple DC subtypes recruited to the implanted vaccine revealed the accumulation of a heterogeneous DC network including CD8þ DCs (important for crosspresentation of tumor antigen to Th1 T cells and cytotoxic lymphocytes) and plasmacytoid DCs (known to generate

Low- and high-power scanning electron microscopy cross-sectional images of a macroporous poly(lactide-co-glycolide) cancer vaccine scaffold containing granulocyte macrophageecolony-stimulating factor, cytosine-guanosine oligodeoxynucleotides, and tumor antigen. The image on the left (low power) is 1X and the image on the right (high-power) is 125X.

FIGURE 41.7

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type I IFNs, helping to activate CD8þ DCs) [78]. Prophylactic vaccination resulted in a 90% survival rate, whereas mice bearing established melanoma tumors and treated twice with the PLG vaccine resulted in 47% survival (Fig. 41.8). Subsequent work investigated the use of alternative adjuvants in the PLG vaccine including MPL-A and poly I:C [79], and alternative recruitment factors such as chemokine (CeC motif) ligand 20 (CCL20) and Fms-related tyrosine kinase 3 ligand [80], which illustrates the ability of this implantable, polymeric vaccine system for use as a versatile macroscale biomaterial platform to deliver multiple immunomodulatory factors. The PLGbased cancer vaccine is undergoing a phase I clinical trial under the name “WDVAX” (NCT01753089), investigating its safety in metastatic melanoma patients.

Injectable Biomaterial Systems as Cancer Vaccines Drawbacks of macroporous scaffolds composed of PLG are their relative stiffness and brittleness, which necessitate a small surgical procedure to enable subcutaneous implantation. To address this shortcoming, macroporous Cryogel scaffolds have been developed that can undergo large amounts of reversible deformation and rapid volumetric recovery, enabling the minimally invasive delivery of these scaffolds through small-bore needles [81]. Bencherif et al. [81] investigated an alginate-based Cryogel system as an injectable cancer vaccine scaffold, incorporating GM-CSF as the DC recruitment factor, CpG as an adjuvant, and irradiated B16eF10 melanoma cells as the antigen source (Fig. 41.9) [82]. To enhance the retention of the irradiated cells within the pores of the sponge-like vaccine scaffold, arginylglyclaspartic acid (RGD) peptides were covalently coupled to the alginate, allowing the preloaded cancer cells to adhere via integrin binding and remain localized to the scaffold microenvironment. Akin to the strategy used by the PLG-based cancer vaccine, this macroporous Cryogel-based cancer vaccine was designed to provide a localized immunogenic niche for recruited DCs and irradiated tumor cells to interact in the presence of the CpG danger signal without the tolerogenic factors present in the tumor microenvironment. Successful encapsulation of GM-CSF and CpG into the sponge-like alginate cancer vaccine construct was followed by sustained release of these factors over 1 month, enabling the recruitment and activation of both CD8þ and plasmacytoid DCs and a strong effector T cell response. Mice bearing B16eF10 melanoma tumors and treated with the Cryogel vaccine at days 3 and 10 after tumor inoculation displayed a 40% survival rate compared with unimmunized mice, illustrating the potent antitumor response that can be generated with this system. Another injectable, materials-based cancer vaccine platform was described by Kim et al. [83] and is based on higheaspect ratio mesoporous silica rods (MSRs). Using the same paradigm of host immune cell recruitment and

FIGURE 41.8 KaplaneMeier survival curves of mice vaccinated with poly(lactide-co-glycolide) (PLG) vaccines. (A) PLG vaccine implanted 14 days before B16eF10 melanoma tumor challenge, comparing groups receiving either blank PLG scaffold or PLG scaffolds loaded with tumor lysate, granulocyte macrophage (GM)ecolony-stimulating factor (GM-CSF), and 1, 10, 50, or 100 mg of cytosine-guanosine oligodeoxynucleotides (CpG). Mice receiving vaccines with 100 mg of CpG had a survival rate of 90%. (B) PLG vaccine implanted 9 days after B16eF10 melanoma tumor challenge, comparing groups receiving either blank PLG scaffold or PLG scaffolds loaded with tumor lysate, GM-CSF, and 100 mg CpG administered either once (on day 9, labeled Vax, 1) or twice (days 9 and 19, labeled Vax, 2). Reproduced with permission from Ali OA, Emerich D, Dranoff G, Mooney DJ. In situ regulation of DC subsets and T cells mediates tumor regression in mice. Sci Transl Med. November 25, 2009;1(8):8ra19, The American Association for the Advancement of Science.

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FIGURE 41.9 (A) Schematic showing how alginate-derived Cryogel cancer vaccines are prepared. Cytosine-guanosine oligodeoxynucleotides (CpG), granulocyte macrophageecolony-stimulating factor (GM-CSF), and arginylglyclaspartic acid (RGD)-containing methacrylated-alginate undergo cryogelation at subzero temperatures. Gels are then seeded with irradiated melanoma cells and incubated for 6 h before subcutaneous injection. (B) Scanning electron microscopic (SEM) image of the macroporous Cryogel. (C) Cross-sectional SEM image of macroporous Cryogel showing interconnected pores. (D) 2D confocal micrograph showing seeded melanoma cells on the Cryogel. Intracellular actin is shown in green, cell nuclei in blue, and Cryogel polymer walls in red. (E) 3D reconstruction of confocal image showing spreading and elongation of seeded irradiated melanoma cells after 6 h in culture. ODN, oligodeoxynucleotides. Reproduced with permission from Bencherif SA, Warren Sands R, Ali OA, Li WA, Lewin SA, Braschler TM, et al. Injectable cryogel-based whole-cell cancer vaccines. Nat Commun August 12, 2015;6(May):7556, Nature Publishing Group.

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modulation in vivo, a suspension of MSRs loaded with GM-CSF, CpG, and antigen was injected subcutaneously, leading to spontaneous nonspecific self-assembly in situ to form a 3D macroporous structure akin to “thrown matchsticks” (Fig. 41.10) [83]. The high pore volume and large surface area associated with the MSRs allowed for the continuous release of bioactive GM-CSF and CpG over several weeks in vitro and promoted the persistence of a model antigen (OVA) at the vaccination site compared with bolus control by an order of magnitude (Fig. 41.11) [83]. The MSR vaccine generated a potent adaptive immune response, not only increasing the frequency of antigen-specific cytotoxic lymphocytes but also eliciting strong and durable titers of antigen-specific serum immunoglobulin G2a (IgG2a) antibodies associated with a Th1 inflammatory immune response. By comparison, the conventional adjuvant alum combined with the model antigen ovalbumin was able to elicit only serum anti-ovalbumin IgG1 antibodies, skewing toward a Th2 more tolerogenic immune response. A single injection with the MSR-based cancer vaccine in the prophylactic setting was able to reduce tumor growth rate significantly and enhance survival over unvaccinated controls in a murine model of lymphoma; this illustrates the potential utility of MSRs for use as an injectable biomaterial scaffold in cancer immunotherapy.

Implantable Biomaterial Scaffolds to Enhance Autologous T Cell Therapy Macroscale biomaterials have also been used to augment other steps in the cancer-immunity cycle [84]. Whereas the work of Ali and coworkers focused on designing systems for the in situ programming of DCs in the context of cancer vaccines, others used a biomaterials approach to address downstream events in the cancer-immunity cycle. Adoptive T cell therapy, for example, involves the infusion of autologous tumor-reactive lymphocytes to target malignancies. Maintaining the presence and antitumor potency of the bolus-injected T cells has been challenging, however. Furthermore, infused lymphocytes have difficulty tracking to the site of the tumor and overcoming the immunosuppressive tumor microenvironment when they reach it, which results in limited efficacy of this modality against solid malignancies [85]. To improve the in vivo expansion and potency of lymphocytes used for adoptive cell therapy, Stephan et al. [86] described the use of a macroporous alginate scaffold to deliver T cells to accessible or resected tumors. The scaffolds were covalently modified with a lymphocyte adhesion peptide and stimulatory cues were presented to the loaded T cells by incorporating lipid-coated silica microparticles containing the soluble factor IL-15 superagonist and bound to costimulatory anti-CD3, anti-CD28, and anti-CD137 antibodies (Fig. 41.12) [86]. Adhesion peptide-modified scaffolds increased the transit of loaded lymphocytes by 6.3-fold in vitro versus unmodified scaffolds whereas the addition of the stimulatory silica microparticles boosted in vitro T cell proliferation by 22-fold. In a mouse 4T1 breast tumor resection model, tumor-reactive T cells were delivered to the resection

FIGURE 41.10 Schematic of mesoporous silica rod (MSR)-based vaccine in situ self-assembly and mechanism of action. MSRs dispersed in

phosphate-buffered saline (PBS) are injected subcutaneously, forming a pocket. As PBS diffuses away from the site of injection, MSRs randomly and spontaneously self-assemble into a macroporous structure, allowing for host immune cell recruitment (via release of recruitment factors from the MSRs), programming (via antigen and danger signals incorporated into the vaccine), and finally emigration of programmed cells to target structures such as draining lymph nodes. Reproduced with permission from Kim J, Li WA, Choi Y, Lewin SA, Verbeke CS, Dranoff G, et al. Injectable, spontaneously assembling, inorganic scaffolds modulate immune cells in vivo and increase vaccine efficacy. Nat Biotechnol January 2015;33(1):64e72, Nature Publishing Group.

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Recruitment factor (granulocyte macrophageecolony-stimulating factor [GM-CSF]) and danger signal (cytosine-guanosine oligodeoxynucleotides [CpG]) are released from mesoporous silica rod (MSR) scaffolds in a sustained manner. (A) In vitro release curve of bioactive GM-CSF from MSRs. (B) In vitro release curve of CpG from MSRs. (C) MSRs allow for enhanced persistence of a model antigen at the vaccination site. Left panel shows near-infrared fluorescent images of mice injected with a bolus of fluorescently labeled ovalbumin (OVA*) versus MSRs loaded with the same antigen; right graph shows relative OVA remaining at injection sites over time. Reproduced with permission from Kim J, Li WA, Choi Y, Lewin SA, Verbeke CS, Dranoff G, et al. Injectable, spontaneously assembling, inorganic scaffolds modulate immune cells in vivo and increase vaccine efficacy. Nat Biotechnol. January 2015;33(1):64e72, Nature Publishing Group.

FIGURE 41.11

Schematic of a macroporous alginate scaffold used to deliver antitumor T cells (blue) to the tumor site via surgical placement. Stimulatory silica microparticles (green) contained in the scaffold promote T-cell activation and expansion, followed by migration of the activated immune cells into the surrounding environment. Reproduced with permission from Stephan SB, Taber AM, Jileaeva I, Pegues EP, Sentman CL, Stephan MT. Biopolymer implants enhance the efficacy of adoptive T-cell therapy. Nat Biotechnol January 2015;33(1):97e101, Nature Publishing Group.

FIGURE 41.12

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FIGURE 41.13 Tumor-reactive T cells delivered to the tumor site using the macroporous alginate scaffoldestimulatory microparticle system

can eliminate residual disease in a 4T1 breast tumor model of residual disease. (A) Bioluminescent imaging of luciferase-expressing 4T1 breast tumors treated with no T cells, intravenously administered T cells, a T cell bolus injected into the tumor resection cavity (both unstimulated and prestimulated before injection), and biomaterial scaffoldedelivered T cells. (B) KaplaneMeier survival curves for these groups, showing 100% survival for mice treated with the biomaterial scaffoldedeployed T cells. All other groups showed tumor recurrence at varying time points. Reproduced with permission from Stephan SB, Taber AM, Jileaeva I., Pegues EP, Sentman CL, Stephan MT. Biopolymer implants enhance the efficacy of adoptive T-cell therapy. Nat Biotechnol. January 2015;33(1):97e101, Nature Publishing Group.

cavity using this macroporous scaffold system loaded with stimulatory microparticles, resulting in no recurrences. In contrast, intravenous injection of T cells had a metastatic relapse rate equal to negative controls, whereas injecting prestimulated T cells directly into the resection cavity had only a modest effect (improving median survival versus controls by an additional 11 days, but not preventing recurrence) (Fig. 41.13) [86]. T cells delivered to the resection cavity via the biomaterial scaffold were found to proliferate at the resection site (167-fold versus injected, prestimulated cells delivered without the scaffold) and maintain a nonexhausted phenotype, whereas intravenously injected T cells accumulated in the spleen and liver instead. In addition, the use of the polymeric scaffold to deliver NKG2D CAR-transduced T cells in a mouse model of stage 3 ovarian carcinoma with peritoneal metastases produced complete tumor clearance in 6 of 10 mice whereas locally injected T cells were unable to show tumor eradication in any animals. The efficacy of this system in preclinical models of improperly resected or metastatic cancer illustrates the potential for a biomaterials approach to provide localized delivery of antitumor immune cells while enabling their proliferation and activation in the face of an adverse tumor microenvironment.

CONCLUSION The unprecedented scale and scope of ongoing investments in immunotherapy by pharmaceutical companies illustrate the excitement regarding this area [87]. Despite this heavy investment, thus far single-modality immunotherapies such as ICIs have proven effective in only a small fraction of cancer patients. Although rational combinations of immunotherapies that affect multiple points in the cancer-immunity cycle have been shown to be more efficacious, our understanding of cancer immunobiology is still far from complete. Nonetheless, rapid progress witnessed in the field of cancer immunotherapy justifies confidence that increasingly effective immunotherapies can be designed as the components required for a robust antitumor response become increasingly clear. The ability of engineered biomaterial systems to control the spatiotemporal distribution of cells and bioactive factors will allow them to have an important key role in learning how biomaterialeimmune system interactions work, but then also to exploit these characteristics to shape an evolving anticancer immune response. It remains to be seen whether the most successful biomaterials-based cancer immunotherapies will be used to improve current therapies such as targeting immune-modulating antibodies and/or adoptively transferred immune cells more effectively to the tumor, enhancing the generation of antigen-specific immune effector cells through better priming, or curtailing the adverse tumor microenvironment. Perhaps an as yeteundiscovered pathway may

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present another opportunity for engineered systems to address. Regardless, the continued development of safer and more effective immunotherapies will depend on the effective integration of ideas across a broad spectrum of disciplines and collaboration between scientists and clinicians to allow for the effective translation of these technologies to the clinic.

List of Acronyms and Abbreviations APCs antigen presenting cells CAR T-cell Chimeric Antigen Receptor T-cell CCL20 Chemokine (C-C motif) ligand 20 CpG cytosine-guanosine oligodeoxynucleotides CTLA-4 Cytotoxic T lymphocyte-associated antigen-4 DCs Dendritic cells DOPE 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine DOTMA 1,2-di-O-octadecenyl-3-trimethylammonium propane EPR effect Enhanced permeation and retention effect FDA Food and Drug Administration GM-CSF Granulocyte colony stimulating factor HER2 Human estrogen receptor 2 ICIs Immune Checkpoint Inhibitors ICMVs Interbilayer-crosslinked multilamellar vesicles IFN-gamma Interferon gamma IL Interleukin IL-12 Interleukin 12 IL-1a Interleukin 1a IL-2 Interleukin 2 LPS Lipopolysaccharide MDSCs Myeloid-derived suppressor cells MPL-A Monophosphoryl lipid A MPS Mononuclear phagocytic system MSRs mesoporous silica rods NK cells Natural killer cells NLGs Nanolipogels OVA ovalbumin PD-1 Programmed cell death protein-1 PDS Pyridyl disulfide PLG Poly(lactide-co-glycolide) Poly I:C Polyinosinicepolycytidylic acid PRR pattern recognition receptor RGD Arginylglyclaspartic acid RNA-LPX RNA liposome complexes (lipoplex) siRNA Short interfering RNAs STAT3 Signal transducer and activator transcription 3 TAMs Tumor-associated macrophages tdLNs Tumor-draining lymph nodes Th1/Th2 T-helper 1/T-helper 2 adaptive immune response TLR Toll-like receptor Tregs T-regulatory Cells TRP2 Tyrosinase-related protein-2

Glossary Active nanoparticle targeting Nanoparticle targeting applications that require the surface conjugation of a targeting moiety (i.e., ligands, antibody) to target nanoparticles to a desired cell or location. Adaptive immune system The response of antigen-specific lymphocytes to antigen, including the development of immunological memory. Adaptive immune responses are generated by the clonal selection of lymphocytes. Adaptive immune responses are distinct from innate and nonadaptive phases of immunity, which are not mediated by clonal selection of antigen-specific lymphocytes. Adjuvant An adjuvant is any substance that enhances the immune response to an antigen with which it is mixed. Adoptive T-cell therapy A specific immunotherapy strategy which involves the isolation and ex vivo expansion of tumor-specific T cells, and then reinfusion to the patient. Antigen presenting cells Highly specialized cells that can process antigens and display their peptide fragments on the cell surface together with molecules required for T-cell activation. The main antigen-presenting cells for T cells are dendritic cells, macrophages, and B cells. Autoimmune responses An adaptive immune response directed at self-antigens is called an autoimmune response; likewise, adaptive immunity specific for self-antigens is called autoimmunity

GLOSSARY

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B Cell One of the two major types of lymphocyte. The antigen receptor on B lymphocytes, usually called the B-cell receptor, is a cell-surface immunoglobulin. After activation by antigen, B cells differentiate into cells producing antibody molecules of the same antigen specificity as this receptor. Cancer vaccine A specific immunotherapy strategy that aims to train the immune function to recognize and destroy cancer cells. This is typically done by providing both an antigen expressed by the cancer cells and an adjuvant to stimulate generation of T cells specific to that cancer antigen. Chemokines Chemokines are small chemoattractant proteins that stimulate the migration and activation of cells, especially phagocytic cells and lymphocytes. They have a central role in inflammatory responses. Clonal T cell expansion Clonal expansion is the proliferation of antigen-specific lymphocytes in response to antigenic stimulation; it precedes their differentiation into effector cells. It is an essential step in adaptive immunity, allowing rare antigen-specific cells to increase in number so that they can effectively combat the pathogen that elicited the response. Corona A serum protein shell that forms around a nanoparticle, especially charged nanoparticles, after introduction into a protein-rich environment such as blood. Costimulatory molecules The proliferation of lymphocytes requires both antigen binding and the receipt of a costimulatory signal. Costimulatory signals are delivered to T cells by the costimulatory molecules B7.1 and B7.2, related molecules that are expressed on the surface of antigen presenting cells, and which bind the T cell surface molecule CD28. Cytokines Cytokines are proteins made by cells that affect the behavior of other cells. Cytokines made by lymphocytes are often called lymphokines or interleukins, but the generic term “cytokine” is used in this book and most of the literature. Cytokines act on specific cytokine receptors on the cells that they affect. Cytotoxic T cells or lymphocyte Tcells that can kill other cells are called cytotoxic Tcells. Most cytotoxic Tcells are major histocompatibility complex class Ierestricted CD8 Tcells, but CD4 Tcells can also kill in some cases. Cytotoxic Tcells are important in host defense against cytosolic pathogens. Dendritic cells Dendritic cells, also known as interdigitating reticular cells, are found in T-cell areas of lymphoid tissues. They have a branched or dendritic morphology and are the most potent stimulators of T-cell responses. Nonlymphoid tissues also contain dendritic cells, but these are not able to stimulate T-cell responses until they are activated and migrate to lymphoid tissues. The dendritic cell derives from bone marrow precursors. It is distinct from the follicular dendritic cell that presents antigen to B cells. Effector immune cells Armed effector cells, most often effector T cell, that can be triggered to perform their effector functions immediately upon contact with cells bearing the peptideemajor histocompatibility complex for which they are specific. They contrast with memory T cells, which need to be activated by antigen-presenting cells to differentiate into effector T cells before they can mediate effector responses. Enhanced permeation and retention effect A phenomenon by which molecules of a certain size, typically nanoparticles and/or drugs, tend to accumulate in tumor tissue more so than they do in normal tissue. Many believe that this is a result of leaky tumor vasculature and inadequate lymphatic drainage of solid tumors. Helper Tcell A class of T cells, a specific subclass of CD4þ T cells, that can help B cells make antibody in response to antigenic challenge. The most efficient helper T cells are also known as Th2 cells, which make the cytokines interleukins-4 and 5. Hydrophilicity/hydrophobicity The physical property of a material to either repel or promote water from binding to its surface. Immunoediting A dynamic process by which tumors survive attack by the immune system. Typically described as three phases: elimination of immunologically susceptible cells, equilibrium, and finally, immunologic escape. Immunogenic Any molecule that can elicit an adaptive immune response on injection into a person or animal is called an immunogen, and thus is classified as being immunogenic. In practice, only proteins are fully immunogenic because only proteins can be recognized by T lymphocytes. Immunologic escape The point at which a tumor is no longer susceptible to immune surveillance and begins to progress in terms of growth and malignancy. Immunologic memory When an antigen is encountered more than once, the adaptive immune response to each subsequent encounter is speedier and more effective, a crucial feature of protective immunity known as immunological memory. Immunological memory is specific for a particular antigen and is long-lived. Immunosuppressive A characteristic of something that promotes the inhibition or downregulation of immune responses. Immunotherapy The prevention or treatment of disease by using or stimulating components of the immune system. Innate immune system Cells that are responsible for the early phases of the host response to an injury or immunologic insult in which a variety of innate resistance mechanisms recognize and respond to the presence of a pathogen. Innate immunity is present in all individuals at all times, does not increase with repeated exposure to a given pathogen, and discriminates between a group of related pathogens. Interferons Cytokines that can induce cells to resist viral replication. Interferon-alpha and interferon-beta are produced by leukocytes and fibroblasts, respectively, as well as by other cells, whereas interferon-gamma is a product of CD4 Th1 cells, CD8 T cells, and natural killer cells. IFNgamma has the activation of macrophages as its primary action. Interleukins A generic term for cytokines produced by leukocytes. The more general term “cytokine” is often used, but the term “interleukin” is used to name specific cytokines such as interleukin-2. Macrophages Macrophages are large mononuclear phagocytic cells important in innate immunity, in early nonadaptive phases of host defense, as antigen-presenting cells, and as effector cells in humoral and cell-mediated immunity. They are migratory cells deriving from bone marrow precursors and are found in most tissues of the body. They have a crucial role in host defense. Monoclonal antibodies Antibodies made from the same clonal cell line and have monovalent affinity, meaning that they bind the same epitope (the part of the antigen that is recognized by the antibody). Mononuclear phagocytic system A network of phagocytic cells located in the reticular connective tissue that are highly responsible for the clearance of nanosized and microsized materials from the blood. Myeloid-derived suppressor cells A heterogeneous group of immature myeloid immune cells that have been shown to induce significant immunosuppressive effects, especially in pathogenic situations such as in chronic infections or cancer. Nanomedicine The applications of nanotechnology in medicine. Nanotechnology The manipulation and engineering of materials that are in the nanometer-size range, typically less than 100 nm. Natural killer cells Large granular, non-T, non-B lymphocytes that kill certain tumor cells. Natural killer cells are important in innate immunity to viruses and other intracellular pathogens, as well as in antibody-dependent cell-mediated cytotoxicity.

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Passive nanoparticle targeting Nanoparticle targeting applications that use the nature physiochemical properties (i.e., size, shape, and surface charge) to target nanoparticles to a desired cell or location. Pattern recognition receptors Receptors that bind to pathogen-associated molecular patterns, which are typical of bacteria or many commercially used adjuvants. PEGylation The process of attachment of poly(ethylene glycol) to molecules, nanoparticles, or macrostructures. Phagocytosis/endocytosis The internalization of particulate material by cells. If the material is being taken up by a phagocytic cell via invagination of the cellular membrane to form a phagosome, it is referred to as phagocytosis. If the material is taken up by a nonphagocytic cell through membrane invagination to form a vacuole, it is referred to as endocytosis. Plasmacytoid dendritic cells A class of dendritic cells that circulate in the blood and are found in peripheral lymphoid organs (i.e., spleen, lymph node). RNA-lipoplex nanoparticles A complex of RNA and lipid material to form liposome-like nanoparticle. Self-antigens By convention, natural antigens in the body of an individual are called self-antigens. Lymphocytes are screened during their immature stages for reactivity with self-antigens; those that respond undergo apoptosis. Spatiotemporal delivery Delivery of drug to its optimal location and with optimal time kinetics. Subunit vaccine A vaccination strategy that uses only part of the disease-causing pathogen. Cancer subunit vaccines typically involve the stimulation of an immune response against a single antigen known to be associated with the cancer cells. T-helper 1 (Th1)/T-helper 2 (Th2) adaptive immune response Th1 immune responses are adaptive immune responses that are primarily driven by Th1 cells. Th1 cells are a subset of CD4 T cells that are characterized by the cytokines they produce. They are mainly involved in activating macrophages and are sometimes called inflammatory CD4 T cells. Th2 immune responses are adaptive immune responses typically driven by Th2 cells. Th2 cells are a subset of CD4 T cells that are characterized by the cytokines they produce. They are mainly involved in stimulating B cells to produce antibody and are often called helper CD4 T cells. Therapeutic cancer vaccine Stimulation of the immune system so as to allow for the targeted attack of antigens present on cancer cells, typically through by generating tumor antigenespecific T cells. Toll-like receptor (TLR) agonists Any agonist aiming to target a TLR. TLRs are a common pattern recognition receptor on antigen presenting cells and other innate immune cells. Binding and activation of TLRs promote inflammatory effects and stimulation of the cell. There are numerous TLRs, each with a different activation potential. They are all named TLR, followed by a number, as in TLR-4. T-regulatory cells Regulatory or suppressor class of T cells that can inhibit T-cell response. Tumor interstitium The space within a tumor that is not taken up by cancer cells; essentially all of the remaining space within a tumor. Tumor microenvironment An intricate network of cancer cells, tumor stroma, and immune cells within a solid tumor that generally promote an overall immunosuppressive state. Tumor stroma The fibroblast, vasculature endothelial cells, pericytes, and other structural proteins that make up the extracellular matrix component of a tumor. Tumor-associated macrophages Macrophages found in close proximity or within solid tumors. Tumor-associated macrophages have been shown to induce significant protumor and antitumor effects and induce major immunosuppressive effects. Tumor-specific antigens An antigen that is overexpressed or specifically located on or within tumors. Often this is an antigen that is unique to the tumor, thus providing a targetable feature.

Acknowledgments The authors would like to acknowledge Aurelie Hanoteau for her reviewing assistance. JMN acknowledges financial support from the National Institute of General Medical Sciences (T32GM088129) and the National Institute of Dental and Craniofacial Research (F31DE026682) both of the National Institutes of Health. AGS acknowledges support from the Federal Drug Administration (R01 #FD-R-05109-01) and Baylor College of Medicine Carolyn Weiss Law Fund for Translational Research. SY gratefully acknowledges support from the National Institutes of Health (R00 DE023577) and the University of Texas System (Rising STARs Award). This content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institutes of Health.

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[79] Ali OA, Verbeke C, Johnson C, Sands RW, Lewin SA, White D, et al. Identification of immune factors regulating antitumor immunity using polymeric vaccines with multiple adjuvants. Cancer Res March 15, 2014;74(6):1670e81. [80] Ali OA, Tayalia P, Shvartsman D, Lewin S, Mooney DJ. Inflammatory cytokines presented from polymer matrices differentially generate and activate DCs in situ. Adv Funct Mater August 1, 2013;23(36):4621e8. [81] Bencherif SA, Sands RW, Bhatta D, Arany P, Verbeke CS, Edwards DA, et al. Injectable preformed scaffolds with shape-memory properties. Proc Natl Acad Sci USA November 27, 2012;109(48):19590e5. [82] Bencherif SA, Warren Sands R, Ali OA, Li WA, Lewin SA, Braschler TM, et al. Injectable cryogel-based whole-cell cancer vaccines. Nat Commun August 12, 2015;6(May):7556. [83] Kim J, Li WA, Choi Y, Lewin SA, Verbeke CS, Dranoff G, et al. Injectable, spontaneously assembling, inorganic scaffolds modulate immune cells in vivo and increase vaccine efficacy. Nat Biotechnol January 2015;33(1):64e72. [84] Chen DS, Mellman I. Oncology meets immunology: the cancer-immunity cycle. Immunity July 25, 2013;39(1):1e10. [85] Redeker A, Arens R. Improving adoptive T cell therapy: the particular Role of T Cell costimulation, cytokines, and post-transfer vaccination. Front Immunol 2016;7(September):345. [86] Stephan SB, Taber AM, Jileaeva I, Pegues EP, Sentman CL, Stephan MT. Biopolymer implants enhance the efficacy of adoptive T-cell therapy. Nat Biotechnol January 2015;33(1):97e101. [87] Brawley L. With 20 agents, 803 trials, and 166,736 patient slots, is pharma investing too heavily in PD-1 drug development? Cancer Lett 2016; 42(37):1e8. [88] Huebsch N, Mooney DJ. Inspiration and application in the evolution of biomaterials. Nature November 26, 2009;462(7272):426e32.

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C H A P T E R

42 Gene Editing in Regenerative Medicine Yunlan Fang1, Xuguang Chen2,*, W.T. Godbey3 1

XenoBiotic Laboratories, Inc., Plainsboro Township, NJ, United States; 2Salubris Biotherapeutics, Inc., Gaithersburg, MD, United States; 3Tulane University, New Orleans, LA, United States

GENOME EDITING TOOLS Genome editing can be defined as the modification of the genome within a living cell through the insertion, deletion, or replacement of one or more segments of DNA. Currently, the commonly accepted view of genome editing entails genome modifications via site-specific nucleases. Although the current view is valued, it is also important to include other genome editing approaches that do not use targeted nucleases, because many downstream applications in regenerative medicine require much higher efficiency editing rates than site-specific editing can supply. These approaches include gene insertion, recombination, and translocation.

Targetable Nucleases All genome editing approaches that use targeted nucleases share a similar mechanism. The function of targeted nucleases is to induce a double-strand break within the cellular genome at a particular site. After the cleavage is established, inherent cellular DNA repair mechanisms are initiated to fix the double-strand break spontaneously. Three major types of targetable nucleases have been developed and widely applied within the genomic engineering field: clustered regularly interspaced short palindromic repeats (CRISPR) and CRISPR-associated systems (CRISPR-Cas), transcription activator-like effector nucleases (TALENs), and zinc-finger nucleases (ZFNs). Among these three systems, CRISPR has attracted considerable attention owing in part to its simplicity, although great work has also been performed using TALENs and ZFNs. Because the three systems share the same DNA repair mechanisms inherent in the cell, therapeutic concepts investigated using one system could be pursued via the other two if certain conditions are met. Clustered Regularly Interspaced Short Palindromic Repeats Genes that can be targeted by a CRISPR system are CRISPR-Cas genes. This system, which was first noticed (but not immediately pursued) in Escherichia coli by Ishino et al. in 1987 [1], involves repeated sequences of DNA that do not appear in tandem. As time went on and more repeat sequences were identified and characterized in different strains of archaea, bacteria, and mitochondria [2], the repeat sequences became a focus of interest, with the term “CRISPR” being coined in 2002 by Jansen and Mojica and the involvement of Cas proteins being observed at the same time [3]. CRISPR-Cas systems are thought to be parts of a primordial immune system in which bacteria acquire immunity to certain viruses by inserting a fragment of the viral genome into a CRISPR locus. The bacterial immune process basically consists of chopping up the genome of a viral invader into small fragments of approximately 20 bases. Focusing on a single fragment, it will be placed into a special location in the bacterial genome. The locus, now modified with the viral gene fragment, will be transcribed and processed into CRISPR-RNA (crRNA), which we will refer to as guide RNA (gRNA) when discussing applied genome editing. * Author was at Rutgers University when the chapter was written.

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Cas proteins are restriction enzymes (endonucleases) that have a tunable recognition sequence. The enzyme specificity is determined by the crRNA (gRNA) introduced earlier. The Cas protein binds to a region on the 30 side of the crRNA (gRNA) known as the “scaffold sequence,” leaving a section at the 50 end of the gRNA, termed the “spacer sequence,” free to act as a probe for locating homologous regions within DNA (Fig. 42.1). In wild-type bacterial systems, the crRNA is used to identify DNA as foreign; in engineered systems, the gRNA is used to locate a region in the genome for editing in a site-specific way. After binding of the Cas-bound crRNA or gRNA spacer sequence to a homologous region of DNA, the Cas protein will cleave the DNA in one of several ways. Three different paradigms for gene editing with CRISPR will be discussed subsequently: double-stranded breaks, nickases, and homologydirected repair. Knockouts via Double-Strand Breaks Isolated cells can be made to express specific gRNA sequences that can be used to target other DNA sequences for knockout. There are three requirements for the transcribed gRNA: (1) It must have a scaffold sequence that will bind to a given Cas. Cas is inactive unless gRNA is bound. (2) The target sequence must be unique; otherwise, several genes may be bound and produce off-target effects. (3) The target must be just upstream of a protospaceradjacent motif, also known as a PAM sequence. (For discussion purposes, we will focus on a specific Cas: Cas 9 from Streptococcus pyogenes, also known as spCas9. The CRISPR/spCas9 system is one of the most wellcharacterized and widely used gene editing systems.) The PAM sequence for spCas9 is 50 -NGG-30 . The PAM sequence is an absolute requirement for this system: If the spCas9/gRNA binds to a target sequence that is not next to a PAM, there will be no cleavage by spCas9. However, with the PAM present, cleavage will take place three to four bases upstream of the PAM (Fig. 42.2). SpCas9/gRNA can bind to any portion of the DNA with a PAM, but it is the matching with the spacer sequence that determines whether DNA cleavage will take place. The spCas9 protein has six domains, including two endonuclease domains: HNH and RuvC. If there is a good match between the gRNA spacer sequence and the target DNA sequence, a conformational change in the Cas9 occurs, thereby activating HNH (which cleaves the target DNA strand) and RuvC (which cleaves the nontarget strand) domains. A double-strand break is thereby introduced. The double-strand break will be repaired by the cell in one of two ways: nonhomologous end joining (NHEJ) or homology-directed repair (HDR). NHEJ is the more common of the two pathways in the cell. It is also more errorprone because insertions and deletions can easily occur, and these can cause frameshift mutations. The result of a frameshift can cause codons to be translated differently, as in different amino acids being assembled onto a growing

FIGURE 42.1 The DNA recognition and cleavage functions carried out in the clustered regularly interspaced short palindromic repeats (CRISPR) gene editing systems come from a complex consisting of a CRISPR-associated systems (Cas) protein interacting with a CRISPR RNA (crRNA). The crRNA, known as guide RNA in applied gene editing, consists of a scaffold sequence that will bind to the Cas protein, and a spacer sequence that will extend from the complex, allowing it to interact with chromosomal DNA as a probe. The wild-type Cas protein has domains that are responsible for inducing double-stranded breaks in DNA after homology with the spacer sequence has been established.

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FIGURE 42.2 Interaction of CRISPR-associated systems 9e guide RNA (Cas9/gRNA) with double-stranded DNA. Although a homologous DNA sequence may be recognized by a Cas9/gRNA complex, DNA cleavage will only take place if a protospacer-adjacent motif (PAM) sequence is present in the DNA. It is the PAM sequences, which appear in numerous, nearly evenly spaced copies in the DNA, that are referred to by the CRISPR acronym (Clustered Regularly Interspaced Short Palindromic Repeats).

polypeptide, the elimination of a stop codon, or the introduction of a premature stop. In each of these cases, a nonfunctional protein is likely, so this type of DNA repair often results in knocking out a gene by lowering or halting the production of a functional protein product. HDR has a much higher fidelity, but it is not used as much in the cell. We will return to HDR in a moment. Nickases One problem with generating knockouts via double-strand breaks is that despite using a 20-base recognition (spacer) segment on the gRNA, sequences with near-homology may also be cut. One logical solution to this problem would be to use gRNAs with longer spacer sequences. The problem with this idea is that the CRISPR-Cas system is set up for approximately 20-nucleotide spacers, so using a gRNA that has a 40-nucleotide spacer sequence would not work well. An alternative approach is to modify the Cas protein so that it cuts only one DNA strand. By creating a Cas9 with an inactivated HNH domain, or a Cas9 with an inactivated RuvC domain, and by loading the mutated Cas9 protein with respective gRNAs that are specific to separate areas on one of the two strands of DNA, one can essentially require homology at 40 nucleotides (20 for each of the two gRNAs) before the formation of a double-strand break (Fig. 42.3). Only one type of mutated Cas9 is required: HNH-mutated or RuvC-mutated. Unmutated HNH cuts the strand that pairs with the guide RNA, while unmutated RuvC cuts the strand with the PAM sequence. It is through the selection of guide sequences with spacers complimentary to sequences on each of the host DNA strands that is used to generate what amounts to a staggered double-strand break. The single-stranded breaks introduced by the mutated Cas proteins are known as “nicks”; hence the term “nickases” used to describe the proteins. Nicks are usually repaired quickly by the HDR system and with high fidelity. With the Cas nickase system, a double-nick, in which both strands are cut, will only occur if both of the spacer sequences line up with good homology.

Nickases working simultaneously to generate single-stranded breaks surrounding a targeted region of DNA. The requirement of homology matching in two distinct guide RNA spacer sequences improves specificity over wild-type clustered regularly interspaced short palindromic repeats (CRISPR)-associated systems (Cas), resulting in a reduction in off-target cleavage. PAM, protospacer-adjacent motif.

FIGURE 42.3

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Homology-Directed Repair HDR is how CRISPR is used for specific gene edits, such as altering the identity of a single base. Three components are requisite to this system: a Cas protein such as spCas9 or Cas9 nickases, gRNA, and a repair template with homology arms. The mechanism used for this precise form of gene editing is similar to what has already been covered, except that DNA (linear or plasmid) containing the desired insertion or repair is codelivered with the Cas9/gRNA. Consider a linear repair template designed to correct a point mutation in the genome. Although only one base pair (bp) will differ from the genomic sequence, the length of the repair template must be longer than one bp. The bases to the left and right of the repair bp, called the “homology arms,” will have the same sequence as the genomic DNA. The length of each depends on the size of the insert (Fig. 42.4).

FIGURE 42.4 Homology-directed repair. In this scheme of genome editing, clustered regularly interspaced short palindromic repeats

(CRISPR)-associated systems (Cas)eguide RNA complexes (Cas9 shown here) are codelivered with a repair template. The template will be introduced into the genome via homologous recombination after the introduction of a double-stranded break. PAM, protospacer-adjacent motif.

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It is important for the repair template not to contain a PAM sequence; otherwise it would be cut by the Cas proteins. If the section to be inserted happens to contain a PAM sequence, the insertion sequence should be reengineered with a silent mutation to remove the PAM. For example, an AGG sequence, which codes for arginine and is also is a Cas9 PAM sequence, could be modified to AGA, which still codes for an arginine but is no longer a PAM sequence. Recall that the HDR, while having high fidelity, is not used by the cell as much as NHEJ, which means NHEJ will still be at work. Because somatic mammalian cells have two alleles, the result of this type of gene editing for a given cell will be a combination of: • wild-type, • NHEJ-mutated, and/or • HDR-inserted (repaired) alleles. It is therefore imperative for the researcher to perform a screen to isolate HDReHDR clones. SpCas 9 Variants and Orthologues Although the SpCas9 system has been widely accepted for laboratory applications worldwide, some limitations still must be overcome. For example, SpCas9 depends on the specific (S. pyogenes) PAM sequence 50 eNGGe30 , and the cleavage site must be close to this PAM. Although NGG is abundant throughout the mammalian genome, the appearance of the sequence might not be close enough to the desired target to effect the desired modification. To address this issue, multiple novel Cas9 variants that use different PAM sequences have been engineered or discovered. Benjamin et al. developed a series of Cas9 variants by mutating the particular site of SpCas9 (D1135/R1335/T1337 or D1135/G1218/R1335/T1337) or screening various Cas9 orthologues from nature [4,5]. The PAM library has been expanded to include 50 -NGCG-30 (SpCas9 VRER variant), 50 -NGAG-30 (SpCas9 EQR variant), 50 NGAN-30 (SpCas9 VQR variant), 50 -NNAGAAW-30 (Cas9 from Streptococcus thermophilus [St1Cas9]), and 50 -NNGRRT-30 (Cas9 from Staphylococcus aureus [SaCas9]) [4]. These Cas9 variants and orthologues provide great convenience to scientists by expanding the editable areas within the genome. Off-target effects pose a potential risk to future therapeutic applications of CRISPR [6,7]. These occur when the nuclease induces a double-strand break at an unwanted site, leading to unwanted cellular changes such as transformation. To reduce the possibility of off-target effects, SpCas9 has been mutated by rational design with apparent success. One such set of mutations, designed to reduce potential interactions between SpCas9 and the phosphate backbone of DNA through direct hydrogen bonds, produced what is referred to as the S. pyogenes Cas9 High-Fidelity variant 1 (SpCas9-HF1). (The specific mutations to the wild-type SpCas9 were N497 A, R661 A, Q695 A, and Q926 A [8].) SpCas9-HF has editing activity that is comparable to wild-type SpCas9, but off-target events have been nearly eliminated. Another set of mutations designed to reduce the affinity between a positively charged groove (located between the HNH, RuvC, and PAM-interacting domains) and the negative charges of genomic DNA have served to neutralize positively charged residues in the Cas9 (K810 A, K848 A, K1003 A, and R1060 A) [9]. The mutated enzyme was named “enhanced specificity” SpCas9 (eSpCas9), and like SpCas9-HF1, it has been shown to have activity similar to the wild-type enzyme but with a clear reduction in off-target events. A novel CRISPR system from the genera Prevotella and Francisella, called Cpf1, has introduced several specific features to these systems [10]. Usually, the double-strand breaks introduced by a Cas protein are blunt ended. The Cpf1 system is different in that it creates overhangs of four to five bases [10]. The overhangs present an opportunity to increase the efficiency of gene insertion. In addition, the Cpf1-associated system does not require the transactivating crRNA to perform the DNA interference function [10]. Cpf1 is able to process pre-crRNA into crRNA without relying on host RNase, which makes it possible to edit multiple genes simultaneously. Because of the large size of a single gRNA expression cassette (w400e500 bp), for SpCas9, each single gRNA must be placed in different vectors or driven by a different set of promoters [11,12]. However, thanks to the feature of Cpf1 being guided by a small (39-nucleotide) crRNA, a single promoter within one vector is enough to drive many crRNA together. Published experiments have achieved simultaneous modification of four different targets by using Cpf1 with a single construct [13] (Fig. 42.5). Another interesting system employs CRISPR-C2c2, which comes from the bacterium Leptotrichia shahii [14]. Instead of working on DNA, CRISPR-C2c2 can serve as an RNA-guided RNase that can cleave certain singlestranded RNAs via pairing with a 28-nucleotide sequence in the crRNA. This system could be used to downregulate gene expression via posttranscriptional knockdown, similar to RNA-induced silencing complexemediated RNA

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FIGURE 42.5 Schematic outlining the components of a multiplex gene preeclustered regularly interspaced short palindromic repeats RNA array. The outline represents a single construct consisting of guide sequences for four genes and Cpf1. Two promoters are included: the U6 promoter directs transcription of the four guide sequences and the cytomegalovirus promoter drives transcription of the Cpf1 nuclease gene. Figure adopted from Zetsche B, Heidenreich M, Mohanraju P, Fedorova I, Kneppers J, DeGennaro EM, et al. Multiplex gene editing by CRISPR-Cpf1 using a single crRNA array. Nat Biotechnol 2016.

interference [14]. Moreover, fluorescent labeling of C2c2 makes it possible to tag specific messenger RNA (mRNA) sequences and linking dC2c2 with a splicing factor makes it possible to affect mRNA splicing [15]. However, the function of the system has been shown only in bacteria. Further investigation is necessary before it can be used reliably in mammalian cells. Transcription Activator-like Effector Nucleases TALENs are chimeric proteins which contain two functional domains: a DNA-recognition transcription activatorlike effector (TALE) and a DNA nuclease domain. They work for gene editing by recognizing a specific sequence, which the user can design, and introducing a double-stranded break with an overhang. TALENs therefore serve as a form of customizable restriction enzyme. The DNA-recognition TALEs used in TALEN technology are secreted by the genus Xanthomona, a type of phytopathogenic bacteria [16]. A TALE consists multiple domains, including a translocation domain at the N terminus, a nuclear localization signal and a transcription-activation domain that are close to C-terminus, and multiple repeats in the center that serve to recognize DNA molecules [16]. The repeat sequences are composed of a highly conserved 33e to 34eamino acid sequence, with positions 12 and 13 being variable. The variable portion of the sequence is termed the “repeat variable di-residue” (RVD). By changing the RVD sequence of a specific repeat, that particular repeat can be made to bind a specific nucleotide [17]. As examples, a monomer with an RVD consisting of asparagine-asparagine will bind with a purine, asparagine-isoleucine will bind with adenine, asparagine-glycine will bind with thymine, and histidine-aspartate will bind with cytosine (Table 42.1) [17]. By manipulating the RVDs of its repeated sequences, a TALE can be made to serve as a probe for a specific site within the genome. The nuclease portion of a TALEN is the catalytically active domain of the restriction enzyme FokI. (The DNA recognition domain has been removed.) FokI can be used in mammalian cells to cut genomic DNA, but it must be dimerized to be functional. Both homodimeric and heterodimeric domains can be used [18]. One challenge in using TALENs is that the enzyme construct is hard to build owing to the numerous repeat sequences that must be customized for targeting. This is why much effort has been directed toward improving the construction of TALENs. One method, called Golden Gate assembly, is widely accepted for its relative simplicity [19]. This method employs type IIS restriction endonucleases, which cleave outside their recognition sequences. By selecting appropriate type IIs endonucleases, the recognition sequence can be removed from the fragment of interest, thus allowing for digestion and ligation in the same reaction mixture. Similarly, because the site where TABLE 42.1 The Identity of the Variable Sequences (Repeat Variable Di-residues [RVDs]) in a Transcription Activator-like Effector Determine the Nucleotide Targets RVD Sequence

Nucleotide Bound

Asparagine-asparagine

Purines

Asparagine-isoleucine

Adenine

Asparagine-glycine

Thymine

Histidine-aspartate

Cytosine

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the DNA cleavage itself occurs is not sequence-specific (is outside the recognition sequence), the compositions of the resulting overhangs are not dictated by the enzyme, so multiple, specific overhangs can be created simultaneously, allowing for the creation of multiple assemblies. Compared with the CRISPR system, which requires only the design of a single gRNA, the TALEN system requires significantly greater effort devoted to sequence design and cloning. As a result, CRISPR is the more widely accepted method for genome editing. TALENs have certain advantages over CRISPR in specific situations, though. For example, TALENs can theoretically target any site within the genome, as opposed to CRISPR, which requires the presence of a PAM site. Therefore, when no PAM site is available for a given application, one can still turn to a TALEN system. In addition, TALENs are associated with lower off-target effects versus wild-type SpCas9. For perspective, a study of off-target cleavage by CRISPR-Cas9 (using SpCas9) versus TALENs was performed using six different genes. The TALENs produced no off-target cleavage, whereas half of the applications (three of six) using SpCas9 nucleases yielded off-target events [20]. Published comparisons of TALENs with the high-fidelity forms of CRISPR were not readily available at the time this chapter was written. Zinc Finger Nucleases ZFNs are another type of DNA recognition/cleavage construct that acts like a restriction enzyme. They combine the Cys2-His2 zinc finger motif as the DNA recognition domain with the FokI DNA cleavage domain [21]. The Cys2His2 zinc finger motif consists of a single zinc atom with about 30 amino acids in a b sheeteb sheetea helix conformation [22]. Each zinc finger contacts three DNA bases through amino acids on the a helix of the motif [23]. The selectivity between a zinc finger and target triplets can be modulated by adjusting the amino acid sequence. A library of zinc-finger domains that can target nearly all of the possible triple nucleotide combinations has been developed [24]. By connecting three to six zinc finger motifs via conserved linker sequences, 9e18 nucleotides will be targeted by a monomer [25]. As with TALENs, because ZNFs use FokI as the nuclease, they are also designed for use as dimers, so a total of 18e36 nucleotides will be targeted. The primary hurdle for using ZFNs is that of enzyme design and construction. The cloning process used to produce the correct amino acid sequence is labor intensive, but the correct primary sequence is also imperative for achieving high targeting specificity. As a result, the number of studies applying this technique has dwindled since TALENs and CRISPR were introduced to the field. However, unlike the TALENs and CRISPR systems, ZFNs originated from mammalian cells, so there has been less concern regarding immunogenicity. Phase I clinical trials that infused ZFN-modified autologous CD4 T cells as a treatment for HIV have been completed, with the treatments being deemed safe (NCT00842634) [26]. This implies that ZFNs may be suitable for other clinical applications.

Other Genome Manipulation Tools Transposons and Transposase There are two types of transposons: class I transposons, also known as retrotransposons, which are transcribed and reverse-transcribed as part of the translocation mechanism, and class II transposons, which do not require transcriptionereverse-transcription but use transposases to catalyze the translocation mechanism. The transposons used for genomic engineering are class II. Combined with the enzyme transposase, a transposon can be moved efficiently between plasmid and genome in a cut-and-paste fashion. Transposase acts by recognizing a sequence of inverted terminal repeats (ITRs) that appear at each end of the transposon and introducing a pair of doublestranded breaks in the DNA. A gene flanked by ITRs can be cut out of a plasmid and transferred to a site within the genome that is flanked by the same ITRs. This method can also be used to delete genes that appear between two ITRs sites in the genome. There is no size limit for transposons. This system has been used, for example, to reprogram primary cells into induced pluripotent stem cells (iPSC) through virus exclusion [27]. Recombinase The major function of recombinase is to exchange strands between two DNA segments that have partial sequence homology. The most widely used recombinase systems for genome manipulation are Cre-loxP and Flp-FRT. The Cre and Flp enzymes are site-specific recombinases that recognize 34-bp recognition sites. The loxP recognition site contains two 13-bp palindromic sequences and an 8-bp spacer (50 -ATAACTTCGTATA-atgtatgc-TATACGAAGTTAT-30 , [spacer shown in lowercase]) [28]. Similar to loxP, FRTconsists of two 13-bp palindromes surrounding an 8-bp spacer (50 -GAAGTTCCTATTC-tctagaaa-GTATAGGAACTTC-30 ) [28]. The Cre and Flp recombinases can be used for gene

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insertion, excision, inversion, or translocation combined with their respective recognition sites and are often applied in genome engineering [29], such as for generating conditional mutants. The International Knockout Mouse Consortium has been launched to produce mutant murine embryonic stem cells for every gene in the mouse genome (more than 20,000 genes) by incorporating the loxP and FRT sites [30]. Using Cre and Flp, knockout mice for any gene can be produced from these cells. Integrase In the past, gene insertion typically was achieved by plasmid integration with stringent antibiotic selection or via infection with lentiviral vectors. Insertions occurred randomly and the copy number, location, and direction of insertion were difficult to control. Homologous recombination with or without a targetable nuclease addressed the problem of directionality, but the size of the inserted gene was limited and integration efficiency was still lacking, especially in primary cells [31]. A technology named “dual integrase cassette exchange” (DICE) provides an alternative route whereby the copy number, location, and direction gene integration are controlled, and with higher efficiency [32]. DICE adopts the phage integrases phiC31 and Bxb1, which have the ability to insert a gene into their own recognition sites by unidirectional recombination. However, these recognition sites do not exist in the mammalian cell genome. To get around this problem, a “landing pad” was introduced into a safe region of the mammalian genome by integrating those recognition sites through TALEN-assisted homologous recombination. As a result, desired genes could be efficiently implanted into the landing pad by codelivery of the two integrases and donor template. It is claimed that there is no size limitation for the inserted gene [32].

DELIVERY CARGO The genome-editing tools just described can be delivered into cells via multiple routes based on whether DNA, RNA, or protein is the tool being used.

DNA DNA is often delivered as a plasmid, which is typically constructed via standard molecular cloning techniques. When properly preserved, DNA has a long shelf life. The efficiency of DNA delivery into cells is the major limiting factor for the associated methods. DNA for genome editing can also be engineered into a viral genome, and plasmids can be adapted for use with viruses for delivery into cells [33]. Although viruses typically generate higher gene delivery efficiencies than do nonviral gene delivery agents, immunogenicity is a concern for in vivo applications, especially if repeated delivery events must be performed. Another concern regarding the delivery of DNA into cells is that the foreign DNA could integrate into the host genome through homologous recombination. If this occurs in an untargeted fashion, vital host genes may be disrupted or inactivated, tumor suppressor genes may be knocked out, or oncogenes might be activated. Whether targeted or untargeted, genomic integration could cause the encoded nuclease to be expressed in a sustained manner, potentially causing continuous formation of double-strand breaks and serious off-target effects.

RNA The editing methods that use delivered RNA avoid the possibility of genomic integration. RNA can be produced by cells transfected in vitro, harvested, and applied to cells of interest in subsequent experiments. With RNA delivery, the efficiency of enzyme production may be enhanced because genetic materials do not have to reach the nucleus and the transcription step is eliminated. However, RNA is more difficult to handle and store than is DNA because RNases are abundant in the cell, on the skin, and throughout most research environments. The immunogenicity of RNA produced in vitro should also be considered [34].

Proteins Instead of delivering genetic materials, delivering proteins into cells for gene editing is now a commercially available option. Multiple CRISPR Cas9 protein delivery kits exist on the market; TALENs and ZFNs can also be

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delivered into cells. For the CRISPR system, the ribonucleoprotein complexes are formed in vitro by combining a Cas nuclease with the appropriate gRNA, and the constructs are delivered into cells without fear of genomic integration. Another advantage to delivering proteins instead of (deoxy)ribonucleotides is that the amount of enzyme that will reside within the cell is more precisely controllable.

DELIVERY METHODS Microinjection, gene gun, electroporation, and hydrodynamic delivery are all methods that have been used to deliver genome editing tools into cells [35e38]. These methods are not widely used because they either need specialized equipment and experienced professionals (microinjection) or they are associated with cell or tissue damage (gene gun, electroporation, and hydrodynamic delivery). These methods are also restricted in their ability to reach host tissues in vivo. Chemical agents such as cationic lipids, polymers, and dendrimers have been widely used for gene delivery, but they are associated with gene delivery efficiencies that are generally lower than those of viral vectors. However, chemical delivery methods are still preferred compared with viral methods in some cases owing to their ease of preparation, lower immunogenicity, and lower expense [39]. All of the chemical gene delivery methods can be applied to genome editing because there is no chemical difference between plasmids encoding targetable nucleases and plasmids encoding other genes. However, different mechanisms may be involved for the delivery of proteins. Cationic liposomes have been used to this end with relative success. The delivery of Cas9egRNA complexes or TALENs via cationic lipids is evident by the multiple commercially available agents that are in the marketplace [40,41]. The cationic polymer poly(ethylenimine) has been used to coat self-assembled DNA cages to delivery Cas9/gRNA complexes successfully into mammalian cells [42]. Cell-penetrating peptides have been used to deliver targetable nuclease proteins, including Cas9 and TALEN [43,44]. Interestingly, ZFN has inherent cell-penetrating capabilities and can pass into cells without a carrier [45]. Viruses have been used for gene delivery for decades [46]. For gene editing applications, several viruses have been evaluated, including lentivirus [47], adenovirus [48], and adeno-associated virus (AAV) [49]. Lentiviral vectors have been shown to have a high transduction efficiency in multiple primary cell types and are capable of having genes integrated into the genome of the host [39,50]. As mentioned, incorporating a gene encoding a nuclease into the genome can lead to constitutive expression of the enzyme, which can lead to off-target events. Even worse, lentiviral integration occurs randomly, which can potentially induce various malignancies [51,52]. One solution to these problems of integration is to knock out the viral gene that encodes integrase, a technique that has been used to transfer genes encoding ZFNs, TALENs, and Cas9 [20,53]. Another solution is to develop lentivirus-derived particles to carry nuclease proteins. ZFN, TALEN, and Cas9 proteins have been successfully packed into such lentiviral particles [47,54]. Donor templates have been copacked into the particles to achieve targeted DNA insertion and gene correction [54]. It has been reported that this approach has lower off-target activity than traditional delivery methods [47,54]. Adenoviral vectors allow transient transgene expression in both dividing and nondividing cells. These vectors carry a much lower risk of genomic integration than their retroviral counterparts. A set of third-generation adenoviral vectors named “helper-dependent adenoviral vectors” have increased cargo sizes (36 kb, as opposed to w8 kbp) [55], which provides more versatility for the delivery of donor DNA templates for gene insertion and correction. However, immunogenicity is still a concern, as it is for all virus methods [56], especially when multiple gene delivery events must be carried out. Like adenovirus, AAV can produce transient transgene expression in both dividing and nondividing cells. In gene delivery, certain AAV serotypes are preferred because their delivered genes will be integrated into the host genome in a targeted fashion. Keeping in mind the previous discussion about problems with constitutively expressed nucleases, recombinant AAV have been produced that lack the viral rep gene to lower or eliminate the frequency of genomic integration [57]. Some strains of AAV carry lower immunogenicity concerns for a single administration, but repeated administrations require different serotypes to avoid a secondary immune response. These viruses are relatively small so they have limited DNA loading capacity (w5 kb) [58]. The ZFNs with double monomers [59] are usually compatible with this limitation, but it is difficult to package double TALEN monomers (w3 kb each) or SpCas9 (w4.2 kb) and associated gRNA within a single AAV vector [60]. For in vitro genome editing, this limitation can be overcome by packaging the functional units separately and cotransducing cells. Edited cells can then be screened and selected. However, this is not practical for in vivo editing because the selection step

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cannot be performed, and every abnormal cell is expected to be treated. One could reduce the size of the gene encoding the enzyme: a Cas9 orthologue, named S. aureus Cas9 (SaCas9), which is 1 kb smaller than SpCas9, was identified [60]. The smaller gene has been successfully copacked with gRNA in a single AAV construct, with precise genome editing still being attainable.

APPLICATIONS OF GENE EDITING IN REGENERATIVE MEDICINE Stem cells are characterized by two properties: self-renewal, in which they are able to divide numerous times while remaining undifferentiated, and pluripotency, which means they can differentiate down any of the three germ cell lineages (ectodermal, mesodermal, and endodermal). Because of these characteristics, stem cells are attractive progenitors to a broad range of regenerative medicine applications. Stem cells can be harvested and isolated from various tissues such as bone marrow [61], amniotic fluid [62], umbilical cord blood [63], or brain [64]. In addition, they can be induced from differentiated cells by introducing genes encoding specific transcription factors [65]. In contrast to embryonic stem cells, adult-derived stem cells and iPSCs can be harvested or produced autologously, thus eliminating the need for immunosuppression and providing a means of personalized medicine as new applications are developed. As genome-editing technologies develop, gene-edited stem cells may be produced for the repair and regeneration of various tissues and organs [66]. Various applications of genome editing to the regeneration of specific tissues or organs are presented in this section.

Liver The liver is a vital organ containing numerous enzymes that are involved in a wide range of bioprocesses such as glucose storage and gluconeogenesis, cholesterol synthesis, bile production, urea production, the synthesis of certain blood proteins, red blood cell turnover, and the detoxification of metabolites and drugs (such as ethanol). Because of the importance of this organ and the numerous diseases that can arise to affect its function, the liver is a prime target for regenerative medicine and gene editing therapy. a1-Antitrypsin (A1AT) deficiency is an autosomal recessive disorder that is caused by a single point mutation (Glu342Lys). Copies of the mutated protein aggregate in the endoplasmic reticulum of hepatocytes instead of being secreted into the blood and body fluids [67,68]. A1AT deficiency may further develop into cirrhosis of the liver and necessitate the need for liver transplantation [69]. ZFNs and PiggyBac transposons have been used in conjunction to correct the point mutation in the A1AT gene in human iPSCs derived from a patient with A1AT deficiency [69]. The cells were then differentiated into hepatocyte-like cells, which displayed normal A1AT function. This pioneering work demonstrated the potential for applying genome-editing and iPSC technologies for an autologous cellbased therapy. Several disorders are related to deficiencies in one or more of the enzymes of the urea cycle, a biochemical pathway that is involved in removing nitrogen from the body. Hyperammonemia is one such condition that can lead to encephalopathy and, potentially, death. Arginase, which is an enzyme of the urea cycle, has a part in ureagenesis. Defects in the gene encoding arginase can lead to hyperammonemia. Gene editing has been used with patient-derived iPSC to restore arginase activity in arginase-deficient patients [70]. In these experiments, CRISPRCas9 was used to add an arginase-1 cDNA expression cassette into the first exon of the hypoxanthine-guanine phosphoribosyltransferase locus in the isolated cells. The cells were then differentiated into hepatocyte-like cells, which displayed both arginase activity and ureagenesis [70]. Another potential cause of urea cycle disruption is ornithine transcarbamylase deficiency, caused by a mutation in the associated gene. Gene editing research toward addressing this malady has included a dual AAV system used to deliver Cas9-gRNA and donor DNA separately into newborn mice with a partial deficiency in ornithine transcarbamylase [71]. Results have been weak but encouraging, with correction of the mutated gene in about 10% of hepatocytes and improved survival in treated mice that were given high-protein diets. Tyrosinemia, which is characterized by a buildup of tyrosine (and some of its breakdown products, depending on the type), is a severe genetic disorder stemming from an inability to break down the amino acid tyrosine. There are three types of tyrosinemia, each characterized by the specific enzyme that is mutated in the degradation pathway (Fig. 42.6). Tyrosinemia type I is caused by a deficiency in the enzyme fumarylacetoacetate hydrolase (FAH), which catalyzes the conversion of fumarylacetoacetate into fumarate plus acetoacetate [72,73]. A deficiency

APPLICATIONS OF GENE EDITING IN REGENERATIVE MEDICINE

751

NH2 HO

CH2 CH

COOH

Tyrosine

TAT p-Hydroxyphenylpyruvic Acid

HPD

p-Hydroxyphenylacetic Acid

p-Hydroxyphenyllactic Acid

Homogentisic Acid

HGD H

O

O CH2 C

H

H

COOH Maleylacetoacetic Acid O

MAI O CH2 C

HOOC

CH2 COOH O HOOC

CH2 CH2

C

Succinylacetoacetic Acid

CH2 COOH

H

O HOOC CH2 CH2

Fumarylacetoacetic Acid

O CH2 C CH2 COOH

O

C CH2 C

CH3

Succinylacetone

FAH Fumaric Acid

Acetoacetic Acid

FIGURE 42.6 Partial degradation pathway of tyrosine, highlighting the steps associated with tyrosinemia type I, catalyzed by fumarylacetoacetate hydrolase (FAH). Tyrosinemia type II has a deficiency in the enzyme tyrosine aminotransferase (TAT) and type III tyrosinemia has a deficiency in the enzyme 4-hydroxyphenylpyruvate dioxygenase (HPD). Figure from Jorquera R, Tanguay RM. Cyclin B-dependent kinase and caspase1 activation precedes mitochondrial dysfunction in fumarylacetoacetate-induced apoptosis. FASEB J 1999;13(15):2284e98.

in FAH results in the accumulation of fumarylacetoacetate and maleylacetoacetate, which induce cellular damage [74]. CRISPR-Cas9 has been examined for use as a treatment for hereditary tyrosinemia type I, in which the gene encoding 4-hydroxyphenylpyruvate dioxygenase was deleted. This deletion caused hepatocytes associated with tyrosinemia type I to be converted into hepatocytes that would be associated with the more-benign tyrosinemia type III [75]. Rather than relying on genetic deletion, a different laboratory group used a gene editing approach to deliver mRNA encoding Cas9 via lipofection, along with AAV containing DNA encoding gRNA, to correct mutant fah in mice with human hereditary tyrosinemia. This method yielded modest recovery of FAH function in approximately 6% of hepatocytes [76].

Angiogenesis Angiogenesis is a process in which new capillaries grow from preexisting vessel networks to form more complexed vessel networks [77]. It is essential for the regeneration of tissues and organs if the component cells are to receive oxygen and nutrients and have carbon dioxide and other waste products removed. Therefore, procedures designed to enhance angiogenesis are related to regenerative medicine. Genome editing using TALENs has been used to induce cells to express hepatocyte growth factor (HGF) to induce angiogenesis [78]. The common theme of editing the genomes of progenitor cells was used on mesenchymal stromal cells isolated from human umbilical cord blood. In this case, a TetOn sequence preceded a gene encoding HGF pUC19-TetOn-HGF expression cassette, and the DNA was inserted into chromosome 19. The TetOn sequence was used as a safety measure. It allowed for control of the hgf gene by requiring that tetracycline be present for the gene to be transcribed. After the establishment of HGF expression in the cells, the cells were encapsulated in alginate and transplanted to correct a hind limb ischemia model. The entire procedure worked well, yielding angiogenesis in the in vivo model.

Muscle: Muscular Dystrophy Muscular dystrophy is family of diseases characterized by a gradual loss of muscular strength and control. There are nine forms of the disease, each with its own characteristic set of symptoms and time of onset. Some affect only

752 TABLE 42.2

42. GENE-EDITING IN REGENERATIVE MEDICINE

Types of Muscular Dystrophy and Some Defining Characteristics [79]

Type

Sexes Affected

Typical Time of First Appearance

Duchenne

Males

Ages 3e5

Upper legs and pelvis; heart and diaphragm in advanced cases

Becker

Males

Ages 11e25

Upper legs, pelvis; similar to, but milder than, Duchenne

Limb-girdle

Both

Teens to early adulthood

Begins in hips, spreads to shoulders, legs, and neck

Myotonic

Both

Ages 20e30

General: prolonged stiffening of muscles after use

Facioscapulohumeral

Both

Teens to forties

Face, shoulders, and back

Congenital

Both

Present at birth, symptoms by age 2

General weakness, progressing to muscle shortening

Oculopharyngeal

Both

Ages 40e60

Muscles of the eye, face, and throat

Distal

Both

Ages 40e60

Hands, forearms, feet, and lower legs

EmeryeDreifuss

Males

Ages 10 to mid-twenties

Shoulders, upper arms, and lower legs

Characteristic Muscles Affected

males whereas others affect either sex. The symptoms of some appear in infancy whereas others present in adolescence or adulthood. The areas affected are also varied and are characteristic of the form of the disease being experienced (e.g., limbs, eyes and throat, face and shoulders, hips, hands, or even heart and diaphragm). A brief description of the types of muscular dystrophy appears in Table 42.2. Duchenne Muscular Dystrophy Duchenne muscular dystrophy (DMD) is an X-linked genetic disease arising from mutations in the dystrophin gene, a large gene consisting of 2.4 million bp and containing 79 exons. Exons 45e55 hold the mutational hot spot for the gene. Frameshift mutations lead to dysfunction of the dystrophin protein, which has an important role in skeletal and cardiac muscle function. In the absence of functioning dystrophin, patients experience progressive muscle weakness and wasting. In conventional virus-mediated gene therapy, owing to the size limitations imposed by the viral capsid, the full-length DMD cDNA must be replaced by a truncated complementary DNA (cDNA) that will yield proteins with compromised function, yielding minimal efficacy and a T cell response after gene delivery in vivo. Exon dys 51 (from Sarepta Therapeutics, Inc.) was developed using an oligonucleotide exon-skipping strategy [80]; after a controversial argument regarding its efficacy [80], it was approved by the US Food and Drug administration in 2016 as the first drug to treat DMD [81]. One drawback of the treatment, however, is that patients must receive regular intravenous infusions. Genome editing provides the ability to correct mutations in genes without the need for introducing overly long sequences. Because of this, editing tools hold a different promise than do conventional gene delivery techniques for providing a more complete treatment for cells with mutations to the dystrophin gene. For instance, paired ZFNs have been delivered to DMD patient-derived myoblasts using electroporation, with the purpose of removing exon 51 of the dystrophin gene. This led to restoration of the dystrophin reading frame in approximately 13% of DMD patient mutations [82]. Long et al. [83], Nelson et al. [84], and Tabebordbar et al. [85] used CRISPRCas9 systems on the mdx mouse model of DMD to delete the mutated exon 23 from the dystrophin gene, resulting in the restoration of expression of functional dystrophin in the murine skeletal and cardiac muscles, plus enhancements in skeletal muscle function in the mice. Long et al. [83] delivered SpCas9 using AAV9, whereas Nelson et al. [84] and Tabebordbar et al. [85] delivered SaCas9 via AAV8 and AAV9, respectively. Iyombe-Engembe et al. used a pair of gRNAs to target exons surrounding exons 51e53 to bring about their deletion from the DMD gene, which restored the correct reading frame with 62% efficiency [86]. The CRISPR-Cas9 system has been delivered by “all-in-one” adenoviral vectors into patient-derived muscle progenitor cells for the purpose of removing mutations from the major DMD mutational hot spot [87]. This blanket approach has been asserted to counteract multiple mutations and provide treatment coverage in more than 60% of patients with DMD. Another investigation that targeted the DMD mutational hot spot deleted up to 725 kb using CRISPR-Cas9 with NHEJ. The efforts reportedly restored the function of dystrophin in cardiomyocytes and skeletal muscle cells derived from edited human iPSCs [88].

APPLICATIONS OF GENE EDITING IN REGENERATIVE MEDICINE

753

TALENs have been used to target a point mutation (A > G) associated with muscular dystrophies in iPSCs isolated from a canine golden retriever model of muscular dystrophy. The corrected canine progenitor cells were implanted into immunodeficient mice after ischemia-induced myocardial injury and cardiotoxin-induced hind limb skeletal muscle injury, yielding encouraging levels of regeneration in both cardiac and skeletal muscles [89]. Limb-Girdle Muscular Dystrophy Limb-girdle muscular dystrophy types 2B and 2D are muscular dystrophies that are induced by mutations in the dysferlin and a-sarcoglycan genes, respectively. The dysferlin nonsense mutation (c.5713C > T; p.R1905X) and the a-sarcoglycan missense mutation (c.229C > T; p.R77C) have been corrected using the CRISPR-Cas9 system in iPSCs isolated from patients with the respective diseases [90]. Another treatment aimed at the dysferlin mutation involved site-specific cDNA insertion via DICE, or TALENassisted homologous recombination for insertion precise, carried out in iPSCs isolated from patients with the disease. Expression of dysferlin was restored in the corrected iPSCs [90]. The typical next step would be to reintroduce the cells into the patient with the hope of restoring muscle function. Although the gene corrections were significant, using the corrected cells to restore function is a task that is more easily said than done.

Blood The b-hemoglobinopathies, including b-thalassemia and sickle cell disease, are a group of genetic blood diseases caused by mutations in the b subunits of hemoglobin (HBB). These mutations are responsible for the decreased production of functional or mature b-globin subunits and therefore reduce the oxygen-carrying capacity of the blood [91]. b-Thalassemia can be caused by different mutations, whereas sickle cell disease is caused by a single DNA base mutation (A > T) in the sixth codon of the hbb gene [92,93]. Allogeneic stem cell transplantation is an effective way to treat b-hemoglobinopathies but it is difficult to find a fully matched donor, so immune rejection of the transplanted cells is often a concern. The combination of genome editing with (autologous) stem cell technology has brought about a potentially improved approach to treat these blood disorders. For instance, a CRISPR-Cas9 system consisting of Cas9 ribonucleoproteins and a homologous, unmutated DNA sequence (carried by AAV) was used to achieve homologous recombination at a site of mutation in the hbb gene in isolated hematopoietic stem cells. These hbb-corrected cells were further enriched and transplanted into nonobese diabetic, severe combined immunodeficiency mice, yielding edited cells in the bone marrow of the mice [94]. In the same study, Dever et al. used a CRISPR-Cas9 system to correct the (A > T) mutation seen in sickle cell disease, following the standard logical scheme of isolating progenitor cells (in this case, hematopoietic stem and progenitor cells [HSPCs]) taken from patients with the disease, and then performing gene editing. The corrected HSPCs were able to differentiate into erythrocytes expressing b-globin mRNA. Two more examples of gene editing directed toward correcting sickle cell disease again used a CRISPR-Cas9 system. This time, a ribonucleoprotein complex consisting of the Cas9 protein and an unmodified gRNA was delivered with a single-stranded DNA oligonucleotide to patient-derived HSPC cells by electroporation. Correction efficiencies of up to 25% were reported [95]. The corrected HSPCs were used to restore the expression of wild-type hemoglobin after differentiation into erythroblasts, with the correction being maintained through 16 weeks after injection [95]. A remarkable correction efficiency of up to 67.9% was reported by Li et al. when they used CRISPR-Cas9 with adenovirus, followed by nucleoporation with a single-stranded DNA template, to correct the sickle cell disease mutation in patient-derived iPSCs [96]. As we have seen, the CRISPR-Cas9 system is a popular route for gene editing for treatment of b-hemoglobinopathies. In fact, this system could be deemed the preferred method for editing. A series of studies have demonstrated that CRISPR-Cas9 nucleases are relatively efficient and specific in gene editing in human iPSCs [91,97e99], with advantages over ZFNs and TALENs, even when used to correct a point mutation such as that seen in sickle cell disease [100]. However, this certainly does not invalidate efforts made with the other editing systems, which have been shown to be superior to CRISPR-Cas9 in certain situations. TALENs have been used to correct separate mutations (the intron 2 mutation IVS2-654C > T and the TCTT deletion) in the hbb gene of patient-specific iPSCs, made to enter cells via electroporation. The corrections were achieved with high efficiency: 68% and 40%, respectively [101]. TALENs and CRISPR-Cas9 have been separately employed, combined with a PiggyBac transposon DNA donor, to correct IVS2-654C > T mutations in the globin gene [102]. In this case, contrary to findings reported in the previous paragraph, TALENs performed with higher efficiency than CRISPR-Cas9 (33% versus 12.3%). In addition, the TALENs generated fewer off-target events. Not all such CRISPR-Cas9 applications display off-target effects, however. Fei et al. employed CRISPR/Cas9 combined

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42. GENE-EDITING IN REGENERATIVE MEDICINE

with the PiggyBac transposon to correct a 28 (A/G) mutation and TCTT deletion seamlessly in the hbb gene in patient-derived iPSCs to restore HBB expression after hematopoietic differentiation [103]. In this case, off-target events were not observed. In addition to therapeutic applications toward b-hemoglobinopathies, genome editing has been employed to increase red blood cell production for transfusion purposes. The SH2B3 gene encodes a protein that serves as a negative regulator of cytokine signaling. Naturally occurring dysfunctional variants of the protein can increase the population of red blood cells. CRISPR-Cas9 has been used to delete the SH2B3 gene in pluripotent stem cells, allowing for the increased production of erythroid cell populations in vitro [104].

Skin Epidermolysis bullosa dystrophica is a genetic disease caused by mutations within the gene encoding the collagen a-1 (VII) chain (COL7A1). Afflicted patients have fragile skin that is susceptible to minor injury and they easily form blisters and skin erosions. Efforts targeted to correct this genetic abnormality have been made by codelivering Cas9 and gRNA into iPSCs generated from patients with dominant dystrophic epidermolysis bullosa [105]. After the editing, the iPSCs were differentiated into COL7-secreting keratinocytes and fibroblasts. Another approach was to deliver TALENs, via electroporation, into primary fibroblasts isolated from a patient with the disease in an effort to correct the COL7A1 mutation [106]. The corrected fibroblasts were then driven into iPSCs and placed into the flanks of SCID mice. Results showed that the cells expressed normal COL7A1 protein and that skin-like structures were formed after implantation.

Nerve As with so many of the applications that have been covered thus far, the CRISPR-Cas9 system has been used to correct genetic mutations associated with neural disorders. The common scheme is to isolate progenitor cells (or to induce isolated cells into iPSCs) from a patient with the affliction and then to correct the mutation with the editing system. The resulting corrected cells are then differentiated and implanted back into the patient. This scheme has been performed on, for example, cells from patients with Alzheimer disease to correct a point mutation in the presenilin 1 gene [107]. It has also been used in conjunction with the PiggyBac transposon system to correct point mutations that result in disruption of tetrahydrobiopterin synthesis or recycling, which affects tyrosine and dopamine levels that are important for proper brain function [108]. The system has also been used to bring about genetic deletions, such as to combat fragile X syndrome, a type of inherited intellectual disability that is caused by an expansion of an area of CGG repeats in the fragile X mental retardation 1 gene [109]. As mentioned in the Nickases section, Cas9 can be mutated to alter its endonuclease activity, resulting in enzymes known as nickases. Whereas wild-type Cas9 allows for genome editing by introducing targeted double-strand breaks in the genome, mutations of the HNH or RuvC domains in Cas9 can yield nickases that cut only one strand of DNA [110,111]. Some have introduced mutations that stripped Cas9 of all endonuclease activity [112,113]. These mutants, called nuclease-deficient Cas9 (dCas9), have been used for transcriptional repression and activation. CRISPR interference technology uses a fusion of Cas9 with repressor domains to yield effective gene repression. CRISPR-mediated gene activation incorporates a fusion of dCas9 with activation domains to recruit transcription activators in the nuclei of mammalian cells [114]. As an example, consider an application that targeted a-synuclein, which is encoded by the SNCA gene and is associated with mediating the pathogenesis of Parkinson disease. Under the mediation of carefully designed gRNA, a dCas9 fused with a hybrid VP64-p65-Rta tripartite activator (dCas9-VPR) [115] was used to obtain upregulation of SNCA expression by eightfold compared with iPSC-derived neurons expressing normal a-synuclein levels, whereas dCas9-KRAB served as a transcriptional repressor to downregulate SNCA expression by 40% in a-synuclein triplication iPSC-derived neurons [116].

Retina Retinitis pigmentosa is a group of inherited eye diseases that are caused by mutations within a group of genes. Patients with retinitis pigmentosa have impaired vision caused by the degeneration of photoreceptor cells within the retina. A point mutation (c.3070G > T) in the retinitis pigmentosa GTPase regulator gene was corrected using

REFERENCES

755

CRISPR-Cas9 with about 13% efficiency in patient-derived iPSCs, which laid the foundation for autologous iPSC transplantation to treat retinal diseases [117]. In another study, CRISPR-Cas9 was used to disrupt a mutation in the rhodopsin gene in the S334ter rat model of retinitis pigmentosa, halting retinal degeneration and improving optomotor reflexes (a possible indicator of improved visual function) [118].

CLOSING REMARKS Gene editing involves the cleavage of double-stranded DNA in a targeted fashion. Although restriction enzymes have been used in molecular biology for about half a century, the field of genome editing has arisen relatively recently. The breakthrough in the field of genome editing came with the discovery of endonucleases that have tunable targeting and can be used in living cells. Three major forms of such systems were presented in this chapter: CRISPR/Cas, TALENs, and ZFNs. For the researcher, the choice of editing system should take into account both the efficiency of a given method and the number of off-target effects generated. A system that produces the desired genetic change in 99.9% of treated cells would not be desirable for clinical use if an off-target insertion, deletion, or base change caused 0.1% of the cells to be transformed into cancer cells. As covered earlier, much work has been performed to generate enzymes that operate with greater specificity in the hope of eliminating off-target effects. The general scheme used for gene editing for regenerative medicine includes isolating cells from the patient, which will greatly reduce the chances for an immune response to the therapy. The cells may be pluripotent, multipotent, or differentiated, but a common theme is to use or produce stem cells (such as iPSCs). The cellular genomes are then edited and the cells are reimplanted into the host in a pluripotent or differentiated form, with the differentiated form appearing to be more common. Differentiation of the cells before implantation gives the investigator a chance to verify that expression of the edited gene is at a clinically relevant level, whereas implantation of undifferentiated cells may allow for differentiation into multiple cell types in the proportions that the body needs. The implantation itself can occur in the form of the injection of cells into a given organ or tissue, or it could include cellular encapsulation or the embedding of cells into a scaffold construct. These implantation methods are options that are addressed elsewhere in this book. The regenerative applications presented in this chapter addressed congenital aberrations, or disease states, that were dealt with via gene editing. In the future, it may be possible to regenerate noncongenital tissue defects, such as those arising from injury, by editing cells to be specifically responsive to an engineered repair environment. Editing may be used to create cells that can be used as tools to guide the differentiation of stem cells in vivo, recruit other cells to an area, induce angiogenesis, or proliferate into a tissue that has an enhanced property to aid the host at the organismal level. With new applications and advances appearing in the literature nearly weekly, it is difficult to predict the precise directions that this burgeoning field will take next.

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C H A P T E R

43 Preclinical Bone Repair Models in Regenerative Medicine Elvis L. Francois1, Michael J. Yaszemski2 1

Department of Orthopedic Surgery, Mayo Clinic, Rochester, MN, United States; 2Departments of Orthopedic Surgery and Biomedical Engineering, Mayo Clinic, Rochester, MN, United States

INTRODUCTION Bone tissue engineering (BTE) is a promising avenue of research that has the goal of providing novel methods to add to our clinical capabilities for treating difficult segmental and contained skeletal defects. Bone tissue demonstrates the intrinsic properties of regrowth and self-repair, which is a process marked by a complex array of biologic, structural and metabolic functions. The primary building blocks of bone tissue regeneration center on the recapitulation of the natural signaling pathways of bone development and healing, which are modified toward developing modalities to stimulate bone formation in a clinical situation in which the skeletal defect may not heal using currently available methods. Of the many tissues under investigation, bone repair models continue to be a promising avenue of study primarily owing to the ever-increasing demand for and short supply of bone substitutes [1]. It is estimated that two million bone graft procedures are performed annually worldwide [2]. In the setting of large bony segmental defects that are often caused by polytrauma, pathological fractures, and osteonecrosis, the capacity for the normal process of fracture healing to repair the skeletal defect and restore load-bearing function to the injured bone is often insufficient, and the result is a fracture nonunion. This further marks the importance of preclinical bone repair models and their value for clinical application. This chapter will provide a general focus on the biological processes of preclinical bone repair in vitro and in animal models. Ethical issues must be addressed when animals are used as preclinical models as part of the testing program. The appropriate conduct of experimentation using live animals is important to progress in the care of both humans and animals. The responsibility for the appropriate use of animals in research is incumbent on the investigators. The three “R’s” of animal experimentation are replacement, reduction, and refinement. Replacement is the process of seeking to replace animals in an experimental design by either using an in vitro method or a phylogenetically lower animal whenever possible. The reduction process is constant assessment of the experimental design in an effort to reduce the number of animals by asking whether the desired data can be obtained in a statistically valid manner using fewer animals. Finally, the refinement process is an attempt to improve existing experimental methods to obtain the desired data with reduced ethical costs in terms of any painful or stressful procedures that are done on the animals.

BIOMINERALIZATION AND BONE REGENERATION The triad of building blocks for bone tissue regeneration are: functionally active osteoblastic cells to secrete new bone matrix (osteogenesis), scaffolds upon which these anchorage-dependent cells attach and from which signaling Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00043-6

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molecules can be delivered to the bone regeneration space (osteoconduction), and growth factors to drive the regeneration process (osteoinduction) [2].

CELL SOURCES Stem Cells in Bone Tissue Engineering Stem cells require two fundamentally important characteristics: the ability to self-renew and generate daughter lineages with identical potentialities, as well as the ability to differentiate along one or more lineages [3]. In the field of BTE, commonly used sources of stem cells include embryonic stem cells (ESCs), induced pluripotent stem cells (iPSCs), and adult mesenchymal stem cell (MSCs).

EMBRYONIC STEM CELLS Human ESCs (hESCs) are derived from embryos generated via in vitro fertilization. Embryos are most often obtained from donors after they are deemed unsuitable for implantation and appropriate consent has been obtained. The preparation of embryos occurs via an antibody application process in which the inner cell mass from the blastocyst stage of the embryo is separated from the trophectoderm. The cells that are derived from the blastocyst inner cell mass are initially plated onto feeder cells. The cells that constitute the inner cell mass then expand to form an hESC cell line [4]. ESCs have two distinctive properties: the ability to proliferate infinitely and the property of pluripotency to differentiate into all three germ layer cell types: ectoderm, endoderm, and mesoderm [5]. Of specific interest in BTE, hESCs can be guided toward differentiation into osteoblast cells in vitro. One method of differentiation was described by Bielby, in which osteogenic cells are derived from three-dimensional (3D) cell spheroids (embryoid bodies [EBs]) [5]. EBs are formed via suspension or hanging drop methods from a single cell suspension. EBs provide the potential for differentiation into precursor cells of all three germ layers. Committed cells in the EB matrix are cultured in monolayer and are induced to osteogenic cells under the presence of exogenous factors (e.g., dexamethasone-ascorbic acid and sodium-b-glycerophosphate). These exogenous factors work in concert to promote an environment conducive to bone regeneration. Dexamethasone stimulates the osteogenic differentiation of precursor cells harvested from multiple tissues. Ascorbic acid facilitates collagen secretion and deposition, and sodium-b-glycerophosphate acts to mineralize the deposited matrix. Working in concert with biochemical exogenous factors, the extracellular matrix (ECM) has an important role in the direction and differentiation of ESCs. Alternatively, undifferentiated ESCs or dispersed EBs are sometimes implanted directly into 3D scaffolds and driven to form multiple tissues for later implantation [6].

Induced Pluripotent Stem Cells iPSCs are somatic cells reprogrammed to exhibit pluripotent properties. Historically, mouse skin fibroblasts were first reprogrammed to iPSCs by overexpression of a set of four key transcription factors [7]. Subsequently, adult human cells were reprogrammed [7,8]. Systematic reviews of the literature on the osteogenic potential of iPSCs suggest that osteo-induced iPSCs demonstrate an osteogenic capability equal or superior to MSCs [9]. Furthermore, studies demonstrate that cell sources for iPSCs do not appear to affect the osteogenic potential of iPSCs [9]. In their early use, it was noted that iPSCs were met with complications of teratoma formation. In an attempt to address these complications, studies have demonstrated that the addition of resveratrol to the osteogenic medium and irradiation after osteogenic induction reduce teratoma formation in animal models [9].

Mesenchymal Stem Cells The most common stem cell type used in the BTE field is the MSC [10]. First identified in 1966 by Friedenstein, MSCs were discovered via the isolation of bone and cartilage-forming cells from rat bone marrow cells that had fibroblast-like morphology [11]. MSCs have since been isolated from multiple tissues (liver, fetal blood, and umbilical cord), but the most well-investigated and readily available source of MSCs is the bone marrow. MSCs harvested from different sources may not be identical in their biological characteristics. A comprehensive report on the

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proteome and transcriptome profiles of MSCs revealed source-specific markers [12]. Furthermore, differences that exist in colony-forming unit efficiency, surfactome profiles, multilineage differentiation, and paracrine functions may suggest a specific optimal clinical application for that particular MSC cell line [13,14]. MSCs comprise roughly 0.001%e0.1% of the total population of bone marrow nucleated cells. In turn, these can be specifically separated from other nucleated cells by their adherence properties to plastic culture flasks. They can then be expanded in vitro without a loss of phenotype. The process of harvesting MSCs and then exploiting their pluripotent ability to differentiate them into osteoblasts, chondroblasts, myoblasts, and tenocytes is primarily why they are often used in musculoskeletal tissue engineering applications [15]. MSCs have the ability to differentiate into a variety of adult musculoskeletal cells as well as an inherent ability to secrete a variety of cytokines that modulate inflammatory and immune response pathways [16]. The synergistic benefit of these immunomodulatory effects is a reduction in the risk for host rejection. Another beneficial property of MSCs is their ability to be driven down various mesenchymal cellular pathways including chondrogenic, osteogenic, and adipogenic lineages when provided the necessary in vitro or in vivo environments [17] or 3D scaffolds [18]. MSCs are driven toward osteogenic differentiation when stimulated with dexamethasone, ascorbic acid, and sodium-b-glycerophosphate. When appropriately stimulated, MSCs upregulate alkaline phosphatase, osteocalcin, and the expression of osteopontin, and they direct calcium deposition within the ECM. In vivo preclinical BTE models also benefit from the use of MSCs, in that the MSCs have been shown to facilitate bone repair when they have anchored onto artificial matrices such as hydroxyapatite scaffolds [19].

SCAFFOLDS BTE scaffolds allow for the adherence of osteogenic cells and serve as an appropriate microenvironment to permit those cells to secrete osteoid, the matrix of newly formed bone tissue. Scaffolds are often subdivided into three classes based on their material composition: polymers, metals, and bioceramics. Scaffolds may contain more than one of these three material types. Further stratification is delineated by the derivation of the material (i.e., natural versus synthetic) and their ability to undergo degradation (i.e., resorbable versus nonresorbable). The benefits of naturally derived scaffolds (i.e., collagens, fibrin, elastin, alginate, hyaluronic acid) are their resorbable qualities in vivo as well as their intrinsic biocompatibility and minimal adverse immunological properties. Synthetic scaffolds have the benefit of often being able to be fabricated with a wider range of degradation rates and mechanical properties compared with their natural counterparts. This ability is based on their polymeric composition, the copolymer ratio, and the interactions of their polymeric side chains. The choice of the synthetic polymer’s chemical composition and 3D configuration often determines its cell adhesion properties and ability to incorporate and deliver bioactive molecules in a controlled fashion [20]. A novel arena within scaffold design has been the adoption of additive manufacturing technologies (3D printing) in scaffold fabrication. The term “3D printing” describes a group of additive manufacturing technologies that are collectively referred to as solid free-form fabrication. These include laser stereolithography, fiber electrospinning, fused deposition modeling, fiber extrusion from the melt, and injection molding [21]. The benefits of 3D printing are conferred by its immense flexibility in fabricating scaffolds of varying structural complexity. This process allows for a great deal of control over the construct architecture and flexibility in scaling up fabrication, and it has the added benefit of technically precise reproducibility, which is sometimes lacking in subtractive fabrication techniques (e.g., milling).

Biochemical Signaling: Growth Factors and Cell Signals BTE cell signals are environmental factors that directly or indirectly influence the regeneration of skeletal tissues. The importance of appropriate cell signals and growth factors cannot be overstated in terms of their importance for bone formation. Common growth factors of particular importance in bone regeneration include bone morphogenic protein (BMP)-2, BMP-7 (also called osteogenic protein 1), fibroblast growth factor, platelet derived growth factor, parathyroid hormone (PTH), PTH-releasing peptide, transforming growth factor-b3, vascular endothelial growth factor, and the Wnt proteins [22]. The interplay between signaling molecules and growth factors is complex and multivariable. Factors critical to cell signaling include the spatiotemporal release of growth factors and their bioactivity. The delivery of biochemical signaling cues is generally categorized as unbound, bound within the implant with a designed controlled delivery, coated on the implant surface, or coded within the cells via gene delivery mechanisms [23,24]. Unbound delivery systems are marked by the rapid efflux of the growth factors followed by rapid clearance from the microenvironment. Bound delivery systems allow the advantage of controlled or specific variable release of the biochemical signal over time. Generally speaking, bound delivery

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systems are more applicable for the requirements of BTE. Hydrogels are the most commonly investigated polymer for cell encapsulation and the in situ delivery of biochemical cues [22]. To mimic the function performed by the ECM, bioactive hydrogels containing protease sensitive sites, cell adhesion molecules such as Arg-Gly-Aspe containing peptides, and/or biological cues in the form of growth factors, inorganic minerals, or drugs have also been developed [25,26]. Of the many growth and differentiation factors in BTE, BMPs possess the unique ability to stimulate the differentiation of mesenchymal precursor cells to chondrocytes and osteoblasts; this in turn allows for the induction of new bone at ectopic and orthotopic sites. Marshall Urist discovered that demineralized bone matrix induced ectopic bone formation in subcutaneous and intramuscular pockets in rodents [27]. Further investigation over the next several decades led to the isolation of BMPs as the causative factor in this induction of the bone formation cascade [28]. Clinically, BMPs are often used at the sites of intended bone regeneration in skeletal defects with the goal of osteoblastic cell proliferation and differentiation. The delivery of BMPs can be either direct application to the intended site of bone regeneration or via controlled release. There are several drug delivery systems under investigation that allow for measured controlled delivery. BMPs as well as other proteins may be encapsulated in poly(D,L-lactic-co-glycolic acid) microspheres [29] or embedded into collagen carriers [30]. Among other delivery systems, these allow for the temporal and spatial release of growth factors and cell signals in a controlled fashion.

PRECLINICAL MODELS OF BONE TISSUE REGENERATION In Vitro Preclinical Models In vitro studies offer the advantage of focused manipulation of specific biomaterials in well-controlled experiments. In targeting specific cascades, investigators are more capable of understanding specific variable responses in a closed environment with the goal of predicting future responses in preclinical animal models and in subsequent translational first-in-human applications. In vitro studies afford the opportunity to simulate complex interactions among cells, scaffolds, and growth factors in relatively controlled environments. Bone-forming cells (osteoblasts) are often procured from three primary origins: • pluripotent stem cells that differentiate into osteoblasts [31], • primary osteoblasts and osteocytes from different species [32,33], and • immortalized cell lines [34,35]. To maintain translational validity, bone in vitro models must incorporate the classic triad of cells, scaffolds, and growth factors in a fashion that closely mimics both the biochemical and biomechanical bone formation characteristics that are seen in vivo. The bone matrix is composed of cells surrounded by an ECM made up of an inorganic phase that is approximately 70 wt% hydroxyapatite and an organic phase that is approximately 30 wt% type I collagen and other noncollagenous proteins [36,37]. Biomechanical studies have demonstrated that greater than 90% of bone cells within the ECM are osteocytes that sense and transduce mechanical forces exerted on the bone, in turn governing the rates of resorption and deposition that occur during bone remodeling [38]. Historically, the study of the interplay between bone physiological and pathological processes has been performed on 2D plastic plates. Although great strides in understanding bone formation, bone resorption, and bone remodeling have been made via this medium, the 2D cell cultures have been shown to mimic the 3D microenvironment of native bone and the interactions between the bone cells and the ECM incompletely [39]. In vitro studies provide great insight into the translational validity of a model, but the biomechanical environment in vivo is much more complex and the stresses exerted on the bone dictate its resorption and deposition rates. This complex interplay between the microscopic cellular level and the macroscopic mechanical level are greatly important in BTE, because understanding these interactions is critical to the success of in vivo BTE preclinical models.

In Vivo Preclinical Models The benefit of in vivo models rests in the ability to assess biomaterials in more complex tissue environments that have variable loading conditions. Among the limitations of BTE animal models is the interspecies variation of bone tissue. It has been well-illustrated that bone composition, density, and mechanical properties of various animal models (i.e., cow, sheep, pig, dog, chicken, and rat) often differ from those properties in human bone [40].

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There are a host of factors to be considered when selecting a specific animal model for in vivo preclinical studies. Chief among these is that the animal model demonstrates physiological and pathophysiological processes that mimic those in humans. The attributes of a “good” animal model are that in addition to being similar to those in humans, as mentioned earlier, the animal’s anatomy and physiology are suitable for the requirements of the study design. The animal chosen should be economical and available. The animal management and care issues must have been considered and are optimal for the chosen study (e.g., housing requirements and ease of handling).

Selection Considerations Based on Animal Species Animal species that are under consideration for in vivo preclinical studies include mice, rats, rabbits, sheep, goats, pigs, dogs, cats, and nonhuman primates. There are pros and cons to the use of any of these animals. Mice and rats are readily available and inexpensive and have minimal housing requirements. On the negative side, their small size is an issue in some experiments, and their life span postskeletal maturity places a time limit on the study. Human clinical conditions that have been modeled by rats and mice include heterotopic bone formation, trabecular bone defects, segmental bone defects, and spinal cord injury. Given their size, robustness, and cost efficiency, rodents are particularly useful in preclinical studies in assessing biomaterials as bone substitutes and are generally regarded as the prime model for in vivo testing of the regeneration of bone tissue [41]. Disadvantages, in addition to those listed earlier, include the lack of Haversian-type remodeling in the rodent bone cortex and thin, fragile cortices in the long bones [42]. The rat, which is the most often used animal model, has been shown to have the most significant difference from human bone compared with other animal models [43]. Rodents are primarily useful for the surgical implantation stage of substitute bone materials. For example, b-tricalcium phosphate (b-TCP), calcium phosphate, and collagen experiments have been commonly conducted in rodents. A prime example of applications of rodent bone defect models is exemplified by investigations of the biocompatibility of highly purified b-TCP bone graft substitutes using a rat femoral defect model. The study demonstrated that purified b-TCP was both biocompatible and resorbable [44]. In another study, investigators successfully used small animal rodent models to establish a 4-mm-diameter calvarial critical-sized defect model in mice. This model was successful in the analysis of the in vivo osteoconductive and osteoinductive abilities of bone substitute materials [45]. Similar to mice and rats, rabbits are also readily available and inexpensive, and have minimal housing requirements. They are larger than mice and rats but are still considered a small animal model. Rabbit models rank as the most commonly employed models in musculoskeletal research [46]. Applications that have used rabbit models include calvarial critical-sized bone defects, posterolateral spine fusion, and cartilage regeneration. Rabbits have similarities to humans in bone mineral density and fracture toughness of middiaphyseal bone. Large animal models allow for the assessment of a larger volume of bone regeneration and repair over a longer time frame than is possible in mice, rats, and rabbits. Large animal models permit the assessment of bone remodeling and implant integration in a manner that better mimics the biomechanics and loading characteristics seen in humans. Large animal models that are often used in musculoskeletal investigations include sheep, goats, pigs, dogs, cats, and nonhuman primates. The use of sheep, goats, and pigs provides an animal model in which bones and joints are more similar to their counterparts in humans than are those in the small animal models discussed previously. Sheep, goats, and pigs have good availability and can serve as an alternative to dogs in some applications. Dogs and cats are companion animals, which often causes their use to receive greater scrutiny. Negative issues associated with the use of sheep, goats, and pigs include the increased cost, housing requirements, and the need for a formal operating room setup to perform surgery on them. Applications for which sheep, goats, and pigs have been used include radius nonunion (and other bone healing or bone defect applications), femoral head osteonecrosis, anterior cruciate ligament reconstruction, and meniscal repair. Dogs and cats share the same pros and cons as sheep, goats, and pigs. In addition, as mentioned earlier, their status as companion animals often attracts greater scrutiny in their use in experimental designs that include surgical procedures. Several applications that have used dogs and cats as experimental animals include many of the same applications for which sheep, goats, and pigs were used: radial nonunions, tibial defects, other fracture healing or bone defect models, femoral head osteonecrosis, and craniomandibular reconstruction. In addition, dogs and cats have been used in surgical studies for total joint arthroplasty, spinal cord injury, and distal radius osteosarcoma. Nonhuman primates have an anatomy and physiology that more closely parallel those of humans than any of the other animal models discussed here. They are not readily available, they are expensive, and they have the highest housing requirements than any of the other animals. The scrutiny that they receive from the Institutional Animal

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Care and Use Committee when their use is requested is even higher than that of the companion animals. Applications in which nonhuman primates have been used include osteoporosis, bone healing, fracture nonunions, prosthetic implant studies, spinal fusion, and organ transplantation studies. Surgical studies using animals are essential for the analysis of novel treatments for both humans and animals. However, we must treat them humanely, take whatever measures are necessary to control their surgical pain, and constantly seek to apply “the three R’s” of replacement, reduction, and refinement to obtain the data that we need, and simultaneously try to do so while using increasingly fewer animals.

CONCLUSIONS No amount or type of preclinical (in vitro and animal) studies can predict with certainty the implant behavior in humans. The first in-human use must consider clinical equipoise: the anticipated balance between potential benefit and potential risk in the human study subject who receives the implant. Consider the regulatory implications of potential preclinical experimental pathways, because different but feasible may pose different regulatory burdens. Plan to speak with US Food and Drug Administration colleagues early and often as the various potential preclinical paths are being considered. Remember that animal models are an essential but insufficient component of the preclinical evaluation for new medical products. The choice of model depends on several factors, including the biologic and structural goal of the study, the applicability of the model to the human condition under evaluation, the cost and technical feasibility of the chosen model, and historical experience with the model. Eventually, the safety and efficacy of new treatment modalities will be determined by well-controlled studies in humans, and postevaluation use in the general population of patients who have the condition to be treated. Be certain that clinicians who care for patients with the disease in question are members of the planning and execution of the preclinical study team. The ultimate determination as to whether your product will become the standard of care for the surgical care of the disease condition under consideration will be made one patient at a time, within the physicianepatient relationship as the treating doctor and the patient discuss the various treatment options available to them to treat the patient’s clinical problem.

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44 Body-on-a-Chip: Regenerative Medicine for Personalized Medicine Aleksander Skardal1,2,3,4, Thomas Shupe1, Anthony Atala1 1

Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States; Virginia Tech-Wake Forest School of Biomedical Engineering and Sciences, Wake Forest University, Winston-Salem, NC, United States; 3Comprehensive Cancer Center at Wake Forest Baptist Medical, Winston-Salem, NC, United States; 4 Department of Cancer Biology, Wake Forest University, Winston-Salem, NC, United States 2

INTRODUCTION Current drug development strategies do not include adequate models to predict drug efficacy or safety in humans. There is a considerable need for the high-fidelity in vitro representation of integrated human physiology to test both the beneficial and potential detrimental effects of drug candidate compounds in the body [1,2]. Animals (rodents, in particular) have been the reference standard of scientific experimentation for centuries. These animal models have served as the cornerstone of research in cell biology, pathobiology, molecular biology, and many other biomedical research fields [3]. Two-dimensional (2D) cell lines and animal models have been used extensively to determine the toxicity of new drugs before the initiation of human clinical trials. It has become clear that these models have significant limitations owing to phenotypic differences in physiology compared with humans. This is especially critical for the assessment of toxic side effects of drugs that might target the liver, heart, and other organs because of the differences in enzymatic expression profiles among humans, rodents, and cell lines, which often results in significant differences in drug metabolism, efficacy, and toxicity. In vitro drug screening platforms use human cells; however, by definition, the genotype of cell lines is altered from the natural state, and primary cells change phenotype after removal from native tissue. These phenotype changes in culture result from the failure of traditional 2D systems to recapitulate several aspects of the native 3D cellular microenvironment [4,5]. As a result, 2D cultures exert selective pressures on cells, significantly altering their phenotype as they adapt to their new conditions. Drug diffusion kinetics are not modeled accurately in 2D tissue culture, drug doses effective in 2D are often ineffective when scaled to patients, and the lack of cellecell/cellematrix interactions in 2D often lead to loss of cell function [4,6,7]. In contrast, advances in tissue engineering, biomaterial development, and microfluidics and electronics have resulted in the successful fabrication of multicellular human tissue equivalents and microorgans (organoids) that demonstrate many of the functional properties of normal human tissue and organs. For example, liver organoids exhibit normal metabolic activity, skeletal and cardiac muscle constructs contract in a physiologically normal manner, lung organoids “breath,” and gut/vessel/brain microvasculature constructs maintain normal barrier functionality [8,9]. “Body-on-a-chip” devices that recapitulate 3D tissue architectures and physiological fluid flow conditions are more effective at supporting normal cell function than static 2D culture [10]. These engineered platforms can include sophisticated hardware systems, potential for scale-up, capacity for high throughput, and user control over physical factors such as fluid shear stress and mechanical deformations. Microfabrication techniques based on a variety of nanotechnologies have resulted in the development of microscale fluidic systems with predictable fluid dynamics

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throughout the entire fluid circuit [11]. Tissues and organoids can be immobilized within platform microreactors using sophisticated hydrogel biomaterials, providing a proper microenvironment and allowing for long-term perfusion. A variety of iterations on these basic concepts are in use by laboratories around the world [12e14]. In addition, a variety of on-chip disease models have been investigated [13]. The most pressing challenge in advancing the body-on-a-chip field is to combine multiple organs within a common microfluidic circuit, to model an entire human, on-chip. Such a system would represent the pinnacle of in vitro platforms for modeling integrated human physiology. Ideally, as in the human body, such a system would recapitulate the interdependent and synergistic functions of all tissues and organs within a cell culture or body-on-a-chip platform. The microfluidic circuit connecting organoid microreactor chambers allows for fluid flow across each organoid type in a sequence that mimics blood flow throughout the human body [8]. Compound metabolites and other secreted factors would likewise be transported to downstream organoid types in a physiologically relevant sequence. As such, these advanced body-on-a-chip platforms would be ideal for testing newly developed drugs and assessing potential toxic side effects in human tissues and organs. Furthermore, the body-on-a-chip platform would offer tremendous benefits for pharmacological studies aimed at determining the specific effects and toxic levels for newly developed drugs, allowing for the better prediction of appropriate doses for human trials. In this chapter, we highlight a variety of organoid-on-a-chip systems for applications such as drug screening and disease modeling and look to the future of multiorganoid body-on-a-chip systems and applications in personalized precision medicine.

ADVANCE OF IN VITRO ORGANOID DEVELOPMENT: PROGRESSION FROM TWO-DIMENSIONAL TO THREE-DIMENSIONAL MODELS The development of novel drugs that are safe and effective in humans has been significantly hampered owing to the inability to model human cell phenotype, function, and intercellular signaling accurately in vitro. Animal models used extensively in preclinical drug studies are traditionally regarded to be the reference standard for drug testing. However, animal models do not reflect human drug metabolism accurately; thus, animals are often not predictive of results in humans. The second type of conventional model system, in vitro 2D culture, fails to recapitulate many aspects of the 3D cellular microenvironment, leading to poor support for cell viability and cellular function [4,15]. In addition, drug diffusion kinetics are completely nonphysiological in traditional cell culture and drug doses that are effective in 2D are often ineffective when scaled to patients [6,7]. Cells grown on tissue culture plastic experience several properties in their environment that are inconsistent with the tissue from which the cells were originally isolated. These include surface microtopography, stiffness, oxygen tension, mechanical loading, biochemical composition, and most important, a 2D rather than 3D architecture. These unnatural characteristics can significantly alter the phenotypic properties of cells, because they are forced to adapt to these new conditions. The functional differences between 2D cultures and 3D constructs have begun to be appreciated. A myriad of studies have demonstrated these differences across a variety of cell types. 3D systems consistently outperform 2D cultures in many aspects including accurately representing in vivo function and demonstrating a physiologically normal response to drugs and toxins [16]. The current drug development pipeline (Fig. 44.1A) has not yet evolved to include newer 3D cell culture technology, which results in countless discrepancies between in vitro drug screening outcomes and later performance in patients, during or even after clinical trials [17]. As an example, our team demonstrated that metastatic colon carcinoma cells adopted an epithelial appearance in 2D tissue culture. However, when transitioned to a 3D organoid form factor, the cancer cells “switched” to a morphology that resembled mesenchymal metastatic cells, which were much more representative of malignant tumor cells in vivo [18]. These kinds of documented benefits of 3D cell culture beg the question: Why are 2D cell cultures still being employed in drug development and toxicology screening? Fortunately, tissue engineering technologies have evolved to the point that microengineered tissue constructs can better mimic the structure, cellular heterogeneity, and function of in vivo tissue. These organ models can often be maintained in viable states for longer periods and are designed to preserve the functional properties of native tissues. They can also recapitulate the microenvironmental roles of cellecell, celleextracellular matrix (ECM), and mechanical interactions that cells experience inside tissues. Oxygenation can be a concern in 3D tissue models. If they are too large in diameter, an oxygen gradient develops across the organoid that can lead to phenotypic changes and potentially a necrotic core. However, oxygen gradients exist in vivo. As such, as long these gradients are taken into consideration and controlled, either by limiting the size of the organoids or by creating perfusable channels within the construct, maintaining an oxygen gradient may actually provide a better representation of native tissue. Overall, these relatively new 3D model systems are greatly superior to their 2D predecessors for drug and toxicology testing.

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FIGURE 44.1 Potential improvements in the drug development pipeline as a result of deployment of organ-on-a-chip and body-on-a-chip technologies into pharmaceutical research and development. (A) The current drug development pipeline requires many years and multiple billions of dollars to bring a drug to market. (B) Plugging in human-based, biofabricated on-a-chip platforms into preclinical stages can potentially drastically improve the efficiency of the drug development pipeline. R&D, research and development.

Fortunately, the general concept of performing research using 3D versus 2D cultures has gained significant traction. However, hurdles and challenges remain. 2D cell culture is an established practice that will certainly remain a widely used tool for many years. This is primarily because it is easy and inexpensive compared with 3D culture systems. Implementing 3D systems in a laboratory can be complicated, requiring the mastery of new technologies including biomaterial development and biofabrication techniques. Furthermore, once 3D culture technologies have been established, processes regarded as trivial in 2D culture, such as cell harvesting and cell passaging, can be difficult and in some cases impossible without harming the cells. For example, if cells are cultured within a 3D hydrogel construct, one must effectively dissolve the matrix to isolate or harvest the cells. Some biomaterials support cell retrieval by building specific features into the material [19], but most do not; instead, they require enzymatic dissolution that in some cases can influence cell viability or phenotype. Also, most cell imaging techniques were developed for 2D cell cultures, environments in which cells exist in a narrow focal plane. In 3D, cells reside in many focal planes. Consequently, high-quality imaging in 3D may be obtained only by confocal or macroconfocal microscopy, requiring expensive equipment to which many laboratories do not have access. In addition, there are a variety of assays that can be significantly more difficult to perform on 3D models or that require significant modification for adapting to 3D models. Finally, some body-on-chip device materials (polydimethylsiloxane, for example) are prone to fouling and drug and protein adsorption. However, new materials for device hardware are being developed to solve this problem. All things considered, it is generally understood that 3D or dynamic on-a-chip platforms outperform static 2D systems in modeling normal human physiology [16]. As a result, these more capable 3D platforms have an immense potential to influence the drug development pipeline positively, decreasing development costs and increasing success rates of drug candidates in clinical trials (Fig. 44.1B). Perhaps just as important, these models can be used to identify nonoptimal drug candidates early before human trials are initiated.

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ORGAN-ON-A-CHIP TECHNOLOGIES AND THEIR APPLICATIONS Advances in biotechnology areas such as tissue engineering [20], biomaterials [21], and microbiofabrication [22] have allowed the derivation of new biological systems with massive potential as test platforms. Researchers have developed a wide variety of human-derived in vitro models that can be used to test drugs, toxins, and drug candidates in a more normal physiological environment [14,23e25]. Furthermore, advances in molecular genetics and tissue engineering technologies have enabled the development of 3D models of specific diseases [13,26e28]. Advances in related technologies such as microfabrication and microfluidic technology have improved 3D cell models by supporting cell and organoid culture, fluid flow, high-throughput testing, environmental sampling, and biosensing. These organ-on-a-chip systems can vary widely in design, allowing for the representation of a range of tissue types. Some of these systems have already been implemented in drug discovery [12] and purport to affect the future of medicine significantly. Here, we highlight just a few microengineering technologies and discuss examples of liver-on-a-chip, vessel-on-a-chip, lung-on-a-chip, and cancer-on-a-chip systems. However, there are many variations of these systems, as well as many additional models for virtually any tissue type.

Microengineering and Biofabrication The cellular content of a body-on-a-chip model is only one piece of the puzzle. Even with a physiologically normal composition of cells, normal cell function and appropriate responses to pharmacological agents are not guaranteed. Another environmental characteristic that must be considered is the interaction of cells with specific ECM proteins. As mentioned, cells receive a lot of information from their immediate local microenvironment that directly influences cell phenotype. Thus, it is critically important to reproduce as many of the components of the microenvironment as possible. Several techniques have been developed to incorporate natural structural and functional components of tissue ECM into cell culture systems. Micropatterning is the precise placement of proteins within a cell culture substrate. Micropatterning can be accomplished by a variety of methods including (1) microcontact printing, in which a stamp is coated with specific ECM components and is pressed against a solid substrate to create a specific pattern; (2) photopatterning using UV light projected through a photomask to catalyze the adherence of ECM proteins, or compounds that bind ECM proteins to a cell substrate in a predetermined pattern; and (3) laser patterning, in which laser light is used to mediate protein binding to a substrate in any desired pattern with very high resolution [29]. Many bioregulatory components of the ECM can be distributed in a controlled manner by using micropatterning techniques. For instance, depositing islands of ECM cell adhesion proteins that restrict cell spreading will induce apoptosis in bovine adrenal capillary cells while maintaining the differentiation state of epidermal keratinocytes. Conversely, micropatterning of ECM proteins that induce cell spreading promoted proliferation of both the bovine adrenal capillary cells and epidermal keratinocytes. These experiments demonstrate a clear link among the local ECM composition of a cell substrate, cell cycle, and differentiation [30,31]. This notion is further highlighted by another study involving micropatterning. Mesenchymal stem cells grown on small micropatterned patches that restrict cell spreading promoted differentiation toward adipogenic lineages, whereas micropatterning of factors that induce cell spreading promoted differentiation toward osteogenic lineages. These studies also showed that modulating cell shape was sufficient to induce the expression of signaling proteins Rac1 and N-cadherin, which have an important role in cell lineage specification [32,33]. Micropatterning represents a powerful tool for precisely controlling the protein composition of the cellular microenvironment within a cell culture substrate. By modulating the biochemical and geometric properties of the microenvironment, global cell phenotype and cell viability may be greatly influenced. Micropatterning has a wide variety of applications in advanced in vitro models and will become increasingly used to tune physiologic output from these systems. Whereas micropatterning technologies are generally used to control the cellular microenvironment in 2D culture systems, bioprinting provides a method for doing the same in a 3D space. Bioprinting involves the layer-by-layer deposition of structural material, cells, and bioregulatory factors in a controlled manner. Bioprinting technologies enable the fabrication of complex cellularized 3D constructs that may include components of the ECM, including both intrinsic and bound bioregulatory factors that modulate cell organization and function within a 3D space. Bioprinting is highly customizable across a wide range of resolutions and biochemical or physical characteristics. Applications requiring a more rigid structure can be printed using biomaterials with high mechanical stiffness. For stiffnesses beyond that which may be directly printed, methods have been developed for cross-linking structural components of the biomaterial subsequent to bioprinting. For example, dental implants have been bioprinted

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from polycaprolactone and hydroxyapatite [34]. On the other end of the stiffness spectrum, soft tissues such as vascular grafts have been printed using very low stiffness poly(ethylene glycol) hydrogels [35]. Because of the increasing resolution and speed of bioprinters, the structures that may be fabricated are becoming highly complex. The ability to precisely reproduce the structural and biochemical microarchitecture of tissues will most certainly result in more physiologically normal cellular function in bioprinted constructs. Several biological constructs have been fabricated using current bioprinting technology. Liver organoids have been generated using microextrusion bioprinting technology that includes several liver cell types suspended in a supportive hydrogel. These constructs demonstrated exceptionally high levels of liver tissue function and maintained viability in the long term, which makes them ideal candidates for tissue-on-a-chip applications [36]. Skin substitutes have also been created using a laser-based bioprinting technology, which allowed the precise placement of cells associated with specific layers of the skin. The resulting skin constructs were implanted into rodent wound models and demonstrated robust neovascularization, differentiation of mature keratinocytes, and the generation of a normal dermal basal lamina, all hallmarks of native skin [37]. 3D printing may still be considered a nascent field. Logistical obstacles continue to limit applications in whole-organ biofabrication. However, the speed, reproducibility, and scalability of bioprinting make it an ideal complement to body-on-a-chip modeling. 3D printing can be used to generate an industrial-scale volume of biological constructs with low run-to-run variability, with the complex structure required for high-function tissue organoid models.

Liver-on-a-Chip Early tissue/organoid-on-a-chip devices were geometrically designed to drive cell aggregation, thereby creating multicellular organoids. For example, devices were designed with microwells with a convergent geometry that terminated in a cell substrate of some type. Based on the microwell design, liver-derived cell lines could be formed into either spheroid or cylindrical constructs in a highly controlled manner. These 3D constructs maintained much better cellular function than did 2D controls [38,39]. In another example, spheroids were created from a cell line using an array of channels connecting inverted, pyramid-shaped microwells, allowing for the delivery of cells and test compound to multiple chambers simultaneously. This integration of microfluidics with an array of microreactors greatly increased the throughput potential for drug screening [40]. Liver-on-a-chip devices have become much more complex. They often employ controlled fluid flow to address nutrient circulation, drug or toxin administration, sample collection, and the integration of liver organoids with other tissue types. The latter will be discussed in detail later in this chapter. In one such liver-on-a-chip, hydrogels were used to encapsulate HepG2 cells with National Institutes of Health (NIH)-3T3 fibroblasts. These arrays of 3D organoids had increased liver function compared with 2D controls and produced an appropriate toxic response to acetyl-para-aminophenol (acetaminophen [APAP]) in a drug screening experiment [41]. Our group employed a versatile photopolymerizable hyaluronic acid biopolymer system for in situ photopatterning of HepG2 cells to generate 3D liver constructs. The constructs were formed in parallel channel fluidic devices that were fabricated by soft lithography and molded polydimethylsiloxane. This system was used for toxicity screening by administering multiple alcohol concentrations within each chip. As expected, alcohol administration resulted in a dose-dependent decrease in viability and cellular function [14]. Efforts within our group are focused on miniaturizing this and other systems to increase throughput further. Miniaturization and microfabrication approaches can be employed to generate more intricate biological microarchitecture such as liver sinusoids. Precise seeding and layering of hepatocytes and endothelial cells within microfluidic circuits can be used to generate structures with the resolution required to produce sinusoid-like models [42]. Another approach to generating biologically relevant microarchitecture involves mating synthetic and biological components. As an example, semiporous membrane to separate two adjacent chambers may be used to partition human hepatocytes from sinusoidal endothelial cells. Such a design was shown to generate higher albumin and urea production compared with traditional hepatocyte cultures; it demonstrates another strategy for recapitulating normal microarchitecture to increase cell function [43].

Vessel-on-a-Chip The term “microfluidics” carries with it the assumption of controlled fluid routing. Thus, microfluidic devices are effective for modeling vascular networks. Moreover, because drugs are generally introduced directly to the bloodstream or enter the bloodstream shortly after oral or airway introduction, fluidic systems that mimic the role of the

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vasculature contribution to pharmacokinetics represent a beneficial component for any drug screening technology. A substantial number of vascular-like fluidic devices have been developed, including both straight channels devices [44,45] and fluidic devices with more complex, branching features [46,47]. One major role of the vascular system, beyond transporting nutrients and oxygen among tissues, is to transport drugs and other molecules to sites throughout the body, where they pass through the endothelium into adjacent tissues. Many microfluidic systems have been designed to model the transendothelial delivery of test compounds to a target tissue. As an example, a device was developed that includes two perpendicular channels that cross at a single point. A semipermeable membrane colonized with an endothelial cell monolayer was positioned at the point where the two channels cross. Fluorescently labeled albumin was introduced into one channel and transport through the endothelial monolayer was quantified by laser excitation of the fluorophore in the other channel [48]. In another example, an endothelialized construct was designed with a mechanism to control shear stress experienced by the endothelial cells. The device was used to determine the effect of fluid shear on nanoparticle translocation across the endothelial monolayer. These studies were intended to define the ability of flow and shear stress to model different types of vasculature in terms of the contribution to pharmacokinetics and drug biodistribution [49]. Other microfluidic devices with integrated vasculature have been developed to determine how the atomic structure of drugs and nanoparticles can influence the rate of translocation across an endothelium [50]. The integration of vascular function in organ-on-a-chip microfluidic design shows great promise in providing more accurate modeling of drug pharmacology in next-generation in vitro cell platforms.

Lung-on-a-Chip The lungs, which represent a fluideair interface between the aqueous in vivo environment and the extracorporeal environment, serve as a common port of entry for drugs, toxins, pathogens, and other xenobiotic compounds. Accurate modeling of the lung in organ-on-a-chip systems is likely important for relevant modeling of the effects of agents that enter the circulation through the alveoli. Significant advancements have been made in the on-chip modeling of lung tissue [51]. Many of these lung constructs consist of lung epithelial cells and endothelial cells situated on opposing surfaces of a semipermeable membrane. The cellularized membrane forms a barrier that can model the transport of aerosols or vapors from the gaseous alveolar compartment into the aqueous circulatory compartment. Contact of the alveolar epithelial cells with air in the alveolar compartment has the added advantage of promoting normal cellular function and maintaining the differentiated state. In more complex devices, multiple independent pneumatic channels were incorporated into the design. Cyclic deformation of the pneumatic channel walls paired with controlled shear within the fluid channel promoted exceptional cell morphology and function [52,53]. These advanced models have proven to be valuable for modeling several lung pathologies including inflammation, pulmonary edema, mucus plug rupture, alveolar epithelial cell damage, and advanced drug screening [54e57]. Although planar airefluid interface models have shown incredible promise for modeling exchange across the alveoli and certain pulmonary disease states, more simple, acinar lung organoids may be sufficient for screening drug toxicity for compounds delivered orally or directly into the circulation.

Heart-on-a-Chip Models for cardiac tissue are generally straightforward in design. The heart’s sole function in the human body is to drive the circulation of blood, and most in vitro cardiac models are designed to model this function. Simple monolayer cultures of human cardiomyocytes will beat spontaneously in culture when grown on Matrigel [58]. Sheets of human cardiomyocytes may be layered to produce 3D cardiac constructs that retain the ability to contract in synchrony [59]. These planar construct designs are sufficient for modeling the heart’s beating action but are not ideal for modeling 3D mechanics such as contractile force. 3D cardiac constructs consisting of human cardiomyocytes embedded in collagen I hydrogels molded into ring structures self-organize into circumferentially aligned cell architecture and support physiologically relevant action potential propagation [60]. These types of constructs have been integrated into microfluidic circuits to form dynamic, contractile heart-on-a-chip systems [61e65].

Cancer-on-a-Chip In addition to modeling normal tissue such as liver and heart, excellent models for tumor tissue have been developed. These models have been integrated into microfluidic platforms to form tumor-on-a-chip devices capable of

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modeling both tumor growth and metastasis. The microenvironments of tumors are often more complex than their normal counterparts. The ECM often varies greatly from tumor to tumor, and cells within the tumor display varying degrees of reliance on the tumor stroma. Tumors possess a range of degrees of vascularization and have different structural and regulatory protein compositions within the stroma. The physical and biochemical characteristics of tumors can be monitored and controlled using microfluidic and microfabrication techniques. Tumor microphysiological systems derived from a patient’s tumor sample can be used to determine the selection of an effective chemotherapeutic agent and an optimal dose on a patient-by-patient basis [66,67]. Advancements in tumor-on-a-chip modeling include the development of integrated hardware to monitor the tumor tissue. These include advanced imaging technology and onboard molecular biology assays that provide detailed characterization of tumor behavior, on-chip. The microscale of on-a-chip systems has been shown to influence cell metabolism significantly. This results from the bioavailability of oxygen within these platforms. Studies demonstrated that microfluidic systems provided greater access to oxygen compared with standard 2D culture systems. This increased oxygen level results in increased Krebs cycle activity and decreased expression of hypoxia-regulated factor-1 [68]. Constant perfusion of with oxygenated media provides a better model for a normal tumor microenvironment. A device was designed with multiple drug gradient mixers and parallel cell culture chambers to facilitate multidose drug screens. This system was paired with a cell-labeling strategy and high content-imaging data collection on-chip [69]. Another tumor model design includes microscale bioreactors that contain hepatocytes, nonparenchymal cells, and breast cancer cells. This system is intended to simulate the hepatic microenvironment. The device contains oxygen sensors, micropumps for controlling nutrient distribution, and real-time sampling capabilities [70]. This device was used to demonstrate that breast tumor cells will spontaneously become dormant when placed within the hepatic niche. This effect was postulated to result from the microenvironmental cytokine profile created by the presence of hepatic cells within the liver bioreactor. Breast cancer has also been studied using a developed system that includes both the ductal and lobular components of breast tissue [71,72]. This system is intended to model the interaction among these two microanatomical compartments during tumor initiation and progression. In another device, HCT-116 human colorectal carcinoma cells and HepG2 hepatoma cells were encapsulated in Matrigel in separate chambers. Myeloblasts were embedded in alginate gels within an additional chamber to simulate bone marrow. Using this platform, the cytotoxic effects of the 5-fluorouracil (5-FU) prodrug tegafur could be determined for each cell type. Interestingly, using 3D tumor organoids, the liver constructs were able to metabolize tegafur to 5-FU, resulting in cell death within the other two tumor organoid types. Tumor models constructed in 2D were unable to metabolize the prodrug to its activated form [73]. Lung tumor models have also been developed in microfluidic devices. Human nonesmall cell lung cancer in both 2D and 3D organoid configurations were evaluated for sensitivity to several common chemotherapeutic agents [74]. In another example, lung cancer spheroids were formed from cell lines or derived from patient lung tumor biopsies, both with and without the addition of a pericyte population. Each of these constructs was tested for susceptibility to the drug cisplatin. Systems that included the pericyte population demonstrated higher levels of chemoresistance to the anticancer drug, which indicated the importance of considering all cell types within the tumor when developing an organoid design strategy [75]. These examples of tumor-on-a-chip platforms demonstrate the benefits of 3D organoid models and microfluidic technologies for cancer research as well as patientspecific drug and dose selection.

BODY-ON-A-CHIP: MULTIORGAN SYSTEMS AND FUTURE APPLICATIONS On-a-chip technologies have gained significant momentum. Although these technologies are relatively new, they have shown great promise for applications in research and drug development. However, systems of increased biological complexity have begun to emerge that feature multiple organoids integrated within a single platform [76e79]. These multiorganoid devices [80], sometimes referred to as “body-on-a-chip” systems, have demonstrated greatly increased potential for modeling relevant physiology compared with single organoid systems. Until recently, these multiorganoid platforms were composed of cell lines and animal cells [73,81] More recent systems have begun to use human primary cells or fully differentiated cells derived from stem or progenitor populations. These human platforms have required the development of more advanced cell substrates and microfluidic devices to support primary cell populations. Several notable published studies have demonstrated complex human multitissue systems. In one such system, a four-tissue circuit was developed in a pumpless microfluidic perfusion platform, housing 2D tissue cultures of liver, cardiac, skeletal muscle, and neuronal integrated within

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a single microfluidic circuit. This platform was designed as a screening tool to determine cell toxicity in experiments using doxorubicin, atorvastatin, valproic acid, APAP, and N-acetyl-m-aminophenol [82]. This pumpless concept was also employed in a dual-tissue guteliver system including transepithelial electrical resistance sensors to monitor gut epithelial barrier function [83]. In another study, a dual-organoid microfluidic platform was developed with integrated intestine, skin, liver, and kidney epithelial barrier tissue. In that system, basic cellular function, appropriate gene expression, and viability were maintained for 28 days [77]. These examples represent important steps toward systems that can mimic complex, multitissue responses and interactions during drug and toxicology testing.

The Importance of Multiorganoid Integration In vitro models that recapitulate human tissues and model disease accurately are rare. Fewer still contain multiple tissues integrated in series on a single platform. Such models are required for testing drug toxicity and effects at the organism level, because tissues do not exist in isolation in the body. Moreover, for cells within a microphysiological platform to function normally, it is essential that they receive signals and support, such as vascular, neural, metabolic, and hormonal cues from other cell types. With respect to drugs, effects in secondary tissues can be as important as those at the intended target site, particularly if they induce toxicity. If undetected, detrimental secondary effects can lead to the failure of expensive clinical trials or withdrawal from commercial or clinical use after US Food and Drug Administration (FDA) approval. Multiorganoid platforms are also useful for disease modeling. Cancer metastasis, in which malignant tumor cells are able to migrate to and establish tumors at a secondary site, may be modeled in multiorganoid platforms. Although they are useful for many applications, single-organoid models have a limited ability to model these types of physiologically relevant events. Here, we describe several examples of multiorganoid platforms that demonstrate the importance of these systems. Cancer As described earlier, cancer metastasis is a disease process that may be modeled only in a multiorganoid system. In developing metastatic potential, certain tumor cells gain the ability to intravasate through endothelium into the bloodstream or lymphatic system. They may then migrate to a distant tissue and extravasate into a secondary tissue site. Few in vitro systems have been developed that employ a multiorganoid approach to model the kinetics of metastasis. One system developed by our team demonstrated that it is possible to recapitulate the metastatic process in vitro. This metastasis-on-a-chip platform was designed to enable tracking of the migrating metastatic tumor cells from a bioengineered colon organoid to a bioengineered liver organoid within a simple recirculating microfluidic device (Fig. 44.2A). It was shown that metastatic colorectal cancer cells were able to migrate out of the colon tumor organoid into the microfluidic circuit and engraft in the downstream liver organoid. Conversely, a nonmetastatic colorectal cancer cell type proliferated at the primary site but never migrated to the liver within the study time frame [84]. Tumor metastasis-on-a-chip platforms, as described previously, may be composed of multiple organoids that enable tumor cells to metastasize from a primary location to a secondary site. On-a-chip devices have also been designed to assess certain discrete aspects of metastasis. For example, one system includes a microfluidic device that can model the process by which multicellular tumor aggregates migrate through both a collagen matrix and an endothelial cell layer [85]. Another device includes an endothelial cell layer that partitions a microfluidic circuit from a chamber that houses a 3D bone construct. This system allows modeling of the extravasation of metastatic tumor cells from the vasculature into bone [86,87]. Other devices include a system to assess the effects of interstitial pressure on cell migration [88] and a system for screening antiangiogenic drugs [89]. These systems illustrate the potential benefits that on-a-chip cancer technologies are capable of delivering. However, there is still a major lack of platforms that model both the primary and metastatic sites, as well as the barriers that separate these locations (i.e., basement membranes, circulation, ECM, and endothelium) in a single platform. By providing circulating flow through a system containing both vasculature and multiple organoids, recapitulating the migration of tumor cells from primary tumor organoids into the microfluidic circulation and engraftment into a downstream target organoid may be accomplished. The results of these systems seem to be well-aligned with what is seen clinically. For example, this system has demonstrated that colorectal cancer cells preferentially engraft into liver organoids, a well-established target tissue for colorectal metastatic tumors [84]. These examples represent several components of the metastatic process that have been modeled in multiorganoid systems; future studies will likely rely on these types of platforms to uncover other factors that influence metastasis.

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FIGURE 44.2 Examples of multiorgan interactions that cannot be modeled with single-organoid systems. (A) Migration and metastasis of tumor cells from one organ or organoid site to another, demonstrated in vitro in a metastasis-on-a-chip device in which colorectal carcinoma metastasizes from the colon to the liver. (B) Reliance of a prodrug therapy such as the anticancer 5-fluorouracil prodrug tegafur on liver metabolism to activate the drug to generate a positive effect by targeting tumor cells successfully. (C) Inflammatory molecules secreted from organs such as the liver and lung upon drug injury can cause detrimental inflammatory responses and cell injury in downstream tissues. BPM, beats per minute; IL, interleukin; TNF-a, tumor necrosis factor- a.

Drug Testing and Toxicology As has been discussed throughout this chapter, a major area of interest addressed by the body-on-a-chip field is understanding how multiple organs and tissues respond to the administration of particular drugs within an integrated platform. A variety of examples demonstrate this concept. For example, 5-FU is a common chemotherapy agent employed in treating colorectal cancer. Unfortunately, 5-FU can induce a variety of detrimental side effects in patients, including cell damage in the gastrointestinal tract. In an attempt to reduce toxicity, several prodrugs have been developed, such as tegafur. Tegafur and other prodrugs are inactive in the administered form. The prodrug is activated by hepatocytes in the liver to 5-FU that are able to kill tumor cells. Consequently, without including a metabolically active liver organoid in the system, no active drug would be produced and experimental results would be irrelevant. Including functional liver organoids, along with intended target tumor cells and potential tissues that may experience unwanted toxicity, a more complete understanding of the benefits and risks associated with administration of a chemotherapeutic agent may be determined (Fig. 44.2B). Additional strides are being made toward deploying organoids and organ-on-a-chip technologies in drug and toxicology screening applications. For example, our group demonstrated the use of 3D cardiac organoids in screening for drugs and toxins. We consistently observed expected changes in cardiac beat kinetics in response to these compounds. By employing a minimicroscope with custom-written software to analyze cardiac beating kinetics, the precise determination of beat frequency and magnitude could be recorded [64]. These toxicity screening efforts were expanded to include both cardiac and liver organoids, assessing toxic outcomes from a panel of environmental toxins and a set of drugs that were recalled from the market owing to unanticipated toxicity in human patients (Fig. 44.3). Fig. 44.4 highlights some of these studies, in which the dual-organoid systems were able to model

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FIGURE 44.3 Highly functional organoids for a multiorganoid body-on-a-chip platform. (A) Acetaminophen (toxicity in liver organoids and reduction in toxicity by N-acetyl-L-cysteine. (B) Cardiac organoids remain viable long-term and support transport of fluorescent dyes (lucifer yellow [yellow stain] and fluorescein [green stain]) through interconnected ion channels suggesting high levels of cellecell communication. (C) Beating analysis of cardiac organoids: An onboard camera captures video of beating organoids, after which beating rates are calculated by quantifying pixel movement, generating beat plots. (D) Vascular endothelium devices respond to changes in endothelium integrity as measured by a transendothelium electrical resistance sensor. Organoid diameter - 250 mm.

these toxicities. In addition, the integration of different tissue organoids into a singular system allows for screening studies that can identify unanticipated toxicities (Fig. 44.2C) [90]. Additional Disease Modeling Research into human pathologies other than cancer may also benefit from the capabilities of multiorganoid systems. Many drugs and that are known to cause inflammatory responses. For example, large doses of the common analgesic APAP (Tylenol) causes significant inflammation and toxicity in the liver. Other drugs, including

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FIGURE 44.4 Drug and toxicology screening in liver and cardiac organoids. Live/dead stains show the presence of dead cells (red) after treatment with lead, mercury, troglitazone, or astemizole. Organoid diameter - 250 mm.

chemotherapeutics such as bleomycin, cause inflammation, toxicity, and irreversible fibrosis in the lungs or other organs. In most of these cases, toxicity and apoptosis lead to the release of proinflammatory cytokines such as tumor necrosis factor-a and interleukin-1 into the circulation. These molecules can cause a series of downstream responses, including recruitment of inflammatory cells, activation of fibroblasts, and changes in vascular protein expression and permeability (Fig. 44.4C). Integrated multiorganoid model systems, including integrated vasculatures, can model these complex, multiorgan drug responses in a reproducible and physiologically relevant manner, whereas single organoid systems cannot.

Cutting-Edge Body-on-a-Chip: The First Highly Functional Multiorganoid Systems As described previously, there is a paucity of truly integrated multiorganoid platforms that are able to model and test the complex responses accurately to drugs, toxins, and disease across a range of human tissue types. However, progress is being made. The Ex Vivo Console of Human Organoids Platform Our team has developed an advanced, modular, multiorganoid integrated, body-on-a-chip system for use in drug development and toxicology screening. The multiorganoid body-on-a-chip platform is named Ex Vivo Console of Human Organoids (ECHO). This platform was initially developed to include four engineered tissue organoid types (liver, cardiac, vascular, and lung), which were developed independently and integrated into a single system that provides real-time monitoring of physiological responses to toxic agents and pharmaceuticals. Using the ECHO platform, a comprehensive set of data was developed providing a characterization of each organoid. In general, the 3D organoids are bioprinted into platform microreactors using tissue typeespecific supportive hydrogels. The result is an array of organoids suspended in a substrate containing tissue-specific, ECM-derived bioregulatory factors [36,91]. Liver organoids are fabricated using liver ECMederived hydrogels that maintain viability and function in vitro for 4 weeks [91]. The production of important liver markers (e.g., albumin, multiple cytochrome P450 enzymes, epithelial cellecell adhesion markers, dipeptidyl peptidase IV, organic solute transporter-a, etc.) have been confirmed and are stable over a month. These organoids respond to toxins such as APAP at appropriate doses and in a dose-dependent manner. The organoids may also be rescued by the clinical countermeasure for APAP intoxication, N-acetyl-L-cysteine (Fig. 44.4A). Cardiac organoids also remain for 4 weeks and beyond. These organoids demonstrate the transport of fluorescent dye molecules among cells within the organoids, indicating a high degree of cellecell communication. The organoids beat spontaneously and change their beating rates appropriately in response to a variety of drugs. These kinetics are captured using an onboard camera system [92,93] and custom software for analysis (Fig. 44.4B). In addition, an engineered vasculature has been incorporated into the platform that responds to agents such as histamine by disrupting the endothelial cell monolayer

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(Fig. 44.4C). This results in increased transendothelial transfer of largeremolecular weight molecules that are normally sequestered within the microfluidic circuit. Several integrated, multiorganoid studies were performed using the ECHO platform. Experiments including integrated liver organoids, cardiac organoids, and endothelial modules in microfluidic devices (Fig. 44.5A and B) under common media demonstrated long-term viability and produced multiorganoid responses to drugs that largely mimic the responses that would be expected in an experimental animal, or even in a human. For example, Fig. 44.5C describes the effects of propranolol and epinephrine on cardiac organoids, both with and without integrated liver organoids in the microfluidic platform. Normally, epinephrine induces an increase in the beating rate in cardiomyocytes. Without liver, propranolol, a b-blocker, blocks the b1- and b2-adrenergic receptors, preventing an increase in the cardiac beat rate after epinephrine administration. However, in the integrated system that includes the liver organoids, propranolol is metabolized to an inactive form, resulting in the recovery of much of the epinephrine-induced increase in beat rate. To our knowledge, these experiments are the first interdependent multiorganoid studies performed successfully in a single integrated system. The ECHO platform has been used to screen drugs that were withdrawn from the market owing to unanticipated toxicities. Because of the lack of accurate models to predict drug toxicity, many drugs have passed through preclinical studies and clinical trials, received FDA approval, and remained on the commercial market for years in some cases, before being recalled because they cause toxic effects in humans. Approximately 90% of drugs that were removed from the market were because of toxic effects in the liver and the heart. A panel of these drugs were tested in the ECHO platform. These include the drug troglitazone (Rezulin), an antidiabetic and antiinflammatory that was recalled for causing liver failure, and mibefradil, an ion channel blocker that was recalled for having fatal interactions with other drugs, including antibiotics. In our platform, troglitazone and mibefradil both result in liver toxicity. We also screened the drug rofecoxib (Vioxx), a nonsteroidal anti-inflammatory drug that was recalled because it caused serious vascular-based pathologies such as heart attack, stroke, skin reactions, and gastrointestinal bleeding. Astemizole, an antipsychotic that caused slowing of potassium channels, torsade de pointes, and QT prolongation. We also tested terodiline, a drug for bladder incontinence that caused QT prolongation and toxicity, in ECHO and showed functional changes and loss in viability among several cell types within the platform. The anticancer drug 5-FU and isoproterenol, a b-adrenergic agonist, both of which are known to induce cardiac toxicity, were evaluated in ECHO. Each of these drugs resulted in increased levels of cell death within the cardiac organoids in a dosedependent manner. Using the onboard camera, beating effects were observed to decrease with dose increases as well. Effects on beat kinetics were detected at doses well below the toxic threshold. This is an important point, because drugs withdrawn from the market for cardiac toxicity are generally not withdrawn for killing cells in the heart, but rather for causing changes in heartbeat kinetics. Other Body-on-a-Chip Programs In addition to the ECHO platform initiative, a variety of other high-profile programs have been working to develop integrated systems. In particular, the Advanced Tissue-Engineered Human Ectypal Network Analyzer (ATHENA) program designed a millimeter-scale multiorganoid system, and a program sponsored by the Defense Advanced Research Projects Agency (DARPA) sponsored a 10-organoid project [94]. These projects have stressed aspects of microphysiological systems that are often different from the main goal of ECHO (i.e., organoid integration within a common microfluidic circuit). Specifically, the ATHENA program, which is based in Los Alamos National Laboratory, developed a system composed of four organs: liver, heart, lung, and kidney [95]. These organoids are three orders of magnitude larger than the ECHO organoids. This scale allows for more relevant mechanical testing and the system fluid volume is sufficient to collect samples that may be analyzed using standard clinical diagnostic equipment. The DARPA program, based in Harvard’s Wyss Institute, is working with a collection of 10 organoids, including representations of endocrine, gastrointestinal, immune, musculoskeletal, and reproductive tissues [96]. In addition, the NIH are supporting a major organ-on-a-chip program through the National Center for Advancing Translational Science, the National Institute for Biomedical Imaging and Bioengineering, the National Cancer Institute, the Eunice Kennedy Shriver National Institute of Child Health and Human Development, the National Institute of Environmental Health Sciences, the NIH Common Fund, and the NIH Office of Research on Women’s Health. However, the NIH initiative differs in that the funding is divided among a variety of individual research laboratories; between them, they are working on organoid representations of a wide range of tissue types [97]. The

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(B)

(A)

(C)

(D)

i)

PMMA

Inlet Liver

Outlet

Heart

DST Brain

Testis

Lung Vascular

PMMA

Membrane Liver Testes

ii) Heart Brain

Lung Blood vessel

Glass slide

FIGURE 44.5 Multiorganoid body-on-a-chip. (A) Depiction of a liver, cardiac, and vascular organoid-containing body-on-a-chip platform. (B) Photograph of the three-organoid system. (C) Description of the effects of propranolol and epinephrine on cardiac organoids, with or without liver organoids, illustrating the importance of multiorganoid systems. Without liver, propranolol, a b-blocker, blocks cardiac beating increases by epinephrine. However, with both organoids present, propranolol is metabolized by the liver organoid, resulting in a measurable epinephrineinduced increase in beating rates. (D) (i, ii) Future body-on-a-chip platforms for increased capabilities for linking multiple organoids within a single circulatory system. BPM, beats per minute; DST, double-sided tape; PMMA, poly(methyl methacrylate).

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program stresses the derivation of organoids from induced pluripotent stem cells (iPSCs), and many of the organoids developed by members of this program are excellent models for specific tissue types. However, the potential for integrating these organoids into a multiorganoid system remains to be seen. Organ-on-a-Chip Systems for Personalized Precision Medicine Numerous in vitro systems are being developed for general drug development screening, but few have been developed to benefit specific patients. This is an unmet clinical need, because finding the most effect drug and dose for a specific patient is often a trial-and-error process. With personalized organoid models (Fig. 44.6A and B), therapies can be screened using a patient’s own cells in a 3D tissue organoid system. For example, accurately predicting a patient’s tumor progression and response to therapy is one of the most challenging aspects of oncology. Prescribed treatments are often made based on the general success rate of a drug within a population, not on the specific response that may be expected within an individual. The concept of precision, or personalized, medicine has evolved to address these problems by using the patient’s genetic profile to identify “druggable” targets for treatment [98e100]. However, in real-world practice, the results of this approach to personalized medicine do not reach the desired goals [101]. After identifying key mutations through genetic profiling, physicians can still be left with a variety of drug options, with no concrete data regarding potential side effects or actual drug effectiveness in the patient. As such, there is a clear need to develop tools that can predict the response of individual patients to drugs [102,103]. Our group is working to develop a multiorganoid platform that contains patient-specific tumor organoids, and in which drug therapies that are selected based on genetic profiling can be tested for efficacy. By using

FIGURE 44.6 Employing biofabricated tissues in personalized medicine. (A) In personalized precision medicine for cancer patients (red

arrows), a potential list of drugs is determined based on mutations found in the tumor genetic profile, from which best-guess therapies are prescribed. In the future, cells from tumor biopsies could be used to create in vitro tumor models specific to a given patient (green arrows). Potentially effective drug therapies can then be screened in the models, thus identifying the optimal drug therapy for that patient in terms of both safety and effectiveness. (B) In genetic diseases, cells can be harvested from alternative tissues, such as skin, translated into induced pluripotent stem cells (iPS), differentiated into cells of the tissue of interest (e.g., lung or heart), and bioengineered into three-dimensional (3D) organoids and organoid-on-a-chip systems, after which generic and genome-specific drug therapies can be screened for the original patient.

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microfluidic devices, a circulatory system containing multiple tissue organoids may be fabricated. This may allow for the prediction of tissues to which metastatic cells will migrate [84]. By combining all of these capabilities described, it may soon be possible to generate accurate prognoses regarding tumor progression and metastases and to develop an informed treatment plan that is personalized to each patient’s tumor. In the near future, iPSC technology may allow for the generation of an array of patient-specific normal tissue organoids that could provide information about the sensitivity of these tissues to a specific treatment regimen. This would enable the selection of an agent that reduces stress on the healthy tissues within a cancer patient. These exciting and powerful new in vitro technologies may soon revolutionize the treatment of cancer and other diseases.

CONCLUSIONS AND PERSPECTIVES Although the benefits of multiorganoid systems are clear, several challenges remain regarding their acceptance and widespread deployment for drug development and personalized medicine. Most single celletype systems do a respectable job in mimicking a few specific aspects of in vivo physiology and are amenable to highthroughput drug screening. The few single- and multiorganoid platforms, including multiple cell types within each organoid, provide a much better model for human physiology but have not yet been optimized for high throughput [104]. As such, these more complex systems are best-suited for evaluating drugs that are in the later stages of development. Many groups are actively developing strategies for multiorganoid systems and automating their production [14]. A significant reduction in size, automated fabrication, improved onboard biosensing, and in-line diagnostic technology would greatly increase the throughput potential for multiorganoid platforms across all potential applications [92,93,105,106]. Another perceived challenge in the development of advanced, multiorganoid, platforms is the requirement of a common cell medium to support a wide variety of cell phenotypes. Typically, human primary cells and iPSC-derived cells require complex, highly specialized media formulations that are tailored to each specific cell type. Surprisingly, there is growing evidence that 3D cell constructs are intrinsically supportive and much less reliant on complexed media supplements or serum. Our group has repeatedly demonstrated the maintenance of a variety of cell types in human cancer models [84,107] using serum-free medium and customized hydrogen substrates [36,91,108]. Even more remarkable is the ECHO platform described earlier in this chapter. In this platform, up to six organoid types, each containing up to five human primary cell types, have been maintained under a serum-free common medium for at least a month with minimal loss in viability and function. It is likely that the organoids themselves produce the autocrine and paracrine factors that are required for long-term viability and function. The compact architecture, scant interstitial space between cells, and accumulation of ECM proteins would allow these factors to become concentrated within each organoid and diminish reliance on exogenous factors delivered through the medium. Multiorganoid body-on-a-chip technology is advancing at a rapid pace and is likely be deployed soon for drug screening [10]. These platforms have significant utility in many areas and will dramatically change the way in which precision medicine, cancer modeling, and drug development are performed.

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Skardal A, et al. Tissue specific synthetic ECM hydrogels for 3-D in vitro maintenance of hepatocyte function. Biomaterials 2012;33:4565e75. Kim SB, et al. A mini-microscope for in situ monitoring of cells. Lab Chip 2012;12:3976e82. Zhang YS, et al. A cost-effective fluorescence mini-microscope for biomedical applications. Lab Chip 2015;15:3661e9. Reardon S. Scientists seek ’Homo chippiens’. Nature 2015;518:285e6. Roark K. In: Los Alamos national laboratory press release, vol. 2016; 2015. The economist; 2015. National Institutes of Health news releases; 2014. Tran NH, et al. Precision medicine in colorectal cancer: the molecular profile alters treatment strategies. Ther Adv Med Oncol 2015;7:252e62. Miles G, Rae J, Ramalingam SS, Pfeifer J. Genetic testing and tissue banking for personalized oncology: analytical and institutional factors. Semin Oncol 2015;42:713e23. Bando H, Takebe N. Recent innovations in the USA National Cancer Institute-sponsored investigator initiated Phase I and II anticancer drug development. Jpn J Clin Oncol 2015;45:1001e6. Hayes DF, Schott AF. Personalized medicine: genomics trials in oncology. Trans Am Clin Climatol Assoc 2015;126:133e43. Cantrell MA, Kuo CJ. Organoid modeling for cancer precision medicine. Genome Med 2015;7:32. Gao D, et al. Organoid cultures derived from patients with advanced prostate cancer. Cell 2014;159:176e87. Esch EW, Bahinski A, Huh D. Organs-on-chips at the frontiers of drug discovery. Nat Rev Drug Discov 2015;14:248e60. Kim SB, et al. A cell-based biosensor for real-time detection of cardiotoxicity using lensfree imaging. Lab Chip 2011;11:1801e7. Shaegh SAM, et al. A microfluidic optical platform for real-time monitoring of pH and oxygen in microfluidic bioreactors and organ-on-chip devices. Biomicrofluidics 2016;10:044111. Skardal A, Devarasetty M, Rodman C, Atala A, Soker S. Liver-tumor hybrid organoids for modeling tumor growth and drug response in vitro. Ann Biomed Eng 2015;43:2361e73. Skardal A, et al. Bioprinting cellularized constructs using a tissue-specific hydrogel bioink. J Vis Exp 2016;21:e53606.

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45 Bioreactors in Regenerative Medicine Jinho Kim, Kelsey Kennedy, Gordana Vunjak-Novakovic Columbia University, New York, NY, United States

INTRODUCTION The goal of regenerative medicine is to restore the structure and function of damaged, diseased, or lost tissues and organs. Most human tissues have limited regenerative capacity after injury and heal with scar formation or inflammation [1]. Accordingly, approaches to regenerative medicine use tissue-specific cells, biomaterials, and biologically active molecules to enhance regenerative responses and promote the formation of matured and functional tissues. In particular, tissue constructs and whole organs can be produced in vitro by providing physiologically relevant culture conditions and by subjecting the cells to biomimetic physical and biochemical cues [2]. By closely recapitulating the native environments of tissue development, bioreactors can provide conditions for growing fully functional and viable tissues and organs for transplantation in vitro (Fig. 45.1).

DESIGN CONSIDERATIONS FOR CREATING BIOREACTORS When appropriate conditions are provided, cells can survive and grow, and eventually form tissue units outside the body. Conventional two-dimensional (2D) cell culture protocols enable the efficient generation of cells with controllable properties and phenotypes, but simple cell assemblies grown on 2D surfaces would be insufficient for applications in regenerative medicine, which typically require three-dimensional (3D) and functional tissue constructs. Using a bioreactor, physiologically functional 3D tissues can be bioengineered in vitro by recreating the microenvironmental and macroenvironmental niche conditions to cells seeded into scaffolds such as the extracellular matrix (ECM) [3]. To this end, several physical and biological aspects must be considered in designing bioreactor systems to produce tissues or even whole organs with desired physiological properties. To promote cell viability, proliferation, and differentiation, it is essential to provide adequate amounts of oxygen, nutrients, and biochemical factors to the cells or tissues cultured in a bioreactor. In addition, metabolic waste generated by the cells must be removed effectively to further facilitate cell and tissue growth. As the cell density and tissue size increase, mass transfer through bioengineered constructs becomes challenging. Under static culture conditions, diffusion is the major mechanism for the transport of nutrients, most critically oxygen. However, diffusional transport is efficient only within a superficial layer (usually about 100e200 mm) of the 3D tissue constructs, which can be at least several millimeters to centimeters thick. Accordingly, to enhance the efficiency of gas and mass transfer, convective flow can be generated within bioreactors by stirring, rotating, or perfusing culture medium [4e6]. Mechanical forces can regulate the physiology of cells, in particular osteocytes, chondrocytes, cardiomyocytes, and skeletal muscle cells, via mechanotransduction pathways by which mechanical signals are converted into biochemical signals [7]. Thus, tissue formation and development can be promoted by mechanical stimuli, motivating the use of various mechanical stimulation regimes to regenerate tissues and organs with desired properties. For example, tissue-engineered cartilage with enhanced functionality was produced by applying physiological levels of compression to hydrogels seeded with articular chondrocytes [8,9]. Similarly, enhanced cell proliferation and tissue organization were achieved by repeatedly stretching scaffolds embedded with heart cells [10] or skeletal muscle cells [11]. Shear forces induced by fluid flow increased osteogenic expression and chondrogenic responses [12]. Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00045-X

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FIGURE 45.1

Schematic of tissue or organ regeneration in vitro using bioreactor systems.

Cell behaviors such as proliferation, differentiation, and migration can also be influenced by electrical stimulation in electrically excitable cells [13]. In particular, when appropriate electrical stimulation is applied, cardiomyocytes cultured in vitro, combined with biomaterials such as collagen, can differentiate and form into functional cardiac tissue constructs [14]. Development of contractile and electrophysiological function in bioengineered tissue constructs can be facilitated by cyclic electrical stimulation and mechanical strain. Tissue contractions resulting from the stimuli can promote cell elongation, orientation, organization of connexin-43, and the formation of structural features similar to those of native myocardium. The electrical stimulation also increases the expression of myosin heavy chain, creatine kinase-MM, and cardiac troponin-I in the bioengineered cardiac constructs. Notably, the positive effects of the electrical stimulation on tissue development strongly depend on the time of its initiation, which indicates that the timing of stimulation must be carefully determined to bioengineer cardiac tissues [15]. In addition to physical stimuli, regeneration of tissues and organs of interest requires the addition of biochemical cues that can modulate cellular and tissue behaviors [16], in particular, soluble biochemical factors such as growth factors and cytokines that are provided to local cells through secretion from neighboring cells or through supply from the blood instruct cells during development and tissue regeneration [17]. Because the concentration and spatial gradient of the biochemical factors strongly influence the microenvironment that dictate the fates of the cells, tight control of the biochemical parameters within the in vitro system is essential to obtain tissue and organs for specific applications. In the body, cells reside in a complex 3D microenvironment that is composed of tissue- and organ-specific cell types, the ECM, and physical and biochemical factors Accordingly, the regeneration of tissue and organs for transplantation requires 3D scaffolds, in which cells seeded are provided with a biomimetic niche microenvironment that can promote the formation of functional tissue or reconstruct viable organs [18]. Various synthetic ECM materials such as poly(ethylene glycol) and poly(vinyl alcohol), which are typically soft and porous and have a high water content, have been employed as scaffolds that can facilitate the organization of cells into a 3D architecture by presenting the necessary structural, physical, and biochemical characteristics [19]. On the other hand, biological scaffold materials obtained from the native ECM have been shown to promote regeneration of functional tissues greatly [20]. Although the mechanisms mediating the cell behavior by the natural ECM are not well-understood, ECM derived from many different tissues such as heart, blood vessels, skin, and tendons have been tested in both preclinical and clinical applications [18]. Methods that allow monitoring of the tissue growth and maturity, and generation of tissue- and organ-specific functionalities are needed for the successful applications of bioengineered tissues and organs in regenerative medicine [21]. Typical assays that are used to analyze the structure and biochemical composition of the tissue, such as histology and microscopy, involve tissue-sampling procedures, which is destructive in nature. On the other hand, advanced imaging techniques such as molecular imaging and deep-tissue imaging techniques allow noninvasive assessment in vitro and enable real-time and long-term evaluation of bioengineered tissues and organs.

LUNG BIOREACTORS Bioengineering Functional Lungs For many patients with end-stage lung disease, lung transplant is the only definitive treatment option available. Although timely organ transplant is important to increase the life span and overall life quality of the recipients, this treatment procedure is significantly hampered by the severe shortage of viable donor lungs [22]. Furthermore, less than 20% of donated lungs are acceptable for transplant whereas the rest are unsuitable because of poor lung

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conditions caused by various reasons including contusion or edema [23]. Accordingly, to increase the number of lungs suitable for transplant, tissue-engineering methods are being developed to salvage donor lungs that are marginally unacceptable and rejected for transplant. In these approaches, endogenous cells in the donor lungs are replaced with fresh cells to restore the lung function to a level acceptable for transplant. Important processes for lung regeneration using tissue-engineering technologies include (1) removal of the cellular components from the donor lungs (i.e., decellularization), and (2) repopulation of the resulting lung scaffold with the new cells such as lung airways cells or stem/progenitor cells (i.e., recellularization) [24]. The 3D lung scaffold provides organ-specific environment with structural and biochemical cues for the transplanted cells to facilitate lung remodeling and regeneration. In particular, functional ECM components preserved in the decellularized lung matrix, such as collagens, elastin, fibronectin, laminin, and glycosaminoglycans, can promote cell engraftment and differentiation during recellularization. For the accelerated regeneration of viable and functional lungs, a bioreactor is an essential component that can offer suitable environments to the bioengineered lung [25]. For instance, normothermic (37 C) perfusion and sterile support of the lungs during reconditioning can be achieved by using a bioreactor. An ideal bioreactor would provide mechanical and biochemical cues to the lungs mimicking physiological stimuli during the native lung development. Furthermore, physically and biologically relevant conditions including the aireliquid interface and breathing movement should be generated in the lung bioreactor to promote precisely regulated differentiation and organization of the transplanted cells in the lung scaffolds created by decellularization.

Bioreactors for Regeneration of Small Animal Lungs Because of their relatively easy access, lungs harvested from small animals such as rodents [26,27] have been widely used to establish protocols for lung tissue engineering (Fig. 45.2A). Accordingly, various lung bioreactors

FIGURE 45.2 Bioreactors for ex vivo regeneration of small-animal lungs. (A) Schematic for lung regeneration process [26]. (B) Photograph of bioreactors for (i) decellularization and (ii) recellularization of rat lungs [28]. (C) Schematic of a bioreactor for selective cell replacement in the rat lungs [29]. De-epith, de-epithelialization.

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designed for small animals have been evaluated for the long-term preservation of explanted lungs during cell-based organ regeneration via decellularization followed by recellularization (Fig. 45.2B). It is imperative that the lung bioreactors provide physiologically relevant conditions (e.g., sterile conditions, isothermic perfusion, and air ventilation) to the lungs throughout the regeneration procedures. While lungs are kept in a bioreactor, chemical or biological decellularization agents can be perfused sequentially through the airways and vasculatures to remove entire cellular components from the lungs [30]. This perfusionbased decellularization requires precisely controlled fluid flow to maximize preservation of the structural integrity and biological activity of the lung ECM. Decellularization agents include detergents such as 3-([3-cholamidopropyl] dimethylammonio)-1-propanesulfonate, sodium dodecyl sulfate, and Triton X-100 and enzymes such as nuclease, trypsin, and collagenase. After cell removal, the lungs are seeded with therapeutic cells (e.g., epithelial and endothelial) through the airway and vasculature. During recellularization, the lung is ventilated and perfused, respectively, via ventilator and pump integrated into the bioreactor, recapitulating physiological lung developmental conditions (Fig. 45.2C). A novel bioreactor for the rat lungs was developed in which one lobe of the lung could be bioengineered while the other one remained intact to serve as a control [29]. This bioreactor allowed removal of the epithelial cell layer (de-epithelialization) from the airways of the small animal lungs (i.e., rat) while the underneath lung tissue was preserved (Fig. 45.3D). The selective cell removal was achieved by introducing a mild decellularization solution through the airway while the solution entering the blood vessels was quickly cleared by continuously perfusing the vasculature. As a result, the vasculature network in the lung could be maintained undisrupted, allowing structural and biological integrity of the bloodegas barrier that facilitated bioengineering of functional lungs after reepithelialization.

Bioreactors for Regeneration of Large Animal Lungs To generate bioengineered lungs on a clinically relevant scale, bioreactors have been developed that could maintain the function of large-animal lungs (e.g., porcine and human) during ex vivo regeneration [31,33,34] (Fig. 45.3A and B). Similar to bioreactors for small animal lungs, these clinical-scale bioreactors allow sterile, normothermic perfusion and ventilation to promote lung regeneration and functional recovery. In addition, large-lung bioreactors have been further optimized to support the lungs for a sufficient time to allow diagnostic and therapeutic interventions that would be needed to improve marginal lungs. Automated lung perfusion and ventilation with integrated sensors (e.g., flow, pressure sensors) and lung monitoring components (e.g., video cameras) into the bioreactor would be desirable to achieve more tightly controlled lung regeneration. Whereas ex vivo lung perfusion (EVLP) has typically been used to preserve donor lungs during transport, its utility for assessment and reconditioning of marginally unacceptable donor lungs has also been pursued (Fig. 45.3C). Benefits of using EVLP include the recruitment of collapsed lung areas, removal of bronchial secretions, and clearing of clots from the circulation. In addition, because the lungs are connected to the EVLP system externally, direct and targeted treatments of the lungs are possible via endotracheal and intravascular routes. Owing to the absence of other organs in the circuit, higher doses of therapeutics for accelerated disease treatment can be administered to the lungs supported by EVLP (Fig. 45.3C). However, the lack of other organs also implies that the metabolic clearance and systemic factors are absent. Nevertheless, EVLP systems have served as excellent platforms for human lung support and bioengineering [35]. The adaptation of EVLP for reconditioning marginal donor lungs increased the total number lung transplants at least by 15% [36].

In Vivo Bioreactors for Lung Regeneration Allogeneic solid organ transplants typically require nonspecific and lifelong immunosuppressive therapy to prevent and treat organ rejection. Although immunosuppression improves graft survival, the patient becomes susceptible to infections, in particular to pulmonary infections that may cause morbidity and mortality [37]. An effective approach to reducing the immunogenic response would be using the recipient as an “in vivo bioreactor,” in which donated organs or tissues are implanted heterotopically, thus taking full advantage of the natural regenerative capacity of the recipient’s body [38]. This elegant method was demonstrated for trachea (Fig. 45.4A), with the donor trachea implanted subcutaneously into the recipient’s forearm for over 9 months. During this time, the neovascularization occurred and the endogenous mucosal lining was partially replaced with the recipient’s own buccal mucosa. After transplantation of the regenerated trachea, the immunosuppressive therapy was suspended and the recipient’s

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(A)

(B)

Ventilator

(C) Temperature probe Pressure probe

Reservoir Leukocyte filter

O2, N2, CO2 Oxygenator

Temperature probe Heater/ cooler

FIGURE 45.3 Bioreactors for ex vivo regeneration of large-animal lungs. (A) Schematic layout of a clinical-scale lung bioreactor and

(B) photograph of the bioreactor setup in an incubator [31]. (C) Schematic drawing of an ex vivo lung perfusion (EVLP) unit [32]. LA, left atrium; P.Eq., pressure equilibrium; PA, pulmonary artery; PEEP, positive end-expiratory pressure; PV, pulmonary vein.

buccal mucosa gradually grew over the graft. Significantly, the allogeneic trachea was stable and functional for several years even after immunosuppression was discontinued [39,41]. Smaller tissues or organs such as trachea and bone segments can be heterotopically regenerated within the recipient’s own body before orthotopic transplantation. However, this in vivo approach can be difficult for the regeneration of solid organs such as whole lungs or liver owing to their large size and the need for perfusion. A potential solution for this can be cross-circulation (extracorporeal circulation), in which donor organs are connected externally to the recipient’s circulatory system during organ recovery or repair [42,43]. Using a preclinical swine model, this novel cross-circulation approach has been demonstrated for prolonged lung support during therapeutic interventions undertaken to recover the injured donor lungs [40] (Fig. 45.4B and C). Because the donor organ is placed outside the recipient’s body, the organ is more accessible for diagnostic and therapeutic interventions that can improve the

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(A)

(B) (C)

In vivo bioreactors (A) Tracheal allotransplant after revascularization in a heterotopic position (forearm) of the patient [39]. (B) Cross-circulation circuit for regeneration of donor lungs ex vivo supported by the recipient and (C) operative setup for the cross-circulation [40]. H, warm water jacket; I.C, incision; I.S., immunosuppression; IJ, internal jugular vein; PA, pulmonary artery; PV, pulmonary vein; TC, thermal camera; TL, time-lapse camera; VB, video bronchoscopy.

FIGURE 45.4

organ conditions, and minimally invasive evaluation of the lung function recovery. In addition, lung regeneration through cross-circulation can be accelerated when combined with the cell-replacement therapy using the patient’s own cells, because the recipient provides the necessary biochemical cues needed by the transplanted cells.

Bioreactors for Study of Lung Biology Lung regeneration and repair involve complex processes for which the underlying mechanisms are not fully understood [44]. More detailed study of lung development can enable the establishment of effective treatment modalities for a number of respiratory complications including immature lungs in preterm babies and inflammation-induced disruption of alveologenesis. In addition, such study would also inform strategies for stem cellebased ex vivo lung regeneration [45]. Microengineered lung models in microfluidic culture devices (i.e., “lung-on-a-chip”) have great value for high-throughput in vitro studies of the lung at the molecular, cellular, and tissue levels [46e48] (Fig. 45.5A). Thus, these biomimetic lung-on-a-chips can recapitulate the structure, function, and microenvironment of human lung, permitting the study of lung development, regeneration, disease, and responses to drugs.

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FIGURE 45.5 In vitro (iv) devices for studies of lung development and regeneration. (A) A lung-on-a-chip model for microengineering small airways [48]. (B) Lung bud organoids grown in vitro in a hydrogel [49]. (C) Bioreactor for lung organoid generation [50]. (D, E) Bright-field microscopy of a lung organoid [51].

When provided with suitable biochemical and physical conditions, stem cells can be formed into 3D tissue structures (i.e., organoids) that contain self-organized cell clusters. For example, human pluripotent stem cells cultured in 3D hydrogel (e.g., Matrigel) formed lung bud organoids resembling branching human airways and early alveolar structures [49] (Fig. 45.5B). Whereas model systems such as 2D monolayer cell cultures, 3D cell spheroids, and tissue explants have limitations in recapitulating the complex and multilevel human physiology, organoids can faithfully recapitulate many biological parameters, providing a unique opportunity to perform more physiologically relevant studies in vitro. Using custom-built bioreactor systems, generation of organoids can be facilitated in terms of organoid number and size, establishing the use of organoids in applications of drug testing, regenerative medicine, and disease modeling [51] (Fig. 45.5CeE).

Evaluation of Bioengineered Lungs Various imaging and tissue processing techniques are used to evaluate the bioengineered lungs and assess their structural and biochemical integrity [21]. For example, removal and repopulation of cells such as epithelium of the

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FIGURE 45.6 Evaluation of the microstructure and composition of the lung tissue. (A) Hematoxylineeosin and immunofluorescence

staining of lung tissue samples of airways. (B) Scanning electron microscopy (SEM) of native and de-epithelialized (De-epith, De epi) lungs and transmission electron microscopy (TEM) of alveoli. (C) Pressureevolume measures of lungs. (D) Microscopic imaging of fluorescent microbeads (diameter ¼ w1 mm) in the lung capillaries [29]. (E) Transpleural imaging of labeled microbeads or mesenchymal stem cells (MSCs) in the lung parenchyma [40]. (F) Imaging of MSCs seeded onto the rat trachea [52]. (G) X-ray and thermal imaging of the porcine lung during regeneration [40]. TR, trachea.

lung can be visually assessed by hematoxylin and eosin staining of the tissue samples obtained from the lung. In addition, immunofluorescence staining can reveal the presence or absence of biological molecules in the lung tissue (Fig. 45.6A). Detail about cell and tissue structures can be probed by using powerful imaging techniques such as scanning electron microscopy and transmission electron microscopy, which provide high-resolution images [29] (Fig. 45.6B). To evaluate the lungs without the need for tissue sampling, nondestructive and real-time assessment have been employed [29,40,52]. A simple method to determine the quality of the bioengineered lung is to measure the pressureevolume relation of the lung, which can provide lung compliance and resistance (Fig. 45.6C). In addition, the structural integrity of the blood vessels can be evaluated by directly monitoring traveling paths of microbeads (diameter of approximately 1 mm) added into the blood flow (Fig. 45.6D). Similarly, cells added into the lung parenchyma (Fig. 45.6E) or airways (Fig. 45.6F) can be visually inspected using a custom-built microscope and opticalfiber imaging probe, respectively. Airway recruitment and blood perfusion of the lung during regeneration can be inspected using X-ray and thermal imaging (Fig. 45.6G).

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BONE BIOREACTORS Bioengineering Bone Bone tissue has an inherent ability to regenerate, except in critical-sized defects caused by congenital abnormalities, trauma, or disease. Autologous grafts are the reference standard treatment, but these grafts are limited by the tissue volumes that can be harvested and patient morbidity. Engineered bone substitutes rely on bioreactors to provide nutrients throughout the large tissue constructs and to deliver the mechanical stimuli needed to promote bone development [53]. In this section, we examine bioreactor design for bone tissue engineering, including an advanced example of clinical translation, and the important role of bioreactors in delivering dynamic loads to cells within the engineered bone. Like lung, bone is a vascularized organ, and we summarize efforts to use the body as its own bioreactor to mitigate the challenge of vascularizing implants. Finally, we consider the use of bioreactors as platforms for biological studies of bone development and disease, and how nondestructive modalities may be used to monitor bone structure and function within bioreactor environments.

Nonperfused Bioreactors for Bone Regeneration Static tissue culture relies on diffusion to deliver nutrients and biophysical stimuli to tissue, resulting in cell viability and matrix deposition only at the border of a 3D construct. To drive nutrient distribution in bone tissue engineering, two types of bioreactors for dynamic culture have been introduced: rotating wall vessels that use concentric, rotation cylinders to create laminar flow; and spinner flasks that use a magnetic stirrer or similar device to create flow through constructs suspended in the flask. Both setups have shown improvements in cell viability and distribution compared with static culture; however, in both cases, flow remains restricted to the periphery and necessary nutrients and signaling cannot reach deep within critical-sized scaffolds [54].

Perfusion Bioreactors for Bone Regeneration Perfusion bioreactors, in which media are pumped throughout the construct, have shown improved cell distribution and viability and matrix production throughout large 3D scaffolds [55,56]. Perfusion bioreactors pump media from a reservoir throughout a construct within a liquid-tight vessel. Flow may be directed around the outside of a construct, or it may be forced directly through a construct by press-fitting the scaffold within the bioreactor casing. The latter setup offers superior mass transport, although flow patterns need to be customized to fit the construct geometry [57]. Perfusion bioreactors have been used for engineering bone from various stem cell sources, including embryonic [58], induced pluripotent [59], bone marrowederived [60], and adipose-derived stem cells [61,62]. Bone marrowederived and adipose-derived mesenchymal stem cells (MSCs) have shown particular promise for clinical translation, because both have demonstrated success in perfusion-bioreactor engineering of large-animal, anatomical bone replacements [60,61] (Fig. 45.7). These custom-shaped bioreactors relied on preoperative imaging of the targeted defect from which a mold was fabricated. Two elastomer pieces were then molded and press-fit onto a decellularized bone scaffold, which was milled to the anatomical shape and loaded with stem cells (Fig. 45.7A). To ensure optimal mass transport and shear stress throughout the construct, computational flow simulations were used to determine the channel size and placement (Fig. 45.7B). The engineered bone closely resembled the resected, native bone after 3 weeks in culture (Fig. 45.7C) and guided generation of new bone over 6 months in vivo in adult pigs (Fig. 45.7D). In contrast, soft issue formed in the condylectomy control (Fig. 45.7E). Custom perfusion bioreactors may be extended to implement multiple chambers for engineering adjacent tissues, such as engineering the articular cartilage surface on bone to form whole joints [63,64]. An important advantage of perfusion bioreactors is the ability to control the media flow throughout engineered bone, not only for mass transport but also to introduce appropriate shear stresses to cells for bone development. Previous reviews reported on the mechanobiology of bone development and the use of bioreactors to recapitulate physiological shear stresses [53,65]. Shear stresses are critical to bone formation and resorption via mechanotransduction in osteocytes, which communicate via paracrine signaling with osteoblasts and osteocytes to regulate matrix turnover [66]. Shear stress was also shown to direct osteogenic differentiation of MSCs [67]. Several studies attempted to recapitulate the physiological stresses seen by cells in native bone using perfusion bioreactors [56]. However, because no method exists for measuring the locally applied shear stress directly at the cellular level in 3D, studies rely on computational simulation of flow rates to determine local stresses. An important study established predictive correlates between the perfusion rate and the local osteogenic behavior of human MSCs in

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FIGURE 45.7 Engineering anatomical bone grafts using a perfusion bioreactor. (A) The bioreactor consists of two polydimethylsiloxane blocks with incorporated channels, two manifolds, and a casing, providing controlled perfusion through the scaffold. (B) Channel diameters and spacing were designed by flow simulation software to provide a desired interstitial flow velocity for a given shape and size of graft. Simulations revealed a uniform flow velocity throughout the volume. (C) Comparison of resected native bone and engineered graft shows closely matched geometry. (D) Movat pentachrome stain of explants from pigs undergoing condylectomy showed fibrous tissue filling the defect; (E) engineered scaffolds guided development of new calcified tissue in the defect [61].

decellularized bone matrix [57]. Computational simulations yielded the independent contributions of oxygen delivery and shear stress to the resulting osteogenic behavior, including gene expression, matrix deposition, and cell proliferation. A range of flow rates from 400 to 1800 mm/s resulted in stresses of 2e20 mPa, respectively; although this is more than an order of magnitude lower than stress values reported in vivo [53], stresses were expected to increase as more matrix was produced and pore size correspondingly decreased. Although many perfusion bioreactors introduce steady, continuous flow to engineered bone, there may be advantages to mimicking the oscillatory or pulsatile loading experienced by bone tissue in vivo [55]. External loads are applied to cells via a combination of matrix strain and interstitial fluid flow through lacunar-canalicular spaces [68]. A study examined the direct effects of oscillatory flow on MSCs using a parallel-plate bioreactor setup, in which the pressure is spatially constant throughout the setup, as confirmed by computational fluid dynamics (Fig. 45.8A). By testing various load magnitudes and frequencies, a regime of 2 Pa at 2 Hz was identified to have the greatest upregulation of osteogenic expression in MSCs [69], similar to the loads experienced by human cells in vivo during jogging. Beyond fluid flow, external intermittent compression may be applied for mechanosignaling for bone development. A compression bioreactor setup, including an array of platens driven by a linear actuator and interfaced with a 24-well plate, was used to introduce periodic strain cycles and investigate mechanical regulation in a bone tumor model [70] (Fig. 45.8B). Computational simulation of the local pressure and flow velocity throughout the porous decellularized bone matrix (modeled as a poroelastic material) allowed the correlation of local pressure and matrix strain with gene expression. A new paradigm for dynamic loading in a bone bioreactor was reported in which nanoscale vibrations were introduced to MSCs via a magnetic plate [71] (Fig. 45.8C). Simulations of the displacement within the constructs revealed a constant force magnitude across the construct surface, and MSCs undergoing nanovibrations generated mineralized matrix despite the absence of physiological matrix stiffness

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FIGURE 45.8 Bone bioreactors with advanced loading regimes and corresponding mechanics simulations. (A) Oscillatory flow bioreactor (left) and stress field simulations (right) [69]. (B) Compression bioreactor (left) and simulated stress field (right) [70]. (C) Vibrational bioreactor, simulated nanoscale displacements of the magnetic plate (top) and six-well plate (bottom) used for culture, and resulting constructs without stimulation, with mechanical stimulation, and with osteospecific media (OSM) [71].

(soft collagen gel served as a matrix material). These diverse approaches to loading bone tissue in bioreactors showed that mechanical effects on bone development must be considered in designing bioreactors.

In Vivo Bone Bioreactors for Solving the Vascularization Problem One of the greatest challenges in producing viable engineered bone substitutes is the generation and anastomosis of a functioning vascular network. Like lung, bone is a highly vascularized tissue and efforts to incorporate vasculature during in vitro development have fallen short of promoting vascular integration upon implantation. A potential approach to overcome this is to use the body itself as a vascularized bioreactor and develop the bone in an ectopic location before orthotopic implantation [72e74]. This “in vivo bioreactor” principle has the additional advantage that the body provides a further supply of stem cells and growth factors to the construct, avoiding overmanipulating

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FIGURE 45.9 In vivo bone bioreactor. (A) The in vivo bioreactor paradigm: ectopic implantation prepares a construct for successful orthotopic implantation [73]. (B) Layered nanofibrouseceramic constructs implanted subcutaneously in nude mice to promote vascularization (left). Hematoxylineeosin histology shows that vascularization occurred all the way through the core of the scaffold [75]. TMJ, temporomandibular joint.

the tissue ahead of implantation, which could simplify the clinical translation of engineered bone replacements [73] (Fig. 45.9A). Since it was originally introduced, in vivo bioreactor studies have taken advantage of the highly vascularized environment in the periosteum, the connective tissue layer surrounding the bone [69]. The successful formation of mineralized, compact bone was demonstrated by 6 weeks after implantation, in rabbit periosteum. Given the relatively invasive nature of this implantation, subcutaneous implantation has also been explored for its vascularization potential. To promote vascular ingrowth, a layered construct was engineered using MSCs within electrospun polycaprolactone nanofibers, an angiogenic layer composed of collagenefibronectin gel with endothelial cells, and an osteoconductive layer containing hydroxyapatite particles. After 4 weeks of subcutaneous implantation in nude mice, vasculature was present throughout the core of the scaffold [75] (Fig. 45.9B).

Bioreactors for Studying Bone Development and Disease Bioreactors for bone were originally introduced to improve mass transport to cells when culturing in large 3D constructs. However, the opportunity to introduce a controlled mechanical and biological environment to realistic 3D bone constructs has allowed sophisticated models of bone development and disease within bioreactors. For instance, advanced mechanobiology studies may be carried out in perfusion bioreactors. A study examined the synergistic effects of matrix stiffness and shear stress on the osteogenic differentiation of progenitor cells. Decellularized bone matrices were coated with various ratios of collagenehydroxyapatite to vary matrix stiffness and were loaded with MSCs [76]. The authors modeled flow through the constructs to estimate shear stresses and found that oscillatory flow is needed to promote cell viability and osteogenic behavior in long-term culture, whereas matrix stiffness had to be optimized separately for the greatest osteogenic potential. The controlled environment provided by bioreactors could also be used to study fracture repair under different biological regimes. For instance, static culture of hypertrophic chondrocytes improved long bone repair via endochondral ossification in the rat model, in contrast to osteoblast-laden scaffolds cultured in a perfusion bioreactor that promoted intramembranous ossification [77].

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Bioreactors have also shown advantages in modeling the bone tumor microenvironment to understand tumorigenesis and drug efficacy better [70,78]. A study that used a dynamic compression bioreactor showed that mechanical loading affects the tumor’s drug sensitivity in an engineered model of Ewing sarcoma [79]. Effects of shear stresses on tumor cell homing to bone during metastasis have also been studied. Spinner flask culture upregulated interleukin-24, which was shown to decrease the viability of prostate cells in engineered bone compared with static coculture [80]. The role of compression in inhibiting tumor-mediated osteolysis and metastatic tumor formation in bone by breast cancer cells was observed in a compression bioreactor for bone engineering [81]. Bioreactors also enable the modeling of joint disease for drug screening applications [82]. By scaling down and simplifying the bioreactor design, the osteochondral interface has been studied “on a chip” using microfluidics to introduce appropriate media to the two tissues [83]. A similar dual-compartment design was implemented to grow osteochondral tissues from human MSCs, and interleukin-1b induced tissue degeneration [84]. This platform is suitable for testing disease-modifying osteoarthritis drugs.

Monitoring the Environment and Tissue Development Within Bioreactors The abilities to sense and visualize the inner workings of bioreactors are crucial for determining their faithfulness in recapitulating the physiologic environment and for evaluating proper tissue development. In the context of bone bioreactors, several structural, mechanical, and biological assessment tools have been proposed to monitor tissues and cells nondestructively. Microcomputed tomography (mCT) is a nondestructive imaging modality that visualizes tissue structure in 3D. Because of the high absorption of X-rays by mineralized tissue, it may be used to quantify formation of bone longitudinally and has been a primary modality for this application [85]. Some studies have performed mCT imaging directly within perfusion bioreactors [86,87] without the need to remove the construct from its media circuit and sterile environment. Such longitudinal imaging can provide feedback to 3D tissue formation during culture, enabling the adaptation of biophysical stimuli as needed. Direct mCT monitoring within a perfusion bioreactor also allowed the relation of flow rate to the formation of mineralized tissue by human MSCs in a decellularized bone scaffold [88]. Longitudinal mCT has been used to investigate the effects of the degree of matrix curvature on new bone formation in 3D silk fibroin constructs [86] (Fig. 45.10A). Scaffolds with channels of varying degrees of curvature, to mimic the matrix turnover in vivo, were cultivated with cells in perfusion and static bioreactors to study fracture healing. Optical imaging, including bioluminescent and fluorescent imaging, offers molecular specificity and may be performed nondestructively to investigate cell behavior within bone bioreactors. Bioluminescence has been used to track stem cell behavior in bone defects in vivo in mice (Fig. 45.10B). MSCs were transfected with two variants of luciferase, one used as a constitutive promoter to quantify cell number and another as a cell differentiation reporter for osteoblast lineage [89]. This approach is a promising avenue to understand stem cell fate better in engineered bone. Noninvasive, optical assessment of oxygenation in a bioreactor has also been achieved by measuring the oxygen-dependent phosphorescence lifetime of microprobes within the media of a perfusion bioreactor [91]. The authors found a decrease in oxygen within tissue constructs under static culture conditions, in contrast to stable oxygen levels in the perfusion setup. However, the technique provides no spatially resolved information; it offers a single oxygen measurement at each time point. Another disadvantage of optical modalities is a limited depth penetration, which restricts imaging to thick, opaque constructs. Owing to both shear stresses and matrix strains, mechanical effects on bone development have received increased attention, but many studies lack a method to visualize the effects of mechanics directly within dynamic bioreactor environments. A bioreactor system was demonstrated that is capable of real-time mechanical conditioning (dynamic, uniaxial strain, and electrical stimulation) of nanofibrous constructs and the simultaneous monitoring of local strains [90]. Fluorescent beads and MSCs were electrospun directly into fibers, and fluorescence microscopy along with digital image correlation techniques were used to map local strains (Fig. 45.10C). Such a system may provide new insights for bone mechanobiology and allow tracking of the mechanical development of bone tissue in future studies. Elastography, which uses medical imaging including ultrasound, magnetic resonance imaging (MRI), and optical coherence tomography, may also have a role in mechanically conditioned engineered tissues [92].

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FIGURE 45.10 Image-based monitoring of tissue and cell behavior in bone models. (A) microcomputed tomography monitoring of bone deposition in silk fibroin scaffolds with and without perfusion. Color-coded overlays of different time points show areas of new growth and resorption [86]. (B) Bioluminescent imaging of human mesenchymal stem cell (MSC) differentiation in demineralized bone scaffolds implanted in a mouse model. MSCs were transduced with reporters for cell number (renilla luciferase [RLuc]) and/or cell differentiation (photinus luciferase [PLuc]) toward osteoblast lineage (OC), endothelial lineage (human platelet endothelial cell adhesion molecule [hPECAM]), or hypoxic expression (HRE) [89]. (C) Optical monitoring of local strain in nanofibrous constructs in a mechanical bioreactor. Fluorescent beads were incorporated into the fibers to track local displacements [90].

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CHALLENGES AND FUTURE DIRECTIONS Promising advances have been made in designing bioreactors that deliver essential nutrients and environmental cues to drive tissue development. However, several areas require improvement to enable the clinical translation of bioreactor technologies for regenerative medicine. As discussed here for lung and bone tissues, vascularization is critical for graft survival upon implantation. Using ectopic locations in the body as an in vivo bioreactor can mitigate some of the issues, but it may not always be practical. Other bio-inspired strategies should be considered that could be carried out in vitro to prime tissues for vascularization and contribute to success upon implantation. For example, harnessing the inflammatory response to promote angiogenesis has been explored in bone tissue constructs. Biomaterial-directed switching of the macrophage phenotype enabled upregulation of angiogenic factors and improved vascularization in a mouse model. Many bioreactor studies do not account for the inflammatory milieu, but it may be necessary to incorporate these cell types for optimal tissue development in the future. In this chapter, we have introduced techniques that have been used to monitor the bioreactor environment and tissue development, but much work is needed in this space to interface bioreactors further with advanced imaging and sensing modalities. Although several modalities have been used to monitor cell fate and tissue development in vitro in monolayer culture [93], relatively few have been extended to monitoring the 3D complex environment within a bioreactor. Part of the challenge is the trade-off between the resolution needed to capture cellular-scale events with the field of view needed to monitor large, 3D constructs. MRI-compatible [94] and computed tomographyecompatible bioreactors have been developed to monitor media flow and tissue structure, but they cannot track individual cells and have not been widely adopted for laboratory use, likely owing to the cost and size of the systems. Optical techniques including bioluminescence, fluorescence, multiphoton techniques, and optical coherence tomography may have an increasing role in real-time monitoring in future studies [93]. Real-time and nondestructive assessment of tissue and organ regeneration will be essential to the eventual automation of bioreactor control. Imaging and sensing readouts may be used in a feedback loop to signal inputs of environment cues (e.g., mechanical actuation, oxygenation) or delivery of biological factors. Toward this end, a computer-controlled perfusion bioreactor with integrated sensors for oxygen and pH was used to grow bone successfully for implantation in mice [95]. Eventual clinical translation will also require standardization and quality control; here, predictive computational models and automation are expected to contribute. Computational analyses and machine learning techniques may enable further the optimization of culture regimes, resulting in improved reproducibility in engineered construct quality.

Acknowledgments The authors gratefully acknowledge the National Institutes of Health support of the work described in this chapter (Grants DE016525, EB025765, HL120046, HL134760, and EB002520).

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46 Bioinks for Three-Dimensional Printing in Regenerative Medicine Javier Navarro1,2,a, Gisele A. Calderon3,a, Jordan S. Miller3, John P. Fisher1,2 1

University of Maryland, College Park, MD, United States; 2Center for Engineering Complex Tissues, College Park, MD, United States; 3Rice University, Houston, TX, United States

INTRODUCTION The advent of three-dimensional printing (3DP) is revolutionizing approaches to fabricating tissue mimics for therapeutic replacement, drug discovery, and fundamental biological understanding. The potential niche for 3DP in tissue engineering is seemingly infinite, because we have at hand the ability to provide on-demand, fabricated, patient-specific designs rapidly and at a low cost. However, 3DP technology was first intended for industrial settings. The translation to tissue engineering applications is hindered by major hurdles that include technical printing issues and, more important, biocompatibility. Consequently, we are limited by the number of materials available that can satisfy both the 3DP and compatibility requirements. Here we provide a historical perspective of 3DP and summarize the different techniques, consider the important characteristic properties a bioink must have to fit bioprinting criteria, summarize bioink and biomaterial advancement used for 3D bioprinting, and discuss future directions to address current limitations for clinical impact. Generally, printable biomaterials, or bioinks, are materials that can be used in 3DP techniques that include or will include biological features. The term “bioink” may lead to confusion because some may consider the material a bioink only if it is cell-laden or contains some matrix or matrix-mimicking component. However, we would like to expand its definition to encompass any printable material that (1) will interface with biological components (e.g., tissues, cells, proteins, growth factors) during or after the actual print, or (2) is involved in the structural construction of scaffolds that will interface with biological components. Bioinks must comply with the 3DP technique as well as provide a biocompatible environment mimicking a desired tissue and ideally degrade controllably with no harmful by-products. Harmful by-products may not originate exclusively in degradation; they may also come from temporary bioinks that have structural roles during the printing process. Unfortunately, to satisfy these criteria, material properties work against each other and require some compromise between desired printability and satisfactory biological features. However, bioengineers have a selection of materials compatible with several different 3DP platforms. Therefore, in our discussion we will cover bioinks in the context of which printing technique is capable of printing with the described material. We will describe the diversity of available bioinks and biomaterials for 3DP under three main categories: (1) matrix or matrix-mimicking, (2) sacrificial, and (3) support [1].

FUNDAMENTALS OF THREE-DIMENSIONAL PRINTING 3DP is an additive manufacturing technique originally applied to plastic and metal manufacturing. It has progressed to adapt to biomedical engineering applications within the past few decades. Over 3 decades ago, Charles a

J.N. and G.A.C. contributed equally to the production of this text.

Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00046-1

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Hull patented a new technology for rapid prototyping, called stereolithography. This described system to fabricate three-dimensional (3D) constructs uses liquid photopolymerization to build desired objects in a stepwise manner and is considered the birth of 3D bioprinting [2,3]. 3DP and bioprinting have revolutionized the ability to create objects with any shape or size on demand. The exciting promise for medicine is that we can customize patientspecific tissue scaffolds, fabricate on-demand medical devices, and reliably reproduce constructs for highthroughput screening. The key term that defines modern bioprinting is control. Casting approaches have surface resolution defined by the mold itself and no control over the internal structure of the casted sample. Similar salt leaching or electrospinning protocols for porous structures have little to no control over the internal pore size and distribution [4]. Modern technologies allow for control deposition and building; there is control over detailed characteristics such as the location and content of material deposition. This includes microstructures to bear stress, cell encapsulation, and growth factor or other biochemical functionalization. Being able to design the bulk shape and internal microstructure fully has allowed structures to be produced that more closely resemble nature’s complexity: controlled internal microstructures (pores, gradients, and layering) [5e10], microchannel and micropatterning for prevascularization [11,12], and the simultaneous deposition of different types of materials [12e14] or different cell lines [6e8,12,15]. The technology allows for control and the mechanics are there to exploit; however, the hurdle to overcome is to define the bioinks that can be adequately used with these systems. An ever-expanding list exists of 3D bioprinting technologies (and bioinks to match) that are differently suited for specific desired materials or applications. These techniques vary based not only according to the materials (e.g., affected by cross-linking mechanisms to determine biocompatibility and/or mechanical properties) but to the resolution of the architecture and the speed of fabrication. Four 3DP methods have been adapted for biomedical application: (1) extrusion-based printing, (2) particle fusionbased printing, (3) inkjet printing, and (4) stereolithography or photopolymerization methods (Fig. 46.1). Technological advances have resulted in novel methods or modifications of these four, including acoustic droplet ejection, direct-write assembly, laser-guided direct writing, and 3D powder printing [3]. Each of these approaches has advantages and limitations. However, given an identified desired application, often one or a combination of technologies proves to match the sought biomaterial and architecture. A 2016 review by the Kaplan Lab [3] summarized the evolution of bioprinting and additive manufacturing technologies.

Extrusion-Based Printing Extrusion-based printing is the most commonly employed method for 3DP, but it requires thermoplastic materials that are cell compatible only if they are printed at physiological temperatures. The technology forces a viscous ink

FIGURE 46.1 Three-dimensional (3D) bioprinting techniques: Extrusion-based printing deposits continuous filaments on a surface to build constructs layer by layer. Inkjet printing adapts inkjet cartridges to contain bioinks so that liquid droplets are formed on a surface that can quickly solidify. Selective laser sintering uses a laser to melt powder particles together to sinter a 3D object within the powder bed. Finally, projection stereolithography uses photosensitive liquid material that can be cross-linked with light exposure with controllable photomasks. Reprinted with permission from Miller JS, Burdick JA. Editorial: special issue on 3D printing of biomaterials. ACS Biomater Sci Eng 2016;2:1658e61. Copyright 2017 American Chemical Society.

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through a nozzle that can solidify once deposited onto a build platform. The extruded material is deposited to form individual lines in a predefined path as dictated by a generated computer model to form a 3D object in a layer-bylayer fashion. Typically, each different ink is extruded out of the nozzle at a specific temperature and pressure so that the material can flow through the nozzle (e.g., polycaprolactone PCL] will extrude at 80 C and a pneumatic pressure of 400 kPa, whereas alginate can be deposited at 20 C [7]). Materials that are most commonly used for this technique possess a sharp solid-to-melt transition such that the material can flow and rapidly solidify once passed through the nozzle. Progress has been made in developing thermoplastics that extrude at lower, more physiological temperatures and pressures by exploiting shear thinning properties (i.e., PCL) [17]. Other approaches include using materials that can be extruded at lower temperatures or do not rely on thermal setting but require additional cross-linking mechanisms such as ionic bonding [18], pH alteration [19], or UV photopolymerization, among others. The inclusion of cells, however, can ultimately affect several aspects of the resulting print. Not only do materials need to be able to flow through the nozzle at low temperature, parameters such as high extrusion forces and narrow nozzle diameters, which usually improve the print’s resolution for synthetics at high temperatures, affect cell viability. Decreasing the pressure and/or increasing the nozzle diameter may improve cell viability by reducing the cells’ experienced shear stress but at the cost of potential nozzle clogging and print resolution. In addition, this technique has difficulty printing overhanging structures without supporting filler structures. The printed filaments can sag or even collapse without underlying support. Although some groups used extrusion printing with PCL or poly(propylene fumarate) (PPF) to create highly controlled porous scaffolds [20e22], others generated structures as a sacrificial template with similar extrusion-based methods to pattern architectures precisely that can be better-suited for some of the biocompatibility issues with this 3DP technique [11,23]. Overall, extrusion-based printing has been applied to many tissue engineering applications and is most widely available.

Selective Laser Sintering Another printing technique, particle fusion printing methods such as selective laser sintering (SLS), uses a directed laser to raise the temperature of a powder material beyond its melting temperature to fuse or sinter the particles locally. The laser fuses each patterned layer in a repeated process to create 3D structures [24]. The construct’s resolution is limited only by the laser resolution and the powder size. This technique has had its most successful applications to bone tissue engineering by sintering composite materials. Examples include ceramics, PCL, and hydroxyapatite [25e28]. Although the tissue engineering field has seen promise with this technique, the printed objects are difficult to incorporate with living cells and instead use a two-step process of first building the volume followed by adding cells. Moreover, there is difficulty in controlling porosity, which can affect the longevity of the biomedical application [3]. However, a major benefit is the technique’s capacity to be self-supportive of complex printed structures, as the powder bed enables the printing of overhanging geometries. In addition, similar to extrusion-based printing approaches, groups use the technique of SLS additively [28] and sacrificially [29].

Inkjet Bioprinting Inkjet bioprinting is considered the cheapest bioprinting technique. It was patented as one of the first strategies for cell printing [30,31]. Originally, inkjet bioprinters were created by replacing traditional 2D ink-based printers’ cartridges with biological solutions. Instead of paper used as the recipient of the ink, a moveable stage added a third dimension to build a construct layer by layer [32,33]. Inkjet bioprinters can deposit picoliter droplets with positional accuracy less than 30 mm, allowing high-precision control in positioning different materials and cells into specific microenvironments [34]. To deposit liquid onto a substrate, some inkjet printers electrically heat the print head to a range between 200 C and 300 C. The high heat raises concerns regarding the viability and function of the cells after printing. Still, some studies using this technique show viability for mammalian cells and attribute survival to the short duration of the high-temperature exposure [35,36]. In inkjet printing, a liquid droplet is solidified after deposition onto a substrate, a process that must occur quickly to control the spatial resolution of the printed volume. Important material properties of the ink are the viscosity and the surface tension to determine the final shape, size, resolution, and accuracy of the print. Therefore, an important mechanism to consider to improve the final print is the cross-linkability of the bioink. Effective methods for this application are chemical, pH, or UV cross-linking [37,38]. There is a delicate balance of cross-linking the deposited droplets for rapid structural organization and to limit toxicity introduced to the cells. The chemical modification to achieve the desired cross-linking capability can decrease cell viability and affect the chemical and mechanical properties of the material. These changes can alter

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46. BIOINKS FOR 3D PRINTING IN REGENERATIVE MEDICINE

the biomimicry that was relevant before modification [39]. In addition, inkjet printing uses a nozzle head to deposit the bioinks and therefore is limited by the possibility of clogging. Bioinks for inkjet printing should have low viscosities (below 10 cP) because the necessary pressures or heat that would eject higher viscosities would negatively affect cell viability [40].

Stereolithography The last 3D printing technique to be discussed, stereolithography, brings us to the origins of 3D bioprinting. Stereolithography-based printing uses light as a tool to solidify liquid materials in a photochemical reaction. With a laser often used as the light source, a projection of light is shown onto the liquid material in a specific pattern to solidify the exposed region [41,42]. Using a light source, photosensitive material, and a controlled axis stage, one can print complex 3D structures. Because stereolithography depends on photosensitive material to print constructs, a limited number of biomaterials can be used. Common materials for this application are poly(ethylene glycol) diacrylate (PEGDA)-based materials and gelatin methacrylate (gelMA). Acrylation modification to PEG and gelatin render the material photosensitive for printing [43,44]. Combinations of already self-assembling proteins, such as keratin or decellularized extracellular matrix (ECM), with photoinitiators have allowed the cross-linking of soft hydrogels [45]. This method has the advantage of excellent structural integrity because no artificial interfaces result from droplets (inkjet printing) or lines (extrusion printing). Also, the electromagnetic spectrum and its multiple energy wavelengths allow for a broad range of chemistry alterations; various wavelength lasers, UV sources, and the visible light spectrum are common energy sources used in this technique. The use of light with all available intensities and wavelengths results in very fast and precise builds. Resolution of this technique ranges from 25 to 200 mm for commercially available printers [3,45] down to w10 mm for two-photon polymerization setups [46]. However, because the material must be photosensitive, many biomaterials cannot be used or they must be chemically modified for photopolymerization [45,47]. A photopolymerizable material such as PEGDA is not in itself good for cell viability. Additional surface modifications are necessary to allow for cell attachment and material degradation by incorporating peptide sequences [48]. In addition, resolution is determined mostly by the laser spot size and therefore has high 3D resolution, but the prints often are warped because the mechanical properties are typically weak [3]. Overall, 3D bioprinting techniques vary in approach and can result in a wide array of medical applications from tissue repair to modeling disease. There has been tremendous progress in the development of the technique as well as the biomaterials synthesized to expand the palette of available 3DP bioinks. 3D bioprinting has exciting translational potential to produce implantable structures for regenerative medicine and high-throughput, reproducible drug screening. However, to realize this medical impact, researchers must continue to explore the architecture, the biocompatible yet printable materials, and the inclusion of proliferating and differentiating cells for fabricated living tissues to reach a desirable function.

BIOINKS 3D bioprinting may offer the potential to fabricate physiological tissue mimics; however, progress toward therapeutic application relies heavily on its integration with bioinks. Therefore, the development of biocompatible yet printable bioinks requires tremendous consideration to match physical and functional aspects of the desired tissue closely. Because 3DP technologies originally were designed for nonbiological applications, some of the materials used as inks for printing, such as thermoplastic polymers, ceramics, and metals, cannot translate to supporting living cells. Hence, one of the greatest challenges of the field is to find materials that are both biocompatible and printable. As defined previously, printable biomaterials, or bioinks, encompass any printable material that (1) will interface with biological components during or after the actual print, or (2) is involved in the structural construction of scaffolds that will interface with biological components. The key difference between bioinks and other printable inks is cytocompatibility; bioinks cannot be toxic or produce any toxic by-products that could be detrimental to living cells, to the physiological function of the printed tissue, or to surrounding or down-stream native tissues. Generally, materials used in the field of regenerative medicine are divided between natural and synthetic materials. Natural materials have all of the advantages of being physiological and inherently bioactive. However, natural materials lack tunability, batch-to-batch consistency, and often the physical properties necessary for printing. On the

809

BIOINKS

TABLE 46.1

Bioink Materials Compatible With Associated Printing Techniques

Natural

Material

Extrusion

Fibrin

[23,51]

Collagen

[23,54,55]

Alginate

[6,7,14,18,23,57e60]

Hyaluronic acid (HA)

[62], Thiolated HA plus thiolated gelatin [63]

Gelatin

[58,62,64,65]

Keratin

Stereolithography

Inkjet

Sintering

[52,53] Multiphoton cross-linking [9]

[53,56] [61]

Acrylated HA [3] [66] [45]

Agarose

[67]

Hydroxyapatite

[68]

Carbohydrate glass

[11]

Modified natural (semisynthetic)

Gelatin methacrylate

[69], Hybrid with gellan gum [70]

Methacrylated HA

[50]

Synthetic

Poly(ε-caprolactone)

[7,20,21,75,76], Hybrid with starch [14,77], co-printed with polyurethane [60], hybrid with hydroxyapatite [78]

Poly(glycolic acid)

[79], Hybrid with hydroxyapatite [78]

Poly(ethylene glycol) (PEG) or PEG-diacrylate

[80], PEG with reactive ends hybrid with multiple proteins [81]

Pluronic F127

[12,86]

Poly(propylene fumarate)

[22]

[10,87,88]

Poly(vinyl alcohol)

[89,90]

[3]

[3,26]

[44,57,71,72]

[73] [73,74] [27,29]

[3,67,82e84]

[85]

other hand, synthetic materials benefit from a high degree of tailoring to specific physical property needs with inherent consistency to meet the printing technique’s criteria. Still, synthetic materials often fail to match the biocompatibility of natural materials, and sometimes lead to toxic degradation products or the lack of cell-binding sites. Some groups have compromised the divided material set by synthesizing a semisynthetic class of materials such as gelMA or methacrylated hyaluronic acid [49,50]. Of our palette of biomaterials, only a subset is also suitable for bioprinting (Table 46.1). For a bioink to be biocompatible as well as printable, the material must have the capacity to be accurately and precisely deposited with spatial and temporal control. Each bioprinting technique may require a different subset of material properties. For example, inkjet printing requires bioinks to possess low viscosity to avoid nozzle clogging; extrusion printing benefits from shear thinning properties to fluidize through the nozzle and quickly solidify once deposited; SLS must be able to become a fine powder and have an attainable melting temperature; and stereolithography requires photosensitive bioinks. These material properties to promote printability often come at the cost of compromising biocompatibility. As an example, a photosensitive material that cross-links in the presence of a photoinitiator that can be highly cytotoxic. Therefore, available bioinks are chosen to meet the demands of the particular printing process, but also for its ability to shield encapsulated cells from a possibly harmful printing process. The requirements for printing depend on a variety of properties, including rheological behavior, the gelation process, or available biological interactions. From a rheology perspective, only specific ranges of viscosities match well with either inkjet or extrusion printing, but shear thinning is an example of a rheological property ideally suited for extrusion. The gelation process, or cross-linking, can greatly influence the geometric integrity of the print. Gelation can occur through ionic, thermal, enzymatic, or photocross-linking mechanisms; these ultimately dictate the printing technique with which the bioink is compatible. Biological interactions might need to be enhanced, especially for synthetic materials, by incorporating cell binding motifs or inclusion of an additional natural material.

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46. BIOINKS FOR 3D PRINTING IN REGENERATIVE MEDICINE

Bioink categories (from left to right): A desired final geometry volume can be fabricated with three different bioink approaches to result in the ultimate final print. The print setup shows how different bioinks would be incorporated during the print fabrication.

FIGURE 46.2

We categorize the available bioinks into three categories: (1) matrix or matrix-mimicking, (2) sacrificial, and (3) support [1]. As seen in Table 46.1, the same bioink can fall into several bioprinting strategies even if the bioink is used in a different way. Each category requires specific workflows, but they are ultimately brought together in the printing process, as illustrated in Fig. 46.2.

Matrix or Matrix-Mimicking Bioinks Matrix or matrix-mimicking bioinks are printed and remain part of the scaffold system. A scaffold can be printed and consist of the matrix bioink material only (acellular scaffold), the matrix material with its surface chemically altered during or after the printing process (functionalized acellular scaffold), or the matrix printed with a loaded cell population (cell-laden scaffold). In all cases, the matrix bioink is the material that provides the mechanical structure to which cells adhere, and which will then be used to enhance cellular communication, proliferation, migration, and differentiation, and ultimately determine the function of the system. As a structural element in a biological environment, there is a delicate balance between achieving the rheological and mechanical properties needed to print a self-supporting structure and the eventual effects that the material may have on the biological development of the cellular component. This balance is specific to the properties and function ultimately desired for the printed sample, anything ranging from soft and porous hydrogels for the in vitro culture and assessment of cells [65,81,91] to very strong and durable scaffolds for in vivo bone regeneration [10,76,92,93]. The desired mechanical and biological properties of the matrices are nevertheless restricted by the capacities of the 3DP technologies and the bioinks associated with each method. As has been reiterated in the literature, synthetic materials can be engineered to provide strong scaffolds with tunable mechanical and chemical properties [76]. Nevertheless, these materials have been traditionally associated with low biocompatibility [91], complex and demanding manufacturing processes (high temperatures, high pressures, strong solvents, etc.) [34,94], very low degradation rates [76,77], and, in some cases, cytotoxicity or harmful by-products [94]. On the other hand, natural materials inherently provide the adequate biological cues that cells need for proper development. The perfect combinations of amino acid sequences, protein ratios, growth factors, and cytokines are found in natural materials from fauna (ECM combinations, collagen, elastin, fibrin, keratin, hyaluronic acid [HA], chitosan, etc.) and flora (alginate, agarose, agar, silk, etc.) [3,91,95,96]. However, natural materials are also associated with weaker mechanical

BIOINKS

811

properties and high batch-to-batch variability [7,46]. Natural variability is unavoidable and perfectly defined manufacturing protocols for natural materials are virtually impossible. Thus, truly tunable properties are difficult to predict and the reason why results in studies using ECM usually differ in a case-to-case scenario. As will be discussed in the next sections, synthetic and natural materials cannot be defined by positive and negative characteristics; researchers have been working on mitigating the weaknesses of both, either by biochemically altering the individual materials or by implementing syntheticenatural hybrids or combined prints that use the strengths of each. Synthetic Materials As mentioned, 3D bioprinting is an adapted technology; the original patent filed by Hull in 1986 [2] proposed using the stereolithographic method to optimize prototype manufacturing of plastic parts for industrial applications. Since then, the processes have evolved and revolutionized the industry and have bled into many other manufacturing applications including bioengineering, regenerative medicine, and tissue engineering. The original Hull patent was intended for synthetic materials, named “U- curable materials,” which could be processed as a “fluid medium capable of solidification in response to prescribed stimulation” [2]. This definition is technologically viable today and could be applicable to most modern bioprinting methodologies, even if the solidification is not via UV cross-linking. The materials have greatly evolved and new ones have arisen, allowing researchers to incorporate synthetic materials and printing technologies into biomedical and tissue engineering. The greatest strength of synthetics is that the manufacturing processes are well-known and can be engineered to specific mechanical and biochemical properties [91,94]. Polymer engineering allows control over molecular weights and distributions, as well as cross-linking densities, which can be tailored to control mechanical properties such as yield stress and strain, ultimate stress and strain, and elastic modulus [94]. This tailoring can occur as part of the polymer synthesis process, but it can be further modified in the printing or postprinting processes with curing or cross-linking steps. In the end, robust mechanical properties can be used to sustain high loads or adequately respond to elastic deformation, which is ideal for the structural scaffolding components of biological constructs. The synthesis of tunable mechanical properties also means that synthetic materials can be used in multiple 3DP techniques (Table 46.1) and result in constructs with consistent macroscopic and microscopic definition. The pore distribution in a biological scaffold is an important parameter that will define the presence or absence of vascularization for oxygen, nutrient, and metabolic waste transport in tissue regeneration [93]. The superior print resolution and fidelity of synthetics have been widely explored to produce complex morphologies that may be applied as biomimicking scaffolds in regenerative medicine (Fig. 46.3) or as structural supports in co-printing applications, a concept that will be further detailed in following sections. In theory, just as the synthetic print resins can be modified to facilitate the manufacturing processes, surface modifications can be implemented to allow better interactions with biologic components. However, the manufacturing and modification processes and variables are usually demanding and work in narrow ranges to achieve specific properties. Often, printing techniques will involve high temperatures, toxic organic solvents, or cross-linking agents, which renders them incompatible with living cells and biological materials such as growth factors and proteins that aid cellular function and survival [94]. Synthetic materials generally do not support cell adhesion without additional surface functionalization for adhesion ligands such as arginine-glycine-aspartic acid, which is widely identified as a binding motif for proteins such as fibronectin, osteopontin, and fibrinogen [45,91,94]. Even with consistent morphology control and compatibility with surface modification strategies, synthetic polymers do not innately mimic ECM, which remains its weakest characteristic for the clinical translation of synthetic bioprinted scaffolds [3]. Still, synthetic materials that are commonly used in bioprinting applications include poly(ethylene glycol) (PEG) and PEGDA, PCL, poly(D,L-lactic-co-glycolic acid) (PLGA), poly(L-lactic acid) (PLA), and PPF, among others. PEG has long been used as a coating on medical devices to control host-immune responses or alter degradation rates in vivo [94]. Furthermore, PEG is commercially available in many physical (linear, branched, molecular weight variation) and chemical variants (diacrylated variant PEGDA) and is approved by the US Food and Drug Administration as a biocompatible material, which makes it a versatile polymer for bioengineering [81]. Here, biocompatibility, a broad term, means the material does not kill cells or induce an aggressive immune reaction, but it does not necessarily mean that it induces cell adhesion or proliferation. As for other synthetic materials, PEG lacks attachment sites that cells need to adhere to a substrate. PEG requires chemical immobilization of binding motifs to support cell adherence and stem cell differentiation [3,94]. These characteristics usually result in PEG being used as a secondary plasticizer component in bioinks: even more so, PEG is often modified with acrylate groups to create photopolymerizable PEGDA, a variation that is commonly used with extrusion or stereolithography approaches, and can be easily coupled with natural biomolecules for cell-laden bioinks [3,81].

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46. BIOINKS FOR 3D PRINTING IN REGENERATIVE MEDICINE

FIGURE 46.3 Synthetic materials in bioprinting applications: (A) Morphology of bioprinted starch- poly(ε-caprolactone) (PCL) scaffolds, including tubular scaffolds made by rapid prototyping and fiber-bonded scaffolds [77]. (B) Variations of three-dimensional (3D) printing (3DP) poly(propylene fumarate) (PPF) assessing scaffold and pore geometry to study the effects of pore geometry on cell viability and differentiation [10]. (C) Design and fabrication process of a 3DP PPF graft to treat a coarctation of the aorta [88]. CAD, computer-aided design; MRI, magnetic resonance imaging; mCT, microcomputed tomography.

PCL is synthesized by ring-opening polymerization of ε-caprolactone [77]. PCL is a highemolecular weight semicrystalline polymer that has good solubility, a low melting point, thermoplastic behavior, and an extended hydrolysis-induced degradation profile in vivo. This polymer is stable in the body for over 6 months [97]; then it exhibits nonenzymatic hydrolysis degradation of 2e4 years (depending on the molecular weight) at physiological pH and temperature, leaving no cytotoxic by-products [76,77,96]. Extended degradation is ideal for providing long-term load-bearing support during healing and regeneration processes [7]. PCL has a melting point close to 60 C, a relatively low temperature in the manufacturing industry, which allows easy processing [77,94,96]. Upon heating, PCL has viscoelastic properties ideal for the extrusion printing of constructs with elastic mechanical behavior, a characteristic that improves the brittle properties of polymers such as PLA and polyurethanes [97]. The Hutmacher group printed cylindrical scaffolds of medical-grade PCLetricalcium phosphate by fused deposition modeling that required added growth factors to increase the osteogenic potential of seeded mesenchymal stem cells (MSCs) [76]. The growth factorecoated PCL scaffolds successfully completed up to 12 months of unrestricted load bearing in vivo within large tibial defects in sheep [76]. The use of additional osteogenic growth factors indicates the biochemical limitations of PCL; other than the hydrophobic nonspecific binding of cells, PCL lacks binding motifs that provide specific binding sites for cells [94]. Combination with natural materials or other functionalized materials is the usual approach to address this limitation. The Atala group, for example, concurrently prints PCL with hydrogels based on gelatin, HA, or fibrin; the hydrogels provide the biochemical cues for cellular adhesion and viability [94]. Starch, a natural polysaccharide, has also been widely used to improve the biocompatibility of PCL [77]. PCLestarch prints can enhance and stimulate osteoblast proliferation for bone regeneration, support hippocampal neurons and glial cells to treat spinal cord injury, or support bovine articular chondrocyte adhesion and proliferation, and glycosaminoglycans for cartilage tissue engineering [77].

BIOINKS

813

PLA is a well-known established aliphatic polymer used in temperature-based extrusion methods [96,97]. It has a melting temperature close to 175 C, so it can be extruded in melt-based systems between 200 C and 230 C [97]. However, PLA glass transition occurs around 60 C and easily interacts with many plasticizers and solvents to change viscosity, a characteristic that allows printing at lower temperatures [96]. The resulting mechanical properties are usually high, with an elastic modulus around 1.5e2.7 GPa, but it tends to be brittle [96,97]. PLA is commonly used in orthopedic implants and drug delivery systems owing to its biocompatibility and biodegradability [96,97]. Nevertheless, its degradation is via the hydrolysis of ester bonds, which releases acidic by-products; in vivo, this may cause the localized decrease of pH through the release of lactic acid, inflammation, and cell death [97]. PLGA is the copolymer of lactide and glycolide, obtained via ring-opening polymerization, synthesized to address individual limitations and uncontrolled degradation of PLA and poly(glycolic acid) (PGA) [96]. Popular polymerization of D- and L-configurations of lactide yield the PLGA variation, which is frequently used owing to its improved toughness and easy manipulation of hydrolysis-driven degradation rates [96]. The Shah group synthesized “hyperelastic bone,” a particleeladen 3D bioink that combines hydroxyapatite, a highly bioactive ceramic, and either PCL or PLGA [78]. The extrusion-printed structures exhibited mechanical and physical properties that allow further manipulation (sheets that can be rolled, folded, or cut). The hybrid with PCL showed highly elastic properties capable of reaching a 61.2%  6.4% strain and a tensile elastic modulus of 10.3  1.3 MPa, a behavior superior to that of the PLGA combination (36.1%  4.3% strain and 4.3  0.4 MPa elastic modulus). In terms of cell interaction, the PLGA combination showed better results; both the hydroxyapatiteePLGA and hydroxyapatiteePCL scaffolds supported human MSC adhesion and proliferation and induced osteogenic differentiation in the absence of engineered growth factors after 28 days [78]. PPF is a biodegradable polymer broadly applied in tissue engineering owing to its ability to form cross-linked networks through its carbonecarbon double bond [10,22]. The Mikos [22] and Fisher [10,88] groups extensively studied PPF and its cross-linking capabilities for 3DP and tissue engineering. Because it is biocompatible and can be photocross-linked, PPF is a prime candidate for 3DP via stereolithography [22,88], but it can also be as a viscous bioink for extrusion and cured using a UV source [22]. In the first case, the printing process is driven by the intensity of the light source and the proportions of photoinitiator and photoinhibitor in the bioink, but the resulting mechanical properties of the constructs heavily depend on the amount of printing or postprinting exposure to UV, which determines the polymer cross-linking density [10,88]. In the case of extrusion, PPF resins exhibit shear thinning behavior and the concentration of PPF drives the viscosity level. Other factors such as fiber spacing during deposition and pressure affect the pore size and fiber diameter, respectively, but interplay among the factors can also alter scaffold architecture [22]. Melchiorri et al. reported that human umbilical vein endothelial cells and human umbilical vein smooth muscle cells were seeded on stereolithography-printed PPF surfaces and proliferated in a 7-day study [88]. Similarly, MSCs were cultured on PPF scaffolds over 7 days and exhibited levels of metabolic activity that were not statistically different from cells cultured on standard tissue culture polystyrene [10]. In vivo, using 3DP PPF grafts to treat a coarctation of the aorta in mice for 6 months, printed PPF experienced a 40.76%  8.37% decrease in mass, and full endothelialization of the inner lumen on the grafts was observed even without preceding cellseeding or surface modifications [88]. Natural Materials Based on the definition earlier proposed, bioinks will interact with biological components (e.g., tissues, cells, proteins, growth factors) during or after the actual print or will serve as structural components during the printing of scaffolds that will interface with biological components. Without considering the specifics of in vivo or in vitro applications, there is an imminent interaction between cells and tissues with the bioink or its by-products. Natural materials are taken from animal or plant sources; these are materials that naturally developed to sustain cellular life cycles, nutrient and waste transport, and healing processes. They are composed of the perfect combinations of amino acid sequences, protein ratios, growth factors, and cytokines, thus intrinsically providing safe and nurturing interactions with cells. The composition provides the proper biochemical environment for cells to adhere or feel attracted to, subsequently allowing individual cells the healthy completion of the cell cycle and then induce cellular proliferation, migration, and differentiation. This is the basic definition of cytocompatibility, and natural materials provide a high intrinsic level of it [3,91,94]. Just as important, the composition and biochemistry of a natural material are designed to be degraded by physiologically viable processes through natural enzymatic and chemical processes, and to be discarded by natural metabolic activity, leaving behind no significantly harmful by-products [91]. Bioinks

814

46. BIOINKS FOR 3D PRINTING IN REGENERATIVE MEDICINE

from these materials can be further biochemically enhanced by encapsulating tissue-specific growth factors, genes, and other controlled-release chemical-regulation factors. Similarly, the surfaces of printed hydrogels can be functionalized by adding the same biochemical factors with both approaches, aiming to recreate environments more like those of in vivo tissues [78,91,94,95]. The balance between cytocompatibility and degradation means that these materials naturally go through the proper cycles and rates needed to induce healthy integration with host tissue [3]. Natural materials can be used as isolated, purified proteins (e.g., collagen, fibrin, keratin, or elastin) or as the natural protein combinations already present in the ECM, combinations that are specific to each type of tissue and determine the type of cells present, the bulk mechanical properties, and its function. Methods to obtain and alter natural materials in laboratory mainly consist of enzymatic cleaving, ionic interactions, and variations in temperature and pH [3]. These methods are used to cross-link the bioinks via multiple 3DP techniques (Table 46.1) and produce hydrogels with theoretically fine-tuned biochemical and mechanical properties. As it will be discussed further (see Cell-Laden Bioinks section), encapsulated cells or those that later migrate into the printed scaffolds are greatly affected by the mechanical cues imparted by the surrounding material. Cellular adhesion, morphology, migration, and especially differentiation, have been widely proven to be affected by the stiffness of the substrate [70,94]. In general, 3DP natural materials result in weak hydrogels difficult to manipulate into specific ranges of physical properties; they are limited by the inefficient or low-energy cross-links achieved by the traditional methods mentioned [94]. Weak mechanical properties have been used as an advantage when attempting to model or regenerate soft tissues or substrate for cell culture, but they are a severe disadvantage when the applications relate to load-bearing hard (e.g., bone, cartilage) or elastic tissue (e.g., muscle, skin, vascular and gastrointestinal tissues, ligaments, tendons). This major weakness has led to combinations with strong and elastic synthetic materials by co-printing or as hybrids, as will be detailed later (see Co-printing and Hybrid Bioinks section). The intrinsic biocompatibility of natural materials is the main reason why these have been used to formulate bioinks for use in a wide range of in vitro and in vivo bioengineering and regenerative medicine applications (Fig. 46.4). The most popular materials are generally proteins from mammalian origin such as collagen, gelatin and gelMA, fibrin, HA, elastin, and keratin; similarly, popular polysaccharides from plant sources include alginate, starch, agarose, and silk, among others.

(a)

(b)

(c)

(d)

(A) (B)

FIGURE 46.4 Natural materials in bioprinting applications: (A) Examples of plotted gelatin methacrylateegellan structures after UV curing: (clockwise) solid pyramid, solid hemisphere, porous hollow cylinder, and 30% porous hemisphere (5-mm scale bars) [70]. (B) Examples of decellularized extracellular matrix (dECM) printing in combinations with polycaprolactone (PCL) framework, including heart dECM (hdECM), cartilage dECM (cdECM), and adipose dECM (adECM) (scale bar ¼ 5 mm). Also, scanning electron microscopy of PCL tissue constructs with adipose-derived dECM bioink (scale bar ¼ 400 mm) [100].

BIOINKS

815

Collagen, particularly type I collagen, is the most abundant ECM protein in tissues [9,94,98]. The most common types of collagen are the fibril-forming collagens (e.g., type I, II, III, V, and XI); they are the main component in the ECM of tissues such as bone, tendon, ligament, skin, muscle, or cornea [9,98]. Because of its natural abundance in all types of tissues, collagen has variations that interact with most types of cells and it performs a wide range of mechanical roles in soft, elastic, or hard tissues. For this reason, collagen is arguably the material that most researchers have tried to adapt for bioengineering applications and has been reported as the most commonly used for cell and tissue culture [94]. It has been used in 3DP techniques such as extrusion [23,55], stereolithography [9], and inkjet [51], mostly employing variations of pH-triggered or temperature-triggered gelation [98] that range from hours to minutes [91]. Cross-linking of collagen by altering pH using solubilized sodium bicarbonate solution has been used to construct multilayered cellehydrogel composites. This method provided a novel approach to printing both fibroblasts and keratinocytes in a single experiment to model dermaleepidermal-like distinctive layers in a 3D hydrogel [15]. Natural material hydrogels have typically been reported to have subpar printing resolution compared with synthetics. Nevertheless, Bell et al. reported printing line widths of about 1 mm using multiphoton cross-linking of type I collagen with a flavin mononucleotide photosensitizer, which confers structural control at a microscale level [9]. Collagen has also been widely used in combinations with other natural and synthetic materials, principally bringing strong biocompatibility to the mixture [13,91,99]. Gelatin is the denatured form of collagen that has undergone partial hydrolysis [9,94]. As collagen, gelatin is characterized by its wide availability, biocompatibility, predictable enzymatic degradation, nontoxic by-products, and inherent cell binding motifs [70]. It has been involved in engineering soft and hard tissues ranging from liver to bone, either by itself or as part as hybrids such as gelatinealginate, gelatinefibrin, and gelatineHA [95,100]. Gelatin is widely regarded as the easiest protein to print, mainly because of thermally responsive behavior that allows extrusion at temperatures below 20 C and hydrophobic cross-linking [64,94]. However, the melting temperature of gelatin (30e35 C) is below physiological temperature, which severely limits its clinical application in vivo [95]. Even with the high resolution obtained by extrusion-based printing [70], gelatin hydrogels are usually soft and limited by temperature, which requires further cross-linking either by postprinting approaches (e.g., using glutaraldehyde or thrombin [95]) or by adding functional groups [94]. Like the acrylate modification on PEG that produces the versatile PEGDA, methacrylamide photoinitiator groups can be used on gelatin to obtain gelMA to produce a photocrosslinkable resin. This modification enables irreversible cross-linking, generally by UV irradiation, that preserves printed architectures under physiological conditions [70]. UV exposure time and gelMA concentration regulate printability, whereas the degree of methacrylation determines the mechanical properties and additional acetylation can be used to influence the rheological properties of the bioink further [70,96]. Fibrin is a glycoprotein composed of fibrinogen monomers; it is synthesized in the liver by hepatocytes. In the body, it has important roles in blood clotting and wound healing [94]. The clotting pathway has been replicated as a cross-linking method for 3DP; thrombin is used to polymerize fibrinogen rapidly into cross-linked fibrin [12,94]. As a glue-like gel, fibrin has been used clinically as surgical hemostatic agents and sealants [94]. Enzymatically quick cross-linking rates have been exploited with extrusion and inkjet-based printing [51,52,94], but the mechanical properties of the constructs have been paradoxically described as both robust [94] and weak regardless of the concentration of the reagents [46]. Fibrin-based hybrids materials with natural or synthetic components are usually reported to fine-tune mechanical properties depending on the application, including cross-linking with PEG and adding PGA fibers, PLGA, hydroxyapatite, or demineralized bone matrix [46]. Alginate is a natural polysaccharide derived from algae or seaweed. Sodium alginate is generally cross-linked in calcium chloride (CaCl2) aqueous solution, via an ion exchange reaction between sodium and calcium [3,101]. This chemically efficient reaction results in biocompatible, lowepolymer density, highewater content hydrogels [3]. Traditionally, cell encapsulation in calcium alginate hydrogels was the main application of alginate in tissue engineering and bioengineering models, despite the controversial effects of CaCl2, the cross-linking reagent, as well as sodium citrate and ethylenediaminetetraacetic acid (EDTA), commonly used chelators, on cell viability [94]. This ionic cross-linking approach has been implemented in bioprinting. It works particularly well in extrusionbased systems that extrude alginate resin into CaCl2 reservoirs [6,7,14,18,60]. Cells can be suspended in a solution of sodium alginate in cell-specific culture medium, after which cross-linking is induced by incubation in CaCl2 and results in a hydrogel construct laden with cells. This approach has been successful in bioprinting, such as in human cardiac-derived cardiomyocyte progenitor cells (hCMPCs) for an in vitro committed cardiac tissue [18]; heterogeneous scaffolds with MSCs and chondrocytes (in alginate with osteogenic or chondrogenic differentiation medium, respectively) for osteochondral tissue engineering [6]; or encapsulated HepG2 liver cells printed directly on a polydimethylsiloxane chamber for a microfluidic pharmacokinetic liver model [101]. Nevertheless, Carrow et al. stated that there were major challenges for bioprinting alginate: (1) the difficulty of controlling the ionically driven process,

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which results in unpredictable microstructures; and (2) the high solubility of alginate, which was a disadvantage when printing thick structures by extruding directly into CaCl2 aqueous solutions [96]. Hyaluronic acid (HA), also called hyaluronan, is a hydrophilic nonsulfated glycosaminoglycan present in the ECM of tissues [94]. The Atala group has regularly used HA in bioprinting processes by adding photocross-linkable methacrylate groups that can undergo free radical polymerization when irradiated with UV light. This modification allows soft hydrogels to be printed via stereolithography or extruded with additional postprinting UV curing [62,63,94]. Although not mechanically robust on their own, HA hydrogels have served in cutaneous and corneal wound healing, prototype vessel structure bioprinting, tumor modeling, and 3DP of cell-laden structures [94]. Other lesser-used natural materials for bioprinting include proteins and polysaccharides such as elastin [98], keratin [45], starch [77], and agarose [3,91]. Despite the success of using isolated natural polymers, there has been growing interest in using the innate combination of proteins in the ECM. The ECM allows structural support and anchoring to cells and also provides a substrate for transport and communication, ultimately affecting the survival and differentiation of cells. CelleECM interactions are extremely complex and cannot be fully and precisely replicated in vitro or engineered from isolated proteins. Several groups have presented decellularized ECM (dECM) bioinks derived from adipose, cartilage, heart, bone, or skin tissues. The combination of proteins in the ECM can be understood as a hybrid of multiple natural materials; therefore, available cross-linking methods and bioprinting approaches have been successful in producing 3D dECM scaffolds. Pati et al. developed dECM bioinks that can be extruded as filaments; printed scaffolds can then undergo gelation at physiological temperatures, remaining in the solution below 15 C and cross-linking by incubation at 37 C [100]. Co-printing and Hybrid Bioinks Approaches attempting to print synthetic and natural materials individually have produced scaffolds with mechanical and biochemical properties that affect cells and tissues differently, and thus can be used in different types of in vitro and in vivo applications. A common generalization in the field is that synthetics are used for their strong, finely tunable mechanical properties; nevertheless, it has been proven that they can provide tunable degradation rates, functionalization capabilities, and various degrees of biocompatibility and print resolution. On the other hand, natural polymers have proven to be highly compatible with a wide variety of cells and biological components, mostly owing to their inherent composition and function. As with synthetics, it is hard to generalize the negative characteristics of natural bioinks, but the properties (mechanical or biochemical) are rarely fine-tunable and usually are presented as ranges and wide error margins commonly associated with batch-to-batch variability. Overall, natural material properties, printing quality, and in vitro or in vivo behavior can be described as unpredictable and difficult to replicate. Combining both types of materials has been an increasingly popular hypothesis that relies on the positive properties of each. In theory, synthetic materials provide structural integrity and printing definition, whereas natural polymers can be used to incorporate cells and other biological components [95,97,100]. Two broad categories for combining synthetics and natural materials as bioinks for 3DP applications are (1) co-printing, the individual but parallel printing of natural and synthetic resins; and (2) hybrid bioinks, in which the resin is a uniform solution of both materials printed as a single construct. Co-printing approaches rely on printing synthetic scaffold structures with robust mechanical properties onto which natural hydrogels can be printed. This addresses the common limitation of natural materials, the inability to maintain uniform 3D structures in vivo (e.g., to allow tissue load bearing or provide a specific porosity or microstructural pattern) or in vitro (e.g., to be handled robustly in bioreactors, or as cell substrates), by integrating a synthetic scaffolding [14]. The main challenge is that co-printing relies on technologies that dispense more than one material during the printing process, sometimes with radically different deposition necessities and cross-linking mechanisms, as illustrated in Fig. 46.5. In extrusion-based systems, for example, the rheology of the materials is the driving principle, and variables such as viscosity, flow rate, temperature, and pressure determine the extruded line width, fabrication time, or print resolution [7]. Shim et al. used a multihead tissueeorgan building system possessing six dispensing heads to dispense thermoplastic PCL and alginate hydrogel individually in the same structure, to produce constructs containing two different cell types for osteochondral tissue regeneration [7]. As they reported, the viscosity of alginate solution was about 10 Pa s and needed low driving forces but high force control to achieve high resolution; on the other hand, viscosity for PCL ranged from 1020 to 2560 Pa s at 80e120 C (a temperature high enough to damage cells) which required high driving forces to extrude [7]. The same PCLealginate approach was reported using a multihead deposition system, they printed PCL and chondrocyte-laden alginate with and without transforming growth factor b [14]. Here, PCL was extruded at 80 mm/min at 80 C using a 650-kPa pneumatic pressure. Sodium alginate was deposited at room temperature between lines of PCL at 400 mm/min, and then cross-linked in sodium chloride

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FIGURE 46.5 Co-printed systems with synthetic and natural bioinks: (A) Synthetic and natural materials are deposited using independent cartridges and print heads. The different properties of each material can be exploited to print them separately and sequentially, but it is necessary to consider force and temperature shocks that may alter one material when the other is printed in contact with it. (B) Bioprinting of poly(εcaprolactone) (PCL)-alginate scaffolds. Scanning electron microscopy images comparing PCL and PCLealginate scaffolds [14]. (C) AlexaFluor 594 red fluorescent dye positively staining type II collagen fibrils deposited by viable chondrocytes encapsulated in the same PCLealginate scaffolds. Also, live/dead assay of chondrocytes showing cellular viability on the natural material but none on the synthetic component [14].

[14]. These cases illustrate the complexity of printing two different materials in the same structure; as stated before, the key lies in the independent extruding heads, in which the variables of the process (temperature, speed, pneumatic pressure, architectural patterns, etc.) can be controlled independently for each different material. The ability to control each material separately but still building a single construct has high impact in resolution. In particular, the ability to place cells and materials with different properties into specific patterns confers high control over the resulting mechanical and biochemical behavior of the whole construct. Complex tissue constructs have been achieved via co-printing, such as the muscleetendon unit (MTU) approach reported by the Atala group [60]. The MTU is the interface between muscle, which is elastic and fibrous in nature, and the tendon, which is stiff and sparsely cellular. In this approach, natural hydrogels were composed of gelatin, HA and fibrinogen. Elastic polyurethane and hydrogel laden with C2C12 myoblasts were chosen for the muscle side, whereas stiffer PCL and National Institutes of Health (NIH)/3T3 fibroblasts gels were selected for the tendon group [60]. The co-printing approach allowed the controlled construction of the interface in which two mechanically different tissues with different cell populations flawlessly met (Fig. 46.5). They reported that the construct was not only able to mimic the complex mechanical behavior of the MTU but successfully retain cell viability in both hydrogel portions (C2C12 cells with 92.7%  2.5% viability and NIH/3T3 cells with 89.1%  3.3% after 7 days [60]). The PCL-hydrogel co-printing approach has also been used to mimic mandible bone, ear cartilage, and skeletal muscle for tissue engineering [92]. Not restricted to producing tissue scaffolds for eventual in vivo applications, co-printing enables the construction of complex models for in vitro testing, particularly vascularization, microfluidic, and tissue-on-a-chip models. Having multiple heads depositing various materials and cells, under strict spatiotemporal control, has allowed researchers to produce highly complex models that more closely resemble the behavior of biological systems in vitro. These efforts usually require the use of cell-laden hydrogels (see Cell-Laden Bioinks section) and complex types of bioinks such as sacrificial (see Sacrificial Bioinks section) or supportive bioinks (see Supporting Bioinks and Supporting Baths section). An interesting example of such structures can be observed in the work of the Lewis group. Those researchers used the co-printing of natural and synthetic materials to develop in vitro models of tissues

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and vascularization [12]. They concomitantly printed natural materials that included cell-laden castable and printable ECM composed of fibrinogen and gelatin, cross-linked via thrombin and transglutaminase enzymatic reactions, whereas the synthetic parts included silicone chip bases and Pluronic sacrificial materials [12]. These materials were deposited and cross-linked independently and sequentially to produce a highly organized vascularized tissue analogue based on the strong characteristics of both natural and synthetic materials. Hybrid bioinks are the second approach to integrating synthetic and natural materials. In this case, there is a single bioink solution that includes both types of materials as solutes. Generally, hybrid bioinks are composed of a synthetic substrate solution with specific mechanical and rheological properties into which a natural component is mixed to alter biochemical and biocompatibility properties, as illustrated in Fig. 46.6. For synthetic materials, adding natural groups to the bioink usually results in improved compatibility with cellular processes, including binding sites and growth factors or reducing the high hydrophobicity of synthetics [14]. For natural materials, the benefits are usually observed as structural or mechanical, but the inclusion of synthetic polymers to the protein chains also enables natural materials to be processed using the techniques and equipment designed for synthetics. The weak ionic interactions or unpredictable enzymatic processes reserved to process alginate, fibrin, or collagen can be changed for optimized and finely tunable techniques such as photocross-linking or high-resolution extrusion [94]. The hybridization of the materials can be achieved by mechanical entanglement of the materials in solution or by chemically joining the polymer and protein chains. The first is a common approach to improving the mechanical or rheological properties of natural materials. Narayanan et al. used human adipose tissue stem cells loaded in alginate bioink with suspended PLA nanofibers [102]. The cell-laden alginate solution could be prepared separately from the nanofibers but they were vortexed together into a single solution, printed, and cross-linked, trapping the PLA within the hydrogel with no cross-linking interaction between the two. This approach was successful in producing constructs that allowed stem cell differentiation down the chondrogenic pathway; more interestingly, it revealed a

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FIGURE 46.6 Hybrid bioinks combining synthetic and natural biomaterials: (A) Hybrid bioinks can be three-dimensionally (3D) printed from a single ink containing both the synthetic and natural components. The two materials can be printed into a single construct and structurally held together either by mechanical entanglement of the polymers or by biochemical cross-linking (a process that can occur during resin formulation or after printing). (B) 3D bioprinting of poly(ethylene glycol) with reactive ends (PEGX)ebiomolecule hybrids. Extrusion printing of PEGXegelatin hybrid. Next, PEGXegelatin (red) and PEGXefibrinogen (blue) can be printed in a single cylinder. Finally, PEGXepoly(ethylene glycol) (PEG) and PEGXegelatin fibers crosshatched, live/dead assay show viable human dermal fibroblasts preferentially adhering to PEGXegelatin, a syntheticenatural hybrid, but not to the PEGXePEG, a synthetic mixture [81]. (C) Optical and live/dead fluorescence images (after 16 days) showing cell viability in bioprinted strands of alginate with adipose-derived stem cells (Alg-hASC) and of alginate-poly(L-lactic acid) (PLA) nanofibers with human adipose-derived stem cells (Alg-Nf-hASC) [102].

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method to use the distribution and alignment of the PLA nanofibers to stimulate orientations within the ECMmimicking hydrogels [102]. The second approach relies on chemically altering and cross-linking the synthetic and natural chains. It is commonly used to improve the biocompatibility of the synthetic portion or print the natural material using synthetic methodologies. The chemical modification allows the personalization and optimization of the resulting bioink chain, which means higher specificity of the printed materials to cell or tissue functions. The Shah group, for example, used functionalized PEG to include a variety of proteins in extrudable, tunable, and cell-compatible bioinks [81]. As illustrated in Fig. 46.6, PEG with reactive ends (PEGX) is used to bridge protein and polymer chains in a variety of configurations, producing mixtures such as PEGXecollagen, PEGXegelatin, PEGXefibrinogen, and PEGXePEG, among many others, that can be successfully loaded with cells and printed via extrusion [81]. Cell-Laden Bioinks Current definitions of bioinks refer to resins that are loaded with cells and printed. As described before, we expanded the definition of bioink to include several categories of printable materials, and do not necessarily consider cells to be the determinant “bio” factor. Nevertheless, the importance of cells for bioengineering and regenerative medicine is undisputable, and cell-laden bioinks are crucial for the development of 3D bioprinting technologies and the goal of printing functional in vivo and in vitro tissues and organs (Fig. 46.7). Synthetic and natural materials have been proven to have various degrees of success in cell compatibility, tissue integration, and tunable mechanical and biochemical properties, so why incorporate the complex additional factor of cells? It is commonly accepted that the acellular scaffold approaches have poor translation in vivo, mostly owing to the limitation of cells adhering only to the surface of the constructs. The success of this approach is unpredictable, locations and concentrations of growth factors or chemoattractants within the constructs cannot be guaranteed, and cell behavior cannot be controlled [14]. We have mentioned before that the key term that defines modern bioprinting is control. Being able to control where cells, matrix, growth factors, and other biological components are placed results in structures with higher orders of specificity and functionality. If materials and cells can be located and properly stimulated to construct gradients, strata, or clusters, there is a higher chance for success without relying on the unpredictable colonization of native cells. Fedorovich et al. exploited this control feature to reproduce the specific spatiotemporal distribution of cells and ECM in osteochondral tissue [6]. The bioinks consisted of alginate solution in osteogenic or chondrogenic differentiation medium, in which MSC or chondrocytes, respectively, were added. After successful extrusion mimicking the adjacent bone and cartilage portions, ionic cross-linking, and subcutaneous implantation in mice, the dual, heterogeneous scaffolds showed two different cell lineage commitments, with each type of cell remaining in its printed position and depositing lineage-committed ECM [6]. Another multiphase approach to osteochondral tissue engineering was presented by the Demirci group, aiming to study tissue interfaces in the anisotropic composition of fibrocartilage [8]. Human MSCs were encapsulated in gels and printed by droplet deposition in an arrangement with zone-specific biochemical factors and ECM components (transforming growth factor-b1 for fibrocartilage and bone morphogenetic protein-2 for bone regions). Again, cells showed different lineage commitment by upregulating osteogenesis- and chondrogenesis-related genes defined by the position and matrix in which they were printed, yet constructing a single heterogeneous scaffold [8]. Cell-laden bioinks are generally hybrid or natural bioinks that can be 3DP into hydrogels. The materials provide innate cell-binding motifs, hydrophilic surfaces, and low cytotoxicity to promote cell adhesion [94]. The hydrogel structures provide soft, degradable, and swelling networks that mimic ECM and allow cell migration, metabolism, and differentiation with minimal restriction [70,97]. The mechanical properties of 3DP hydrogels can be modified by regulating cross-linking density, the linking chemistry, or polymer concentrations to match properties close to those of native ECM [94]. Structural properties of the microenvironment, such as stiffness or composition, can deliver biochemical cues by mechanotransduction to regulate cell shape, migration, and differentiation lineage selection [81,91,94]. As an example, bioinks used for high-resolution prints generally produce stiffer gels ideal from a 3DP and structural standpoint, but the high elastic modulus will drive stem cell differentiation toward the stronger tissue lineages (bone and cartilage), making it an obsolete approach to produce soft tissues [70]. Hydrogels seem to have the ideal characteristics for cell adhesion and sustenance, but the materialecell tandem must also work with 3DP methods and account for the impact of the processes on cell function after printing. First and foremost, no part of the bioink, printer setup, additional cross-linking mechanisms, or by-products can be cytotoxic; and they have to be sterile-compatible. This seems straightforward, but it considerably reduces the available materials and processes that can be used [3]. 3D bioprinting methods mostly rely on physical forces or temperature to deliver materials. The most popular methods, extrusion-based and inkjet printing, rely on some mechanism of pressure that pushes the bioink through a nozzle. This setup translates pressure on the cells first as a compressive

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Applications of cell-laden bioinks: (A) Micropatterning and bioprinting the anisotropic three-dimensional (3D) fibrocartilage phase, composed by merging two human mesenchymal stem cell (hMSC)-laden gelatin hydrogels with either tendon or bone growth factors [8]. (B) Macroscopic and live/dead images of printed human cardiac-derived cardiomyocyte progenitor cells in 5% alginate scaffold [18]. (C) Fluorescence image of alginate hydrogels laden with chondrocyte (green) and osteoblast (red) cells and dispensed into a poly(ε-caprolactone) (PCL) framework. Every other pore is empty for oxygen and nutrient transportation [7]. BMP-2, bone morphogenetic protein-2; TGF-b, transforming growth factor-b.

FIGURE 46.7

force within the cartridge and then as shear stress while they are moving through the nozzle. Varying the pneumatic pressure, extrusion speed, and nozzle diameter regulates the stress delivered to cells and has been proven to affect cell viability during and after printing [14,94]. A variety of printing protocols have studied these parameters and successfully printed multiple types of cells and materials with a very high viability rate: extrusion of MSCs and chondrocytes in alginate with 89% viability 5 h after print [6] or 97% cell viability after thermal inkjet deposition [34]. High shear rates and shear stress have been proven to harm the cells, but high viability may be explained: (1) Stress causes protein denaturation by damaging the tertiary or quaternary structures of the chains, a process that is reversible with time [34]; or (2) natural or hybrid gels have shear thinning behavior, which decreases stress on the cells even at high shear rates or pressures [6,100]. Bioink viscosity and surface tension can also be modified to reduce shear stress using solvents or surfactants [34,95], although these must comply with minimum cytotoxicity requirements. In the case of stereolithography, the UV intensities required to initiate and sustain photoinitiator and photocross-linking reactions can negatively alter cell morphology and viability [3]. The amount of energy radiated can also cause irreversible damage to the cells as a result of increasing temperatures and dehydration. Temperatures different from physiological 37 C, both high or low, and drastic changes will alter cell metabolic activity and may cause cell death [94]. This is a critical consideration for co-printing approaches, in which the

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temperature required to deposit synthetic materials could easily be above 60e100 C and then solidify on cold surfaces. Both extremes having the potential for irreversible cell damage [4,6,12,14]. Chemically or pH-induced crosslinking by ionic and physical mechanisms can also harm cells. Additional cross-linking or postcuring approaches, such as postprint glutaraldehyde or EDTA chemical cross-linking or additional UV irradiation, which guarantee 3D geometries and mechanical properties, are commonly seen as additional negative factors [4]. A critical element to maintain cells in the bioinks is the permanent need for proper oxygenation and metabolic transport. In vivo, any cell will be located within 100e200 mm of a capillary, the maximum distance for adequate gas and metabolic exchange [4,7,91]. Thick-casted or printed hydrogels without pores do not allow diffusion deep enough to supply oxygen or nutrients and result in necrotic cells encapsulated in the center of the structure. Generally, the design of 3DP architectures is envisioned with pores or channels that provide open transport pathways and open vascularization channels. Interconnected pores, constructed by weaving strands or layered patterns, with dimensions ranging from tens to hundreds of micrometers usually allow transport through the scaffolds and report high cell viability and development [7,10,64,95]. Another common approach is to induce fast vascularization of the constructs, generally by adding endothelial progenitor cells or growth factors and using bioreactors [4,12]. Overall, accounting for the strengths and limitations of printing cell-laden bioinks, numerous approaches regularly report positive effects of this approach on cell behavior for a wide range of applications. As an in vitro example: Gaetani et al. printed a model of undifferentiated but committed cardiac tissue [18]. Here, sodium alginate was dissolved in a culture medium and mixed with hCMPCs. 3D constructs were obtained by printing strands into layers, stacking them to obtain different degrees of porosity, and cross-linking with CaCl2. Compared with regular 2D cultures, bioprinting had no effect on cell viability and proliferation, but it increased cardiac lineage commitment by upregulating early and late cardiac transcription factors and markers [18]. On the other hand, in an in vivo application aiming to produce an implantable bioartificial liver, Wei et al. used a gelatinefibrinogen matrix loaded with rat hepatic cells to produce 3D porous constructs via extrusion printing and thrombin-induced gelation [51]. After extrusion, about 98% of the hepatic cells were reported to be viable, steadily producing albumin, and dissolving the surrounding gelatin matrix throughout the culture time [51].

Sacrificial Bioinks Sacrificial bioinks enable the fabrication of complicated structures and open geometries without dealing with many of the difficulties related to satisfying biological requirements. Using a sacrificial bioink for a print material, the print volume is initially created and will subsequently be washed away, as shown in multiple approaches in literature illustrated in Fig. 46.8. The bioink provides space-filling volume and support that will be evacuated. Some groups refer to their sacrificial bioinks as fugitive inks to suggest its temporary role in printing in the scope of the final structure [103]. Therefore, a sacrificial bioink only needs to be nontoxic and will not introduce harmful by-products; however, no further biological features are necessary, such as cell adhesiveness or biodegradability [1]. Here, by nontoxic, we idealize success cases in which the by-products are also noncytotoxic. Still, sacrificial bioinks ideally match these specifications: high print fidelity, ease of removal, and the lack of toxicity (Fig. 46.8). To enable ease of removal, an important material property is its gelation process: in other words, the conditions under which the printed material will wash away. Some examples of bioinks, such as Pluronic F127 or gelatin, have a thermally reversible gelation process. Therefore, although printing can occur at one temperature, the printed material can evacuate when another temperature is attained. The Lewis laboratory is one of the leaders using Pluronic F127 as a fugitive ink to create perfusable networks in tissue mimics. Pluronic F127, a poloxamer, is solid at 37 C but it can be liquefied when cooled to 4 C. They take advantage of this material property to flush out the fugitive bioink with cold cell media, leaving behind perfusable channels. The resulting print structures resemble thick vascularized tissue mimics with endothelialized lumens viable after 45 days of perfusion [12]. The same group applies this strategy to other tubular tissues; the technique was also applied to fabricating renal proximal tubules [86]. Other common sacrificial bioinks include agarose [67], alginate [57], and gelatin [66]. The Khademosseini laboratory applied sacrificial templating strategies with different materials. For agarose, the geometry was extruded to form a solid network at 4 C. Subsequently, the agarose fibers could be manually removed or lightly vacuumed because the material does not adhere to surrounding photocross-linked hydrogel [67]. With a different material, the Khademosseini cohort used sodium alginate to fabricate a sacrificial network. The gelation of alginate occurred ionically with calcium chloride and could be removed with EDTA treatment [57]. Others applied this strategy to 3DP other tissue mimics such as aortic valves and bone; however, those groups did not print with this bioink sacrificially

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Sacrificial templating strategies: (A) The Lewis laboratory extrudes Pluronic F127 filaments in a renal proximal tubule geometry that is cast with a gelatinefibrinogen gel. The extruded Pluronic can then be flushed out by cooling to 4 C to leave behind perfusable renal tubules that can be seeded with proximal tubule endothelial cells [86]. (B) Miller and colleagues similarly extrude a lattice structure with carbohydrate sugar glass that can be encapsulated with a number of different extracellular matrix (ECM) or ECM-like materials and later dissolved for a perfusable vasculature. Human umbilical vein endothelial cells are seeded in the casted gel and remain viable owing to the perfusion of cell media through the vascular architecture [11]. (C) Lee et al. print collagen layers with intermittent sacrificial and cell-laden gelatin filaments via inkjet. Gelatin’s gelation process relies on thermal transition; therefore, it can be liquified to evacuate open geometries. When seeded with endothelial cells, junctional markers confirm their barrier function after long-term perfusion [66]. (D) Kinstlinger et al. demonstrate that selective laser sintering can also be implemented with a sacrificial templating strategy. Vascular geometries can be formed with poly(ε-caprolactone)(PCL) and casted over with polydimethylsiloxane (PDMS) (shown here) or other materials and leave behind open channels [29]. 3D, three-dimensional; DAPI, 40 ,6-diamidino-2-phenylindole; PTECs, proximal tubular epithelial cells; RFP, red fluorescent protein; VE-cadherin, vascular endothelial cadherin.

FIGURE 46.8

[58,104]. An additional naturally derived sacrificial bioink, gelatin, is thermally reversible, which proves useful for a sacrificial templating strategy. Lee and colleagues used inkjet printing to form vascularized tissue constructs dropwise and layer by layer with gelatin and collagen. Collagen layers were printed and polymerized at 4 C with a pHaltering cross-linking agent (sodium bicarbonate) so that the thermal responsive nature did not take precedence over the intended material evacuation. Within the collagen layers, gelatin was left uncross-linked and therefore was removed by liquefying the printed structure when the final print structure was raised to room temperature [66]. These natural sacrificial bioinks are all able to be easily evacuated, which fits one of the crucial characteristics of a sacrificial bioink.

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So far, the sacrificial bioinks that have been discussed have poor mechanical strength potentially leading to complications in maintaining print fidelity after the print. Some researchers have addressed this complication by using an alternative sugar material that has high mechanical stiffness and is water soluble [11,105]. Miller and colleagues developed an approach to using carbohydrate sugar glass to fabricate complex vascular designs. This strategy circumvented issues related to poor mechanical strength. However, a similar evacuation approach is employed in which the printed perfusable network is evacuated by running cell media through the channels to provide a fluidic network. Although the carbohydrate glass lattices were printed at high temperatures (110 C, a temperature physiologically unviable), the finished print volume could be brought to physiological temperature with complete cell media for 10 min to dissolve the carbohydrate glass [11]. After evacuation, these open channels could be perfused with cell suspensions for long-term cell culture. However, the extrusion-based printing technique used to fabricate structures has limitations. One example of a limitation is that because the build volume relies on depositing material layer by layer, this technique is unable to create overhanging structures without support. Some in the bioengineering community employ SLS as a different 3D bioprinting technique extending the list of bioinks while addressing the overhang limitation of extrusion-based approaches. SLS-based prints have been commonly used with materials such as PCL and PLA to produce bone-mimicking scaffolds, for example. However, emerging research for this printing strategy falls under the scope of sacrificial bioinks [106,107]. Limited in resolution only by the smallest powder size of the bioink and the laser used, SLS exhibits micron-sized fabricated structures with complicated overhangs with a sacrificial templating workflow [29]. Using PCL, a lattice structure can be selectively laser sintered and dissolved to leave behind a fluidic network. However, to provide perfusable channels, the printed construct requires dissolving in dichloromethane, which may or may not induce toxicity to future seeded cells [29]. More gentle evacuation for future cell encapsulation may be necessary to translate to a cytocompatible workflow. Although others use sacrificial templating strategies with plastics (i.e., acrylonitrile butadiene styrene) to fabricate intricate microfluidic systems and subsequently remove them with acetone, these removal approaches may not lend well to tissue incorporation [108]. Although there is a short list of sacrificial bioinks, these materials share one thing in common: the need to have multimaterial integration for a finalized structure. Inevitably, for the sacrificial ink to be evacuated, a requirement is an encapsulating material in which the flushed material can leave behind a space. The selection of the encapsulation material, the matrix, or matrix-mimicking bioink (see Extrusion-Based Printing section) is crucial to incorporating cells, or otherwise an interface with biological tissue might require cell-adhesive sites or degradation properties. Moreover, the gelation process of the encapsulation should ideally mismatch the sacrificial bioink so that the intended material is the one being sacrificed at the point of dissolving.

Supporting Bioinks and Supporting Baths Supporting bioinks and supporting baths are used during 3DP to improve the mechanical properties and expand the geometric capacities of the bioprinted scaffold, as shown in Fig. 46.9. High viscosity is commonly a desired material property for supporting baths to provide structural integrity as the print is fabricated. This material property functions to hold the print in place so that the printed material alone (without the support bath) would be unable to be maintained (i.e., overhanging features or ultrathin features). Consistent with requirements for other bioinks, the material should not possess toxicity, to ensure biocompatibility. The Feinberg laboratory takes an innovative approach to bioprinting by taking advantage of material properties to extrude bioprinted features within a hydrogel support bath. Their approach, coined freeform reversible embedding of suspended hydrogels, uses a gelatin slurry hydrogel bath embodying thermoreversible properties and Bingham viscosity to print soft, fragile constructs within the bath material. With the bath’s Bingham plasticity, the material acts as a solid unless a yield shear stress is attained, in which case the material will flow as a viscous fluid locally to the applied shear. In this way, extruded cross-linked printed structures (i.e., fibrin, alginate, and collagen type I via enzymatic, ionic, and pH mechanisms- respectively) within the hydrogel bath will immediately be surrounded by the supporting viscous material- enabling improved complexity in fabrication without collapsing or deforming under the printed structure’s own weight. In addition, the gelatin slurry possesses thermoreversible and biocompatible properties. The temporary support bath can be removed by raising the temperature from room temperature maintained during the printing process to a physiological range at 37 C (for gelatin specifically) to release the 3DP object supported within. Because the gelatin slurry is biocompatible, cells can be incorporated in the extruded bioink within the bath without concern regarding its effect on cell viability[23]. Similarly, others have achieved the bioprinting of intricate structures that otherwise cannot be achieved without support baths. Carbopol is another support bath material possessing desirable rheological properties in that the

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FIGURE 46.9 Supportive baths used in three-dimensional (3D) bioprinting: (A) The Feinberg laboratory developed the free-form reversible embedding of suspended hydrogels method to 3D printing (3DP) structures within a support bath. The bath is a Bingham plastic gelatin slurry that provides structural integrity to the print; an example is a human femur computed tomography scan that was downsized to create a printed version composed of alginate (scale ¼ 1 cm, 4 mm) [23]. (B) Angelini’s group demonstrates the capacity of cells to remain functional (viable, migrating, and proliferative) with their 3DP method [110] that expanded from their previous work in (C). In (C), a 3D object is printed within another gel support bath. This support bath, based on carbopol, fluidizes locally to the point at which shear is applied (i.e., the needle head extruding print material). The print can be released and form intricate structures such as the fine tentacles of a jellyfish [89]. (D) Christensen et al. use calcium chloride as a supporting and cross-linking bath for sodium alginate bifurcated objects with inkjet printing. This material supports cell encapsulation and allows for overhanging geometries [61]. (E) The Burdick laboratory uses photosensitive hyaluronic acid bioinks printed within a self-healing support hydrogel to form free-standing geometries, as shown by the spiral enveloping a channel [50].

transition from locally fluidized to solidified afforded by its Bingham plastic nature allows for the support of printed structures. Although the writing medium is more permanent (via photocross-linking or other means), the support bath can be dissolved with water [89,109]. The Angelini group has even expanded their carbopol support material to be cytocompatible for a range of cells (i.e., MSCs, human aortic endothelial cells) with enough elasticity to support cell division with the inclusion of cell media adjusted to pH 7.4 [112]. For those using inkjet printing for their bioprinted constructs, a common support bath system employs a CaCl2 solution that can dually function as a cross-linking agent and a support material. In this way, the alginate inkjetted

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deposition with or without cells is fabricated with the support bath instantly polymerizing the material to allow for complex structures with overhangs [61]. Yet another extrusion-based bioprinting approach, the Burdick laboratory developed a hydrogel bioink format that is directly written into a self-healing support hydrogel based on guestehost complexes. Their system is based on noncovalent and reversible bonds with the application of shear. By directly writing into this material, the bonds of the support bath are disrupted by the physical stimulus but are quickly reformed after the shear is removed. This enables the capacity of injectable hydrogels and extrusion 3DP. Their writing bioink based on HA was chemically modified for photocross-linking; more important, the support bath possesses shear thinning properties to provide structural support to print freestanding structures [50] (Fig. 46.9). Support baths must interact with the finalized structure in such a way that the material itself must still provide a minimal baseline of compatibility. Certainly, some approaches have been using the support bath as a reservoir of cells or perhaps cross-linking agents that will ultimately become part of the printed construct; therefore, the level of cytocompatibility is going to increase correspondingly. Regarding toxicity, beyond the baseline of cell compatibility in cell-laden prints, future studies might want to consider long-term effects on cytotoxicity. If residual components of support baths are cytotoxic, a few days after the print is not long enough to prove long-term cell viability and functionality. Finally, in parallel to sacrificial inks, these materials must be easily removable to decouple from the finalized printed construct.

Current Translation of Three-Dimensional Bioprinting 3D bioprinting has opened the door for opportunities in directly translating to the clinic. 3DP provides us with the ability to print tissue analogs by controlling the delivery of living cells and matching the appropriate material in a defined and organized manner. The control is promising because it is beyond defining the exact location as well as the sufficient number of cells in a multimaterial environment. 3DP has applications to tissue-engineering scaffolds, constructing cell-based sensors, physiological screening for drug and toxicity, and modeling tissue disease and tumors [100]. Here we will broadly characterize translation applications using in vitro and in vivo examples. In Vitro Applications In general, lab-on-a-chip style models can represent tissue analogs well by incorporating the many fluidic networks of our body (i.e., vasculature, lymphatics). An example of lab-on-a-chip, miniaturized, functional tissue mimics was discussed in the sacrificial section from the Khademhosseini laboratory. In their work, they incorporated embedded vascular networks in their 3DP constructs [67]. Many others have also dedicated their research to incorporating vascular fluidic networks, because vascularization remains a critical challenge in tissue engineering [11,29,86,103]. As discussed, beyond a few hundred microns of the diffusion limit, cells will not remain viable. However, even with many research groups studying vascularization strategies, many challenges remain, such as which vascular geometries to print and how to reach truly multiscale vasculature. By miniaturizing human tissue functional elements, studies can more accurately predict drug toxicity over animal models. Khademhosseini and colleagues delved into more complex in vitro models by developing liver-on-a-chip constructs. By developing a liver model, one can assess drug toxicity more adequately because the liver has the most important role in drug metabolism. Their liver model was biofabricated as a perfusable bioreactor with a direct-write printer to create hepatic spheroid-laden hydrogel structures [65]. In addition to better drug screening, in vitro models fabricated with 3DP approaches can better represent tumor models to help researchers elucidate cancer mechanisms. Efforts are being made to represent tumors 3D in models, because the progression of tumor metastasis is significantly different from that of a 2D counterpart [111]. The West laboratory modeled tumor angiogenesis with layer-by-layer tunable PEG hydrogels [112]. 3DP can provide cancer researchers with control over specific 3D microenvironments influencing nutrient transport and fluid shear stresses. With this kind of physiological mimicry, researchers can identify mechanisms directing tumor metastasis. A major challenge for 3DP is scale for tissue engineering translation: Printers often have difficulty achieving large, clinical-size organ analogs with the microdetail of the cellular organization necessary to be functional. The Atala laboratory reported an integrated tissue-organ printer that prints cell-laden hydrogels of desired mechanical stiffness with sacrificial hydrogels. Fabricated constructs presented include mandible bone, cartilage, and skeletal muscle. The printed anatomical shapes can be composed of multiple biomaterials and cells resulting in structures that in the future can be vascularized and included in more complex, solid organs [113]. Until then, these constructs may be difficult to incorporate in the clinic because encapsulated cells will quickly necrose without proper vascularization.

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For in vitro 3DP, clinical translation is mostly limited to drug screening and disease modeling. These applications can provide important information regarding addressing drug toxicity and to develop drugs. In Vivo Applications The promise of 3DP to fabricate tissue parts for implantation in the body is taking shape in in vivo applications. Tissue engineering has the potential to facilitate tissue regeneration by replacing injured parts and encouraging regrowth with the appropriate healing environment (i.e., growth factors, vascularization, stem cells). With 3DP, a diversity of critical-sized defects can be addressed tailored to the needs of the individual. In this section, we discuss studies that use 3DP to fabricate scaffolds applied in vivo. As mentioned previously, rapid vascularization is a limitation in applying tissue mimics in vivo, because cells will necrose soon after implantation without access to nutrients and oxygen. Several studies achieve rapid vascularization in vivo enabled by prevascularization strategies fabricated with 3DP techniques [114,115]. Using sacrificial templating with a sacrificial carbohydrate bioink, an open microvascular network can anastomose in line with the rat femoral artery with a surgical technique [11,114]. Similarly, others employ 3DP biodegradable scaffolds to anastomose built-in vasculature directly and surgically in an AngioChip to the host vasculature of the rat hind limb femoral vessels [115]. Patency is maintained through the fabricated vessels within the scaffold. These studies demonstrate the potential of prevascularizing tissue analogues with 3DP for rapid vascularization in vivo. Other tissues have been fabricated and implanted in vivo, such as bone. Bone scaffolds 3DP with MSCs, bone morphogenetic protein growth factor, and PCL as the bioink have been implanted in critical-sized bone defects of sheep. After 3 and 12 months, bone healing progresses with signs of vascularization ingrowth, mineralization, and appropriate mechanical stiffness [76]. By considering important aspects of biocompatibility and degradability, many of the bioink principles described earlier here have resulted in the success of 3DP tissue mimics for in vivo implantation. The opportunities to apply 3DP to in vivo applications are still being explored. Integration with the host tissue is of utmost importance for all in vivo studies. For researchers, great care is taken to select the most appropriate bioink during fabrication to allow the best possibility for successful integration. Future work is necessary to ensure the functionality of the intended tissue analog taking into consideration the role of the parenchyma.

CONCLUSION AND FUTURE DIRECTIONS This chapter introduces current 3DP techniques and describes the materials used to enable bioengineering applications. For bioprinting, the materials, or bioinks, used for fabrication must be biocompatible as well as printable. Bioinks are printing materials that will interface with biological components or are part of the fabrication process of a construct that will come into biological contact. We describe three separate categories for bioinks: (1) matrix or matrix-mimicking, (2) sacrificial, and (3) supporting. Although there has been great progress in synthesizing biocompatible and printable materials, future work is needed to address the greatest limitations in bioengineering. For one, vascularization must be rapidly integrated. Many of the challenges for vascularization strategies involve scale to realize multiscale hierarchical vessels. Reaching capillary-sized vessels requires microscale resolution in printing. However, many printing techniques that can even reach these small levels are not able to print rapidly. Another limitation in current studies is the use of physiologically relevant cell types. Fortunately, induced pluripotent stem cells hold much promise in alleviating this limitation. Materials available as bioinks are limited. Generally, the bioinks used are a compromise between structural strength and biocompatibility. Therefore, current and future work follows a multimaterial approach to achieve the desired properties. The advent of 3DP applied to bioengineering and tissue engineering has pushed researchers to create complex biological structures with incorporated fluidic networks with multiple cells and materials for the most physiologically mimicking environment. This is an exciting time for the field to realize approaches that have a clinical impact. From rapidly reiterating studies for drug screening to achieving high spatial resolution to fabricate tissue analogs, 3DP holds much promise for therapeutic translation. The choice of bioinks will continue to have a fundamental role in determining ultimate biocompatibility.

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List of Acronyms and Abbreviations 2D Two-dimensional 3D Three-dimensional 3DP Three-dimensional printing CaCl2 Calcium chloride dECM Decellularized extracellular matrix ECM Extracellular matrix EDTA Ethylenediaminetetraacetic acid gelMA Gelatin methacrylate HA Hyaluronic acid hCMPC Human cardiac-derived cardiomyocyte progenitor cell MSC Mesenchymal stem cells MTU Muscleetendon unit PCL Poly(ε-caprolactone) PDMS Polydimethylsiloxane PEG Poly(ethylene glycol) PEGDA Poly(ethylene glycol) diacrylate PEGX Poly(ethylene glycol) with reactive ends PGA Poly(glycolic acid) PLA Poly(L-lactic acid) PLGA Poly(D,L-lactic-co-glycolic acid) PPF Poly(propylene fumarate) SLS Selective laser sintering UV Ultraviolet

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47 Three-Dimensional Tissue and Organ Printing in Regenerative Medicine Gregory J. Gillispie, Jihoon Park, Joshua S. Copus, Anil Kumar Pallickaveedu Rajan Asari, James J. Yoo, Anthony Atala, Sang Jin Lee Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

INTRODUCTION Attempts have been made in tissue engineering to develop biological substitutes to address the high shortage of tissues and organs for transplantation. Although tissue engineering has progressed rapidly over the past 2 decades, conventional fabrication methods are limited in their ability to create clinically applicable tissue constructs with well-interconnected pores, complex architectures, patient-specific geometries, and heterogeneous material distributions. Over the past few years, three-dimensional (3D) bioprinting strategy has been applied to overcome these limitations. It enables the fabrication of tissue constructs with unique spatial control over the deposition of cells, biomaterials, and bioactive molecules such as growth factors, cytokines, peptides, and small molecules, resulting in higher regenerative capability after implantation [1,2]. 3D printing technology was developed in the 1980s and included various approaches to creating objects from a computer-generated file [3]. This technology quickly became a powerful tool in tissue engineering and biomedical research [4]. In structurebased bioprinting, bioinert or bioactive materials such as metals, ceramics, and polymers are used to develop a tissue structure followed by precisely depositing cells and bioactive molecules onto it [5,6]. In cell-based bioprinting, a high density of cells is patterned spatially with a prescribed organization in a layer-by-layer fashion, forming tissue constructs [5]. Thus, additive biomanufacturing or 3D bioprinting allows the creation of tissue-specific architectures with precise geometries that have been limited using conventional fabrication methods. Landers et al. first introduced 3D bioprinting as an extrusion-based method to dispense cells continuously within a hydrogel material (bioink) from a dispensing head to a stage-based patterns predesigned through computer-aided design/computer-aided manufacturing (CAD/CAM) tools [7,8]. Various types of 3D bioprinting methodologies are available to meet specific requirements in tissue engineering applications.

BIOPRINTING STRATEGY: FROM MEDICAL IMAGE TO PRINTED TISSUE Bioprinting aims to achieve reproducible, complex tissue structures that are well-vascularized and suitable for future clinical use. Human tissues and organs have arbitrary 3D shapes composed of multiple cell types and an extracellular matrix (ECM) with a functional organization. CAD/CAM processes are important technologies necessary for future clinical applications of 3D bioprinting because these processes provide an automated way to replicate the 3D shape of a targeted tissue structure. In general, the process starts by scanning the patient to produce 3D volumetric information about a target object using medical imaging modalities such as computed tomography

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FIGURE 47.1 Three-dimensional (3-D) bioprinting strategy from medical imaging to printed tissue construct: (A) schematic diagram and (B) example of computer-aided design (CAM)/computer-aided manufacturing process for automated printing of 3D shape imitating target tissue or organ. A 3D CAD model developed from medical image data generates a visualized motion program, which includes instructions for XYZ stage movements and actuating pneumatic pressure to achieve 3D bioprinting. CT, computed tomography; DICOM, Digital Imaging and Communications in Medicine; MRI, magnetic resonance imaging; STL, stereolithography.

(CT) and magnetic resonance imaging. These imaging tools acquire information from cross-sectional slices of the body and the data are stored in the Digital Imaging and Communications in Medicine format, which is a standard format for digital imaging in medicine. This information is transformed into a 3D CAD model by reverse engineering. This process starts by interpolating points within and between image slices to improve the resolution and generate voxels from the measured data. This CAD model is created by extracting localized volumetric data from a targeted tissue structure to generate a surface model. In this step, sophisticated reconstruction of the CAD model is required for bioprinting owing to the complexity of the tissue or organ. A motion program, which is instructional computer codes for the printer to follow designed paths, is generated with a CAM system. This CAM process is divided into three steps: slicing, the tool path, and the generation of the motion program. Slicing obtains information about sliced 2D shapes of an object for the layer-by-layer process. Then, tool path generation creates a path for the tool to follow to fill the cross-sectional space of each layer. The printed tissue-specific structure has the proper inner architecture constructed with multiple cellular components for efficient tissue regeneration. Therefore, a well-organized strategy for tool path generation is required to have high efficient tissue regeneration and is an important process for 3D bioprinting. Fig. 47.1 shows the 3D bioprinting strategy from the medical image to the printed tissue constructs developed by CAD/CAM process and automated printing of a 3D shape imitating a target tissue or organ.

BIOPRINTING MECHANISMS Researchers have developed a variety of printing mechanisms, all of which aim to accomplish the same goal: printing 3D human tissue or organ structures for reconstruction. The effectiveness of each printing mechanism relies heavily on the choice of biomaterials and the targeted applications. Bioprinters consist of three main components: a three-axis stage, printing cartridges, and the dispenser. Stage controllers are used to move the printer head in the X, Y, and Z directions. Printing cartridges, usually in the form of a syringe, store either the polymeric components of the scaffold or the cellular/hydrogel components. They contain the nozzle, which determines the amount of material dispensed at set printing parameters. The dispenser is the final component; it varies among printing methods, but it is the mechanism that deposits materials (Fig. 47.2).

Jetting-Based Printing Jetting-based printing, or “inkjet” printing, is the most commonly used printing mechanism for both nonbiological and biological objects [9]. Similar to an inkjet printer used to apply ink to paper, jetting-based bioprinters dispense a controlled volume of liquid to a predefined location through noncontact deposition (Fig. 47.2A). The “ink” in this instance is usually a hydrogel that may or may not contain cells and can be dispensed in

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FIGURE 47.2 (A) Microscopic view of a three-dimensional multicell “pie” construct printed by a jetting-based mechanism. Individual cells were printed in predetermined locations and then located by microscopy at those locations after printing [124]. (B) Auricular implant was printed through extrusion-based mechanisms with cell-laden hydrogel, poly(ε-caprolactone), and Pluronic F-12 [47]. (C) Human umbilical vein endothelial cells were printed precisely in straight lines through laser-assisted bioprinting to study cell migration rates [125]. PBS, phosphatebuffered saline.

volumes between 10 and 150 pL, depending on the dispensing modules used [10]. The two main dispensing methods for jetting-based printing are thermal and piezoelectric inkjets. Thermal inkjets use an electric heater that generates small bubbles in the printhead. These bubbles collapse to create pressure pulses that force droplets of liquid out of the nozzle. Although thermal inkjets use heat at around 200e300 C, the duration of heating is typically around 2 ms, which studies have found to result in only a 4e10 C rise in hydrogel temperature [9]. Piezoelectric inkjet printers use a polycrystalline piezoelectric ceramic to create the pressure pulse that ejects the droplet [10]. The volume of liquid dispensed depends on the temperature gradient, the frequency of the pressure pulse, and the ink viscosity. To determine where the material will be dispensed, the bioprinter can control the X and Y positions of the nozzle heads to a scale of micrometers, which allows for precise, high-resolution printing. A major drawback of jetting-based printing is the amount of shear stress induced on the cells as they are forced through the nozzle. Shear stress creates a huge risk to damage cell membranes and causes cell lysis [10]. Another drawback of this printing is that to dispense materials, it has to be in a liquid form, so to form a solid structure, the liquid has to be cross-linked during or very soon after printing. This causes a significant reduction in printing speed, because a time for cross-linking must be determined. It is also considered that high cell concentrations (greater than 10 million cells/mL) often result in nozzle clogging and some instances may even alter the properties of the hydrogel so much that it can no longer be cross-linked [11,12]. Despite these drawbacks, jetting-based printing offers many advantages over laser-assisted and extrusion-based printing methods. Commercial inkjet printers are readily available, and because of this, many laboratories can modify the printers so that they can experiment with bioprinting at a low cost. The printers also offer high printing resolution, which allows concentration gradients of cells, biomaterials, or bioactive molecules to be introduced throughout the structure by altering drop sizes and densities [9]. The droplets can be arranged into patterns such as lines 50 mm wide or single droplet patterns containing one or two cells [13]. The droplet size can be controlled electronically and can range from less than 1 pL to more than >300 pL in volume with deposition rates from 1 to 10,000 droplets per second [14].

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Extrusion-Based Printing Unlike jetting-based printing, extrusion-based printing is an additive manufacturing mechanism that relies on fused deposition modeling and requires direct contact with a surface to print (Fig. 47.2B). These printers mainly include a temperature-controllable dispensing system and a stage, with one or both capable of movement in the X, Y, and Z axes. Some 3D extrusion printers have multiple dispensers that allow them to switch printing material without going through the process of retooling [15]. Extrusion-based printing methods offer improved printing resolution, speed, and spatial control. These printers also differ from inkjet printers in the manner in which they deposit material, because they dispense in a continuous string as opposed to individual droplets. The printer extrudes a 2D pattern designed by CAD/CAM software, and then each layer serves as the foundation for the next layer above it. Three systems can be used to extrude the material: pneumatic, piston, or screw-based dispensing systems [9]. Pneumatic dispensing systems can deposit high viscous materials, but the drawback is that there is a delay in deposition owing to the time it takes to compress the gas in the cartridge. The ultimate force of pneumatic systems is limited only by the air-pressure capabilities of the system. Mechanical systems such as a piston and screw-based dispensers provide more spatial control at the cost of reduced maximum force capabilities [9]. Extrusion-based printing methods have the widest range of biomaterials that can be printed; they vary from hydrogels to polymers and cell aggregates. Materials with viscosities ranging from 30 mPa/s to greater than 6  107 mPa/s have been successfully printed with extrusion-based printers, with high viscous materials acting as structural support and low viscous materials providing an adequate environment for cell printing [16]. This printing method is also able to print a high density of cells; however, cell viability may be decreased with smaller nozzle sizes and a high pressure level. The main challenges to extrusion-based printing mechanisms are the low printing resolution and low printing speed: 5- to 200-mm pattern resolution at linear speeds of 10e50 mm/s [17].

Laser-Assisted Printing Laser-assisted bioprinters operate by focusing a laser pulse toward an absorbing layer, typically gold or titanium, to generate high-pressure bubbles that propel cell-containing bioinks toward a collector substrate [9]. A standard laser-assisted printing system usually consists of a pulsed laser beam, a focusing system, a “ribbon” that has a transport support made from a laser energyeabsorbing layer, a layer of hydrogel or cellular material, and a substrate-receiving layer (Fig. 47.2). Laser-assisted printing method has successfully transferred peptides, DNA, and cells [18]. The main advantage of this method is that it is nozzle-free; therefore, it avoids nozzle clogging issues found in other bioprinting systems. Laser-assisted methods can deposit materials with viscosities ranging from 1 to 300 mPa/s and cell densities close to 108 cells/mL, with resolutions close to a single cell per drop without significant effects on cell viability or function. In addition, many factors affect the printing resolution. Some of these factors include the energy delivered per unit area owing to the laser, the surface tension, substrate wettability, the gap between ribbon and substrate, and the thickness and viscosity of the biological layer [9]. There are also a lot of disadvantages associated with the laser-assisted printing system. One drawback is that each ribbon must be prepared following a time-consuming process that may become overwhelming if multiple cells or hydrogels must be codeposited [9]. It can also be difficult to target and position the cells accurately because of the nature of the ribbon cell coating. Metallic residues may be present in the final construct owing to vaporization of the laser-absorbing layer, although there have been methods to reduce this contamination, including using nonmetallic absorbing layers and altering the printing process so that an adsorbing layer is not needed [19].

Hybrid and Other Mechanisms Hybrid fabrication is a combination of technologies that are practiced individually. Despite rapid growth over 2 decades, current 3D bioprinting techniques and systems fall short in integrating rigid and soft multifunctional components. In some instances, the advantages of one technology might not be sufficient to meet the requirements of creating targeted functional tissue constructs. Merging technologies may overcome this limitation and improve the bioprinting process. The combination of technologies can be at the cellular, bioink, or bioprinter level. Tan et al. introduced a 3D bioprinting strategy by integrating conventional scaffold-based fabrication and the cell-based bioprinting approach [20]. They explored the use of hydrogel-encapsulated, cell-laden microspheres as the bioink for 3D bioprinting. The hydrogel lubricates and glues the microspheres during printing and fuses them together

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after printing upon gelation. Kucukgul et al. developed computational algorithms to generate support structures for cylindrical cell aggregates to print fully cell-based vascular structures [21]. Xu et al. proposed a combination of electrospinning and inkjet bioprinting techniques to develop layered cartilage constructs [22]. Combining the two principles of electrospinning and bioprinting resulted in the fabrication of a cartilage construct with appropriate cell function and mechanical properties. Shanjani et al. developed a new hybrid 3D bioprinting technology called Hybprinter by integrating soft and rigid components [23]. This technique employs a digital light processinge based stereolithography component and molten material extrusion techniques. Poly(ethylene glycol) diacrylate (PEGDA) was used as soft hydrogel and poly(ε-caprolactone) (PCL) as the scaffold’s structural support. Mendoza-Buenrostro et al. reported a hybrid bioprinting technique for the fabrication of scaffolds with topography scales ranging from nanometers to millimeters [24]. Scaffolds were produced by extrusion-printing PCL embedded with nanofibers.

BIOMATERIALS FOR BIOPRINTING: BIOINKS Printability When used in bioprinting, hydrogels that can encapsulate and deliver cells and bioactive molecules through printing mechanisms are referred to as bioinks. In extrusion-based bioprinting, bioinks should have strong shear thinning behavior, which means that viscosity decreases under shear strain [25]. They should also have low surface tension so that extrusion is possible without the formation of droplets at the needle tip. In addition, bioinks should have proper viscoelastic properties that allow them to be stable enough to be printed in multiple layers before cross-linking. Traditional hydrogel precursors are low-viscous solutions that are cross-linked either during or after the printing process. Cross-linking before printing increases the shear stress on the cells, resulting in cellular damage and nozzle clogging. Cross-linking after printing affects the resolution because of the spread of bioink in the time between extrusion and cross-linking and can lead to incomplete cross-linking deep in large, multilayered constructs [9]. Fig. 47.3 shows a schematic diagram of variables, including printability, in 3D bioprinting. There are structural properties that must be satisfied so that the printed construct maintains shape, accuracy, and integrity after printing is finished. The printed construct should maintain its predesigned structure, which includes shape, resolution, orientation, and mechanical properties. Because bioprinting aims to build a 3D tissue construct with cells, cell viability after printing is one of the main criteria for the printability of bioinks. For the cells to survive, a biologically favorable microenvironment is required so that the cells can be well-preserved not only during the printing process but also afterward in culture. In hydrogel-based bioinks, cell density, diffusion coefficient, temperature, and humidity can significantly affect the printability of the bioinks [26e28].

Hydrogel-Based Bioinks for Cell Printing Hydrogels are naturally derived or synthetically produced. Hydrogel networks consist of polymer chains or peptide chains. They are highly moist and ideal for absorbing high levels of nutrients and oxygen, enabling cells to survive within the construct and diffuse waste. Therefore, hydrogels have been printed by a variety of mechanisms to deliver cells in the printed constructs. In the extrusion-based printing, the required properties of hydrogel-based bioinks during the printing process are (1) high viscosity to provide homogeneous cell suspension and initial structural integrity, (2) strong shear-thinning behavior to minimize cell damage, and (3) rapid gelation mechanism to stabilize a 3D-printed structure. If the gelation time is too long, the spatial resolution is lost and the layer cannot be printed correctly. To control the gelation time, the cross-linking mechanisms can be manipulated by chemically modifying the material, introducing a cross-linking agent or changing the polymer content [29]. Synthetic Hydrogels Synthetic hydrogels used as bioinks have low cytotoxicity, controlled biodegradability, and good mechanical properties; however, most synthetic hydrogels have low biological properties that interact minimally with cells. Pluronic F127, a thermosensitive hydrogel, undergoes a phase transition at room temperature and becomes a viscous substance [30]. When the concentration of Pluronic F127 is 25% w/v or more, it can be dispensed with high printability. Although high-resolution of the printed structures can be achieved using Pluronic F127, the structure can

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FIGURE 47.3 Schematic diagram of variables critical to three-dimensional bioprinting strategy. The hydrogel-based bioinks determine the viscosity, gelation mechanism, and printing parameters, and eventually the bioprinted tissue structures.

easily be collapsed in the culture condition. To improve its mechanical stability, Pluronic F127 has commonly been used chemically modified as a photocross-linkable hydrogel [31]. Poly(ethylene glycol) (PEG)-based hydrogels can be chemically modified to control their biological and biomechanical properties. Therefore, PEGs have been used as bioink materials in 3D bioprinting by introducing various functional motifs attached to the terminal end of PEG [32] or combining them with other hydrogels [33,34]. Naturally Derived Hydrogels Natural hydrogels are classified into mainly proteins and polysaccharides. Most of them are present in the body, so they show high biological properties and do not cause a severe immune response. Collagen (type I) is the most abundant component of ECMs; it contains cell-guiding chemical cues such as the cell adhesion peptide sequence arginine-glycine-aspartic acid. Under the appropriate temperature and pH, a pure collagen solution physically forms a gel with properties dependent on the solution concentration. However, collagen itself has rarely been used as a bioink material owing to low viscosity and poor mechanical stability. To overcome these limitations, collagen has been mixed with various other hydrogels such as agarose, chitosan, fibrin, and hyaluronic acid (HA) [22,35e37]. An approach to printing skin cell-containing collagen solution from separate nozzles using an inkjet microvalve dispensing method was reported [38]. In the printing process, the collagen solution remained acidic and cooled. Multiple layers of cell-laden collagen bioink were printed and then treated with aerosolized sodium bicarbonate (NaHCO3) to buffer the pH toward neutral for gelation. The results showed high cell viability at 1 day after printing for keratinocytes and fibroblasts, which indicated the survival of cells and spatial control of the printing approach, which is needed to offer a functional skin replacement. Gelatin is a substance in which water-insoluble collagen is made soluble by high temperature or acidebase treatment. It forms a thermoreversible hydrogel. Moreover, chemical modification of gelatin with unsaturated methyl methacrylate results in gelatin methacrylate (GelMA), which can form covalently cross-linked hydrogels

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under UV exposure [39]. Photocross-linkable GelMA is the most popular bioink material for cell bioprinting because of high printability and tunable mechanical properties. GelMa was used to print a complex architecture that contained various cell types and vasculature [40]. A cell-laden GelMa and sacrificial Pluronic F127 were dispensed and embedded within the GelMa block in a predetermined 3D structure. Afterward, the printed structure was cross-linked by UV exposure. The sacrificial Pluronic F127 was removed at 4 C by phase transition to create open microchannels within the GelMa block. Finally, a suspension of human umbilical vein endothelial cells (HUVECs) was seeded into the open microchannels. This approach allowed for the viable deposition of cells in a 3D structure with a microvessel-like structure that was covered by endothelial cells (ECs) to provide nutrients to surrounding cells. HA is the most abundant of the glycosaminoglycan (GAG) family; it has the repeating structure of glucuronic acid and N-acetyl-glucosamine disaccharide. HA has a high molecular weight and a large amount of branching, which allows for intermolecular hydrogen bonding and high viscosity. Like other polysaccharides, HA can support cell survival but has low cell-binding motifs. Moreover, HA itself has low structural integrity and shape fidelity after printing; therefore, HA has been chemically modified for cross-linking or mixed with other hydrogels [41]. Fibrinogen, which is a glycoprotein, is reacted with thrombin to convert into fibrin network self-assemblies [42]. Fibrin has many cell-binding motifs that enable cell attachment and vulnerability to proteases for remodeling. Fibrin-based bioink has been used by laser-induced forward transferebased printing for a 3D multicellular array [43]. To improve the gel stability, HA was added to a fibrinogen solution to print the arrays. Endothelial colony-forming cells (ECFCs) were printed along with adipose-derived stem cells (ASCs) in a 3D array to observe cell-cell interactions. These droplet arrays were printed onto a layer of fibrinogeneHA that was spraytreated with a thrombinecalcium chloride solution to induce gelation. The cell-laden droplets were converted into fibrineHA as they encountered the treated substrate with a residual thrombin solution. Results showed that ASCs initially migrated toward ECFCs with no evidence of ECFC sprouting or migrating. Once ASCs contacted the ECFC aggregates, an explosion of ECFC network sprouts began to extend from the initial droplet position and remained stable networks for several weeks. Alginate is a naturally derived anionic polysaccharide that exhibits gelation in the presence of bivalent ions such as Ca2þ [44]. Alginate hydrogel has served as a cell delivery material for many tissue engineering and drug delivery applications owing to its ease of preparation and relatively good cell compatibility; however, a major drawback is the lack of mammalian enzymatic degradation, which limits tissue remodeling when implanted. Because the solubility of alginate is low, it is often mixed with other materials to increase the printing resolution as a bioink. For example, alginate solution mixed with cellulose nanofibers improved shear fluidization and viscosity, resulting in high printing resolution. In addition, alginate was cross-linked with the divalent cation Ca2þ to stability the printed construct. Tissue-specific, ECM-based bioinks have been developed and tested. Technically, ECMs obtained from various decellularized tissues are pulverized and solubilized as bioinks [45]. The approaches combined with tissue-specific ECM-based bioinks and tissue-specific cell types have been shown to be valuable for recapitulating anticipated tissue features [45,46]. However, the ECM solution inherently has a low viscosity and exhibits low shape fidelity and structural stability. Thus, various attempts have been made to enhance their chemical and physical properties in 3D bioprinting.

Biodegradable Synthetic Polymers for Structural Integrity Synthetic polymers have some advantages for applications in tissue engineering and 3D bioprinting. These polymers can be synthesized with reproducible quality and fabricated into various structures with predetermined bulk and surface properties. Additional advantages include the ability to tailor their biomechanical properties and biodegradation kinetics for various biomedical applications. In bioprinting, synthetic polymers such as PCL, poly(lactide-co-glycolide) (PLGA), and poly(lactic acid) can provide mechanical strength, thereby overcoming limitations of the hydrogel-based constructs regarding size, shape, and structural integrity [47]. For the extrusion method, melted polymers or polymer solutions with proper viscosity are needed. Therefore, PCL is the most commonly available polymer for extrusion-based bioprinting because of its low melting point of 60 C and high printability. Many research groups are exploring the synthesis of polymeric biomaterials for 3D bioprinting.

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Scaffold-Free Cell Printing 3D tissue structure can be printed using multicellular aggregates without supporting materials [48]. This method enables cell aggregates to be dispensed from a capillary to form a 3D structure. Norotte et al. reported a scaffold-free cell printing method using multicellular spheroids consisting of smooth muscle cells (SMCs) and fibroblasts and agarose hydrogel for temporary support [49]. After printing, the printed cell aggregates were fused and formed a small-diameter vessel-like tube ranging from 0.9 to 2.5 mm in diameter. Li et al. fabricated vertical channels without using a support material by taking advantage of the gelation behavior of gelatin combined with cross-linking alginate or fibrinogen [50]. Itoh et al. proposed a method to fabricate a scaffold-free vascular tube using cell aggregates [51]. Predesigned 3D tubular structures were constructed with cell aggregates consisting of ECs, SMCs, and fibroblasts. The multicellular aggregates self-organized and fused into a tubular structure that was perfused with a bioreactor. To avoid droplets from forming at the end of the nozzle owing to surface tension, Ozler et al. developed a quantitative model to predict the success of cell aggregate extrusion [48]. This approach can be repeated for different cell types after obtaining their respective rheological properties. In addition, cell viability can depend on the compression ratio applied during printing. Thus, it is necessary first to investigate the impact of compression on their survival rate and cellular functions.

THREE-DIMENSIONAL BIOPRINTING IN REGENERATIVE MEDICINE APPLICATIONS Three-Dimensional Bioprinted Vascular Structures A major limitation of bioengineered tissue constructs is the lack of proper vascularization into the implanted constructs. Fully vascularized tissue constructs are required to attain long-term cell survival and tissue functions. 3D tissue constructs packed with metabolically active cells can rapidly form necrotic cores in the absence of a vascular network [52]. The transport of nutrients and other physiologically relevant molecules toward and away from a large tissue construct is still limited [53,54]. When a 3D tissue construct with cells is implanted, efficient mass transfer requires an intact microvasculature to maintain the metabolic functions of cells deep inside the construct. Indeed, the ingrowth of a microvascular system into the implanted bioengineered tissue constructs in a timely manner is the key to success in clinical use [9]. Hence, many attempts have employed 3D bioprinting as a promising technology to fabricate vascularized tissue constructs (Fig. 47.4) [40,49,55].

FIGURE 47.4 Examples of the three-dimensional bioprinted vasculature. (A) Vascular unit printed by Miller et al. [56] The construct shows

three compartments consisting of the vascular lumen, the endothelial cell (EC) lining, and the matrix-encapsulated cells. Endothelialized channel walls and the intervessel junction surrounded by human fibroblasts. (B) Engineered tissue construct developed by Kolesky et al. [40] showing channels lined with human umbilical vein endothelial cells (HUVEC) (red) and human neonatal dermal fibroblasteladen gelatin methacrylate (green). (C) Hydrogel construct with embedded, connected hollow channels reported by Attalla et al. [57] and cross-section of hollow channel embedded within HUVEC-laden hydrogel. (D) Deposition of ECs and fibrin channel scaffold using thermal inkjet printer described by Cui and Boland [55]. Fluorescent image shows the printed microvasculature after 21 days.

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One of the earliest and simplest approaches to fabricating a vascularized tissue construct is to use a sacrificial component that acts as a structural role during printing and is later removed to create a hollow tunnel. Miller et al. reported a vascular casting approach with carbohydrate glass as the sacrificial template; however, vascular networks developed by carbohydrate glass are limited in size and culture duration because of the practical difficulty of direct perfusion [56]. Kolesky et al. attempted to overcome these limitations by fabricating 3D constructs complete with a vasculature including multiple cell types and ECM proteins [40]. Attalla et al. modified a printhead with a microfluidic unit and generated instantly perfusable vascular networks integrated within a cell-seeded hydrogel [57]. The vascular structures were made from alginate using a coaxial nozzle. Cui and Boland reported an inkjet-based printing method with a precise droplet volume of 130 pL to fabricate microchannels [55]. Human microvascular ECs combined with the printing of biomaterials were found to align themselves inside the microchannels and form a confluent microvascular lining. Although bioprinting technologies have evolved to a level of creating complex vascularized constructs with multiple cell types and ECM proteins, reconnection of the vascular structures to the host circulatory system remains challenging. Lee et al. developed a novel bioprinting approach to creating a network from capillaries to large perfused vascular channels [58]. This was achieved by angiogenic sprouting from the large vessel through a natural maturation process. However, these printed constructs are often thin or hollow structures, so they are nourished by diffusion from the host vasculature but not by anastomosing [59]. When the diffusion limit needed by engineered tissues exceeds 150e200 mm, a precise vascular network must be embedded into fabricated constructs, a feat that has not yet been accomplished [60,61]. Efforts to simplify the complex fabrication methods and find new technologies for bioprinting vascular structures are in great need by tissue engineering as a whole.

In Vitro Tissue Models Drug discovery is an inefficient process with a high failure rate and an extreme financial burden. Animal studies do not always indicate the results in human trials, and the regulatory environment is becoming stricter as time progresses. In addition, from a moral standpoint, attempts should be made to reduce the number of animal studies conducted. In vitro studies, mostly 2D cell culture methods are also severely limited. Drug response, gene expression, migration, morphology, and viability have all been shown to differ between 2D and 3D environments. Much emphasis has been placed on creating in vitro 3D tissue models to overcome these limitations. Typically, this is done by suspending cells or organoids (or cell aggregates) in a 3D culture within a singular or entire array of microfluidic devices. 3D in vitro assay systems have advanced immensely to a level in which living constructs can closely mimic the native tissue environment in a high-throughput platform. Although several fabrication techniques have been used to develop these models, 3D bioprinting technologies are advantageous owing to their low cost and efficiency, high throughput, excellent reproducibility, and ability to create complex geometries. The two major areas in which 3D printed in vitro tissue models have been applied are cancer research and drug screening systems. Table 47.1 summarizes bioprinting applications for in vitro biological systems. Tumor Models 3D bioprinting of cells as tumor models are helpful for studying the interaction of immune and tumor cells and for screening new treatments [62]. Xu and Celli et al. introduced a 3D in vitro cancer model using human fibroblasts and ovarian cancer cells on a Matrigel matrix; they showed precise and reproducible control over the cell density and spacing compared with manual ejection by micropipettes [63]. Snyder et al. introduced a similar model in 2011 with human hepatic carcinoma cells and mammary epithelial cells. Cells and the microfluidic device were printed to test the radiation shielding of the prodrug amifostine. Radioprotective benefits for the liver were seen in their in vitro model [64]. Huang et al. examined tumor cell migration in a honeycomb structure with different channel widths (25, 45, and 120 mm) to mirror those of natural blood vessels [65]. HeLa and 10T1/2 cells were seeded within the device and were evaluated in the different channel sizes. HeLa cancer cells showed less morphological changes between channel sizes than did 10T1/2 cells and migrated at higher rates as the channel size decreased. Also looking at bioprinted HeLa cells, Zhao et al. examined the cell response after extrusion in a gelatinealginateefibrinogen hydrogel compared with a 2D culture model. Cells in the 3D model showed higher proliferation, matrix metalloproteinase expression, and chemoresistance [28]. Hribar et al. used 3D projection printing to create concave PEG structures that formed and maintained breast cancer spheroids for long-term culture [66]. The breast cancer spheroids exhibited necrotic, hypoxic cores, which are key components of the tumor in vivo microenvironment.

TABLE 47.1

Applications of Bioprinting In Vitro Biological Systems

Tissue/Organ

Bioprinting Method

Cell Type

Tumor/cancer

Pneumatic cell droplet patterning

Fibroblasts and ovarian cancer cells

Tumor/cancer

Temperature controlled, pneumatic extrusion

Tumor/cancer

Encapsulation Material

Testing

Outcomes

References

Matrigel

Reproducibility and precision of cell density and spacing

Bioprinting methods showed improved performance compared with micropipette ejection

[63]

Hepatic carcinoma and mammary epithelial cells

Matrigel

Radiation shielding capabilities of the prodrug amifostine

Amifostine provided radioprotection to cells, with greatest benefit seen in dual-cell model

[64]

Projection stereolithography (UV exposure)

HeLa and 10T1/2

Poly(ethylene glycol) diacrylate

The effect of channel width on tumor cell and 10 T1/2 migration and morphology

Tumor cells showed increased migration speed and less change in morphology with smaller channel sizes compared with 10 T1/2

[65]

Tumor/cancer

Extrusion

HeLa

Gelatinealginatee fibrinogen

Viability, proliferation, MMP expression, and chemoresistance versus 2D culture

3D cell culture increases cancer cell proliferation, MMP expression, and resistance to chemotherapy

[28]

Tumor/cancer

Continuous 3D projection (with nonlinear exposure)

Breast cancer cells

Poly(ethylene glycol)

Long-term culture and validation of breast cancer spheroids

Spheroids showed hypoxic cores and signs of necrosis, key features of tumor environment

[66]

Liver

Extrusion

Hepatocytes

Alginate

Validation of liver cell activity and metabolic performance

Cells were viable, proliferative, synthesized urea, and metabolized 7-ethoxy-4-trifluoromethyl coumarin to 7-hydroxy-4trifluoromethyl coumarin

[67,68]

Bacterial Infection

Inkjet

Escherichia coli

Alginate

Treatment of E. coli with several common antibiotics

Similar results to current lowthroughput, less reproducible, and more expensive methodologies

[69]

Brain

Extrusion and subsequent dissolution of sacrificial resin

Mouse brain endothelial cells

Type 1 collagen microchannels

Model validation with transendothelial permeability measurements and hyperosmotic mannitol disruption test

Permeability decreased over 3 weeks of culture and then recovered over 4 days after hyperosmotic mannitol disruption

[70]

Lung

Extrusion

Endothelial and epithelial cells

Matrigel

Cell viability, distribution, morphology, and permeability compared with manually placed cells

Printed constructs resulted in more homogeneous cell distributions, proper cell morphologies, lower permeability, and similar viability

[71]

3D, three-dimensional; MMP, matrix metalloproteinase.

THREE-DIMENSIONAL BIOPRINTING IN REGENERATIVE MEDICINE APPLICATIONS

841

Drug Screening Systems The number of in vitro drug screening systems has increased immensely in both quality and quantity, ranging from hepatic cells suspended in a microfluidic device to integrated, multiple-tissue, body-on-a-chip systems. Chang et al. used direct cell writing bioprinting to create a 3D microorgan housed using soft lithographic micropatterning [67]. Alginate-encapsulated hepatocytes printed in the microfluidic device were viable and proliferated and were capable of synthesizing urea. This work was furthered by infusing a hepatocyte containing a microfluidic device with 7-ethoxy-4-trifluoromethyl coumarin, which was metabolized into 7-hydroxy-4-trifluoromethyl coumarin, mimicking the in vivo behavior of the liver [68]. Rodriguez-Devora et al. developed an inexpensive drug-screening platform via inkjet bioprinting deposition. Escherichia coli were printed in an alginate solution with different antibiotic droplets patterned on the cells, resulting in similar bacteria inhibition compared with the current screening process [69]. An in vitro model of the bloodebrain barrier was developed by Kim et al. Mouse brain ECs were cultured within an array of type I collagen microchannels fabricated using microneedles on a 3D printed frame [70]. The model was validated by measuring transendothelial permeability and a disruption experiment by hyperosmotic mannitol. In addition to the bloodebrain barrier, an aireblood barrier was developed to model the lung. Horva´th et al. developed a model by bioprinting epithelial cells and ECs separated by a basal membrane layer [71]. The bioprinted model was more reproducible and had thinner cell layers than that which could be manufactured using traditional manual methods. In addition to these already successful in vitro models, several areas for future work stand out in this new field. Work is in progress to incorporate an array of tissue types into the same drug screening platform. Bioprinted in vitro models are also in good position to evaluate gene therapy techniques; however, standardized model systems and industry standards are needed to facilitate comparison across studies. The use of these in vitro models show promise in increasing our understanding of biology, disease progression, organ cross-talk, and many other areas as the field progresses.

Tissue Engineering Applications Bioprinting has been used in the laboratory to fabricate constructs targeting nearly every tissue types in the body. Although clinical implantation is still rare in this relatively new technology, there have been many successes in vitro and in vivo. Highly detailed, anatomically correct, and patient-specific tissue constructs have been fabricated for a number of tissues and organs. A wide range of cells has been shown to maintain viability, gene expression, and functional capabilities after the printing process. Various stem cells have demonstrated the ability to preserve their differentiation potential and also have been directed by various cues applied during the printing process [72,73]. This section will highlight a select few of the many tissue-specific regenerative medicine applications that have been studied with bioprinting technologies. Bone Bone regeneration is a natural target application for bioprinting given the importance of anatomical structure to its in vivo function. Conventional 3D printing technologies are in use clinically as patient-specific metal implants [74]. Bioprinting offers a unique and promising alternative to bone grafting because of the wide variety in anatomic location, defect size, and patient-specific morphology for bone pathologies [75,76]. The advantage of bioprinting is especially apparent for bone defects, which also feature a significant cosmetic function such as in craniofacial reconstruction [77]. Fedorovich et al. printed a mixture of Matrigel and alginate hydrogels with endothelial progenitor cells and mesenchymal stem cells (MSCs) [78]. The constructs were implanted subcutaneously in immunodeficient mice. The researchers were able to demonstrate that incorporating biphasic calcium phosphate microparticles caused the MSCs to differentiate into an osteogenic lineage and caused bone formation within 6 weeks after implantation. Phillippi et al. used inkjet bioprinting to pattern bone morphogenetic protein 2 on primary muscle-derived stem cells on fibrin-coated coverslips [79]. The stem cells differentiated into an osteogenic lineage even under myogenic differentiation media conditions. Keriquel et al. demonstrated in situ bioprinting by delivering nanoscale hydroxyapatite and osteoblasts into mouse calvarial defects, with positive outcomes [80]. Using CT scanning, Yao et al. were able to print anatomically accurate PCLehydroxyapatite mandible scaffolds that supported physiological loads [81]. Wang et al. examined the degradation profile of printed poly(propylene fumarate) scaffolds as it pertained to pore size, porosity, and mechanical properties [82]. They also developed a novel test for cytotoxicity of the degradation products and determined the scaffolds to be suitable for bone tissue engineering applications. To incorporate biological

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47. THREE-DIMENSIONAL TISSUE AND ORGAN PRINTING IN REGENERATIVE MEDICINE

materials with bioprinting technology, Pati et al. cultured MSCs onto a printed PCLePLGAeb-tricalcium phosphate (TCP) scaffold [83]. The cells deposited ECM during a brief culture period, after which the scaffold was decellularized. In vivo, the ECM-enriched scaffolds induced greater bone formation than did unadorned scaffolds of the same composition. Many limitations exist, including for large-sized defects and in higheload bearing applications. Lack of perfusion and neovascularization prevent large defects from being treated with bioprinting strategies; further research is needed in this area [84]. In addition, the discovery of new bioprinting-compatible materials and unique structural designs could increase maximum load-bearing applications for these constructs. More work is also needed to match the degradation profiles of scaffolding materials closely with those of the bone remodeling rate. New bone formation is obstructed if the scaffolding material degrades too slowly, but the defect site is left without a load-bearing material if degradation occurs too quickly, damaging nearby tissue. Altogether, bone tissue is one of the more promising target tissue applications for bioprinting owing to its many advantages relative to other tissue engineering strategies and the natural ability of bone to remodel in vivo (Table 47.2).

Cartilage Articular cartilage is imperative to reducing friction and absorbing compressive forces in load-bearing joints with little to no capacity for self-regeneration. Current cartilage tissue engineering strategies are insufficient for reproducing tissue that is equivalent to healthy cartilage [85]. However, there has been greater interest in the zonal differences found in cartilage matrix and cellular composition [86]. Bioprinting presents an appealing tool for constructing stratified scaffolds, especially in patient-specific size and shape of individual lesions [87]. Gruene et al. used laser-induced forward transfer to generate MCS grafts, showing good cell viability, density, and functionality [88]. MSCs were able to differentiate into osteoblasts and chondrocytes and the graft maintained good structural integrity in vitro. Cui et al. loaded chondrocytes into poly(ethylene glycol) dimethacrylate hydrogel and inkjet bioprinted them into an osteochondral plug [89]. The implant had mechanical and biochemical properties similar to native cartilage, and Safranin-O staining revealed good integration with surrounding cartilage tissue. The same group also used their experimental setup to investigate the effects of fibroblast growth factor 2 and transforming growth factor b1 on cartilage generation. Samples were cultured up to 4 weeks, and the highest GAG content was found for samples containing both growth factors, suggesting a synergistic effect between increased cell proliferation and increased chondrogenic phenotype expression [90]. To address some of the limitations of bioprinting, Xu et al. alternated between inkjet-bioprinted layers of rabbit elastic chondrocytes suspended in a fibrin-collagen hydrogel and electrospun PCL [22]. The construct formed cartilage-like tissues in vitro and in vivo demonstrated by type II collagen and GAG deposition. In addition, the scaffolds with electrospun layers showed improved mechanical properties compared with those that were only bioprinted. In particular, engineering the external ear has been a notably successful area of bioprinting cartilage tissue. The ear is almost completely avascular and aneural, it has a complex geometry, and it serves a largely aesthetic function, which places a greater emphasis on individualizing each prosthetic to the specific patient. Mannoor et al. developed a bionic ear that can translate sound waves into an electrical output [91]. The scaffold was extrusion bioprinted with sodium alginate, silver nanoparticles, and chondrocytes in an ear-shaped geometry around the conductive, sound-translating coil. Lee et al. extrusion printed PCL with PEG as a supporting sacrificial layer [92]. Chondrocytes and adipocytes were differentiated from adipose-derived stromal cells, encapsulated in alginate hydrogel, and dispensed into their respective regions. After 7 days in in vitro culture, immunostaining analysis confirmed chondrogenesis and adipogenesis. We also applied the extrusion-based bioprinting to fabricate a complex shape by making human-sized ear cartilage tissue construct (Fig. 47.5A) [47]. After implantation, the printed ear shape was well-maintained, with cartilage tissue formation upon gross examination. Histological analyses showed the tissue formation of cartilage tissue as confirmed by GAG and collagen type II staining. Quantitatively, GAG content increased over time, reaching 20% of that of native ear GAG content at 2 months after implantation. At this stage, the next challenge for bioprinting as a means for cartilage regeneration is to conduct translational studies. Few in vivo studies have been conducted. The long-term stability of bioprinted cartilage constructs has yet to be demonstrated and no studies have compared these strategies with practices used clinically. However, research in cartilage bioprinting is growing exponentially and exhibits many promising results for the future.

TABLE 47.2

Three-Dimensional Bioprinting Technologies for Tissue Regeneration Applications

Tissue/Organ Testing Model

Printing Method

Bone

In vitro viability and differentiation studies

Cell Type

Bioink

Outcomes

References

Extrusion

Endothelial progenitor and multipotent stromal cells

Matrigel and alginate hydrogels

Viability and differentiation capability were unaffected by printing process, and the two distinct cell populations were maintained within a single scaffold

[78]

In vitro differentiation studies

Inkjet

Primary muscle-derived stem cells

Fibrin

Incorporation of bone morphogenetic protein-2 caused spatially controlled osteogenic lineage differentiation even in myogenic media conditions

[79]

In situ bioprinting

Laser

Osteoblasts

Glycerol and n-Ha slurry

Successful in situ bioprinting into mouse calvarial defects with minimal side effects

[80]

In vitro

Extrusion

e

Polycaprolactonehydroxyapatite

Reconstructed from CT scans, anatomically accurate and supportive of physiological loads

[81]

In vitro degradation, mechanical, and cytotoxicity

Laser

Fibroblasts

Poly(propylene fumarate)

Scaffolds maintained their mechanical stability, and degradation products did not induce significant cell death

[82]

In vitro and in vivo bone formation

Extrusion

Mesenchymal stromal cells

PCL/PLGA/B-tricalcium phosphate

Scaffolds which were decellularized after brief culture period induced greater bone formation in vivo

[83]

In vitro

Extrusion

Spheroids of HUVECs and cardiac cells

Type 1 collagen

Viable cells, fusion and beating at 70 h with early signs of vascularization

[98]

In vitro

Inkjet

Cardiomyocytes

Alginate

Viability in thickness as high as 1 cm and contraction was observed at macroscopic and microscopic levels

[99]

In vivo cardiac infarct patch

Laser

HUVECs and human MSCs

Polyester urethane urea

Increased function and vessel formation compared with cell-only treatment 8 weeks postinfarction

[100]

In vitro

Extrusion

Cardiac-derived cardiomyocyte progenitor cells

Alginate

Cells demonstrated viability, phenotypic cardiac expression, and ability to migrate from hydrogel

[101]

Cardiac valve In vitro

Extrusion

Porcine aortic valve interstitial cells

PEGDA and alginate

Anatomical accuracy was confirmed; a range in mechanical properties was obtainable by varying the concentrations of the hydrogels

[6,103,104]

Cartilage

In vitro

Laser

MSCs

None

Good structural integrity and osteoblast and chondrogenic differentiation

[88]

In vitro

Inkjet

Articular chondrocytes

PEGDMA

Similar mechanical and biochemical properties of native cartilage and good integration with surrounding tissue; fibroblast growth factor 2 and transforming growth factor b-1 synergistically improved GAG deposition

[89,90]

Cardiac Muscle

Continued

TABLE 47.2

Three-Dimensional Bioprinting Technologies for Tissue Regeneration Applicationsdcont’d Printing Method

Cell Type

Bioink

Outcomes

References

In vitro and in vivo

Inkjet

Rabbit elastic chondrocytes

Fibrinecollagen

Combined PCL electrospinning and bioprinting technique facilitated type 2 collagen and GAG deposition with improved mechanical properties

[22]

In vitro

Extrusion

Chondrocytes

Alginate

Biomimetic ear could translate sound waves into an electrical signal and coexist with viable chondrocytes

[91]

In vitro

Extrusion

Chondrocytes and adipocytes

PCL and PEG (sacrificial) and alginate (encapsulation)

Chondrogenesis and adipogenesis confirmed by immunostaining

[92]

In situ bioprinting

Extrusion

Amniotic fluid derived stem cells

Fibrinecollagen

Amniotic fluidederived stem cells outperformed both MSCs and acellular graft

[105]

In vitro and in vivo

Laser

Fibroblasts and keratinocytes

Collagen

Early indicators of stratum corneum formation and blood vessels after 11 days

[106,107]

In vivo

Inkjet

Fibroblasts, keratinocytes, and microvascular endothelial cells

Collagen

10% improvement of wound contraction compared to allogeneic skin substitute

[108]

In vitro and in vivo

Extrusion

MSCs and chondrocytes

Alginate

Distinct tissue regions were found after 21 days in culture and 6 weeks after subcutaneous implantation

[113]

In vitro

Extrusion

Osteoblasts and chondrocytes

Type 1 collagen and hyaluronic acid

Cells showed better proliferation, migration, and function on hydrogels made from their native ECM and performed well in 14-day coculture

[37]

Musclee tendon

In vitro

Extrusion

Myoblasts and 3T3 fibroblasts

PCL & Polyurethane

Cells were viable after a week in culture and scaffolds showed an appropriate trend in mechanical properties

[96]

Pancreas

In vivo

Extrusion

INS1E b/islets

Alginate and gelatin

Scaffolds were formed and embedded while maintaining cell viability and morphology

[123]

Adipose

In vitro

Laser

Adipose-derived stem cells

Alginate

Cells maintained viability, differentiation potential, and adipogenic gene expression after 10 days

[121]

Neural

In vivo

Extrusion

Bone marrow MSCs and Schwann cells

e

Grafts underperformed autograft controls, but provide a proof-of-concept for future work

[114]

In vitro

Inkjet

Retinal ganglion cells and glia

e

Good cell viability and growth properties of cells was found after printing

[115]

In vitro and in vivo

Laser

Neuronal, Schwann, and dorsal root ganglion cells

After 3 weeks, the nerve guide supported reinnervation across a 3-mm gap equal to that of an autograft

[116]

Tissue/Organ Testing Model

Skin

Bonee cartilage

PEG

CT, computed tomography; FGF-2, fibroblast growth factor-2; GAG, glycosaminoglycan; HUVECs, human umbilical vein endothelial cells; MSCs, mesenchymal stromal cells; n-Ha, nano-hydroxyapatite; PCL, polycaprolactone; PEG, poly(ethylene glycol); PEGDA, poly(ethylene glycol) diacrylate; PEGDMA, poly(ethylene glycol) dimethacrylate; PLGA, poly(lactide-co-glycolide).

THREE-DIMENSIONAL BIOPRINTING IN REGENERATIVE MEDICINE APPLICATIONS

845

FIGURE 47.5 (A) Bioprinted ear construct: The shape was well-maintained, with substantial cartilage formation upon gross examination.

Histological and immunohistochemical analyses showed the typical cartilage tissue formation [47]. (B) Bioprinted organized muscle construct: The retrieved muscle constructs showed well-organized muscle fiber structures, the presence of acetylcholine receptor clusters on the muscle fibers (myosin heavy chain (MHCþ) and a-bungarotoxine (a-BTXþ)), as well as nerve (neurofilament) contacts with a-BTXþ structures within the implants at 2 weeks after implantation [47]. (C) Bioprinted muscleetendon unit (MTU): fluorescently labeled dual-cell printed MTU constructs (green: DiO-labeled C2C12 cells; red: DiI-labeled National Institutes of Health/3T3 cells; yellow: interface region between green and red fluorescence) [96]. (D) Bioprinted tracheal construct: Ingrowth of ciliary respiratory epithelium from the normal region was observed in the lumen of the bioprinted construct. Typical morphologies of respiratory mucosa and pseudocolumnar ciliary epithelium with goblet cells were well-developed at 8 weeks after implantation [120]. DAPI, 4,6-diamidino-2-phenylindole; L, lumen.

Skeletal Muscle and Tendon The organized ultrastructure of skeletal muscle is required for muscle contraction and force generation [93]. Because 3D bioprinting mechanism enables control of the spatial organization of cell-laden bioinks, we were able to fabricate highly oriented, muscle-like bundles to engineer skeletal muscle construct (Fig. 47.5B) [47]. The printed, aligned cellular construct began stretching along the longitudinal axis at 3 days in culture, and the constructs underwent compaction from polymeric pillars, keeping the fibers taut during differentiation. The aligned muscle fiber-like structures were observed at 7 days in differentiation medium condition. This bioprinted skeletal muscle construct maintained the tissue organization, followed by tissue maturation and host nerve integration in rats. The results demonstrate that the 3D bioprinting is capable of producing promising the structural and functional characteristics of skeletal muscle constructs in vitro and in vivo. Tendon has a hierarchical architecture, and tenocytes are aligned along with a dense collagen fibrous structure [94]. To mimic these structural characteristics of the tendon, an electrohydrodynamic jetting printing was introduced to generate a tubular-shape, multilayered tendon construct with highly porous, oriented microscale PCL fibers [95]. The cultured human tenocytes on the bioprinted structure showed a highly cellular orientation, metabolism, and type I collagen expression. 3D bioprinting technologies are particularly useful for composite tissue constructs such as muscleetendon. We used our integrated tissue and organ printing system to print four different components to fabricate a single integrated muscleetendon unit (MTU) construct (Fig. 47.5C)

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47. THREE-DIMENSIONAL TISSUE AND ORGAN PRINTING IN REGENERATIVE MEDICINE

[96]. The printed MTU construct was composed of mechanically heterogeneous polymeric materials that were elastic (polyurethane [PU]) on the muscle side and relatively stiff (PCL) on the tendon side, in addition to having a tissue-specific distribution of cells with C2C12 myoblasts on the muscle side and National Institutes of Health/ 3T3 fibroblasts on the tendon side. Results showed that cells were printed with high cell viability and cellular orientation as well as increased musculotendinous junctional gene expression. It was demonstrated that 3D bioprinting technology enabled a 3D heterogeneous tissue construction with region-specific biological and biomechanical characteristics. Cardiac Tissue and Heart Valves The heart is a complex organ in both shape and tissue organization, both of which are difficult to replicate by other fabrication methods. The ability to control the distribution of different cell types and growth factors spatially make bioprinting an attractive option for cardiac engineering, although only proof-of-concept successes have been accomplished thus far [97]. Jakob et al. were able to use extrusion-based bioprinting to pattern spheroids of HUVECs and cardiac cells on collagen [98]. The cells proved viable after the process; they fused at 70 h into a beating tissue and showed early signs of forming vascularization. Using inkjet bioprinting, Xu et al. printed a half heart shape (with two connected ventricles) with cardiomyocytes encapsulated in an alginate hydrogel [99]. Cell viability was preserved in constructs as thick as 1 cm owing to the designed porosity within the structure, and contraction was observed in vitro at both the microscopic and macroscopic levels. Instead of targeting whole-heart reconstruction, Gaebel et al. developed a cardiac patch for regeneration after cardiac infarction [100]. Using laser-induced forward transfer, HUVECs and human MSCs were patterned on polyester urethane urea and transplanted into the infarct zone. After 8 weeks, increased vessel formation and function were found, compared with a control treatment of bioprinted cells alone. A patch by Gaetani et al. used human cardiac-derived cardiomyocyte progenitor cells; they were extrusion-bioprinted in a sodium alginate mesh pattern [101]. They demonstrated cell viability, phenotypic expression of cardiac lineage, and the ability to migrate from the alginate, which suggested that bioprinting can be used to define cardiac cell delivery. Patients with heart valve failure must receive a replacement valve that can be mechanical requiring a lifetime of anticoagulant treatment, or biological, which typically fails within 10e20 years [102]. Bioprinting has gained momentum as a potential heart valve fabrication strategy to mimic the complex geometry and nonhomogeneous material makeup, mechanical properties, and cell distributions that naturally occur in heart valves [102]. Hockaday et al. used a dual ionic and physical cross-linking hydrogel strategy by using a PEGDA sodium alginate composite [6]. The printing accuracy of the aortic valve root wall and trileaflets was confirmed via microCT scanning. By varying PEGDA and alginate concentrations, elastic moduli were found to range from 1.5 to 5.3 kPa. Porcine aortic valve interstitial cells were seeded and cultured on the scaffold for 21 days with nearly 100% viability. Later studies by the same group printed the cells directly within the hydrogel, as opposed to seeding the scaffolds afterward, also with good geometric accuracy, cell viability, and mechanical properties [103,104]. Although these studies are far removed from use in the clinic, they demonstrate that bioprinting technology is amenable to cardiac tissue regeneration and open the door for many future studies focused on improving the current methodology and outcomes. Skin Bioprinting is an excellent technology for depositing distinct layers. It has been used in an attempt to mirror the layers of native skin, and research in this area has increased significantly. Skardal et al. performed direct in situ printing of amniotic fluidederived stem cells suspended in the fibrinecollagen hydrogel [105]. Compared with an acellular graft and an MSC graft, amniotic fluidederived stem cells showed increased microvessel density and capillary diameter. Laser-assisted bioprinting has been used to embed fibroblasts and keratinocytes in collagen [106]. Histology revealed a high density of both cell types and the expression of laminin protein. The same group grafted their construct onto mice and reported early indicators of stratum corneum formation and blood vessels after 11 days [107]. Yanez et al. printed keratinocytes and fibroblasts in collagen as well, but they included human microvascular ECs [108]. When implanted onto the backs of mice and compared with allogeneic skin substitute as a control, wound contraction improved by 10% and histological results appeared similar to those of normal skin. Sweat glands and hair follicles remain elusive, as does commercial and regulatory viability [109]. Nonetheless, skin bioprinting has shown many encouraging successes, and the clinical bioprinting of skin appears to be an impending reality [20].

CONCLUSIONS AND FUTURE PERSPECTIVES

847

Other Tissue Types Many other tissue types have been targeted with bioprinting technology, albeit to a lesser extent than those discussed thus far. This could be caused by a lesser clinical need, a higher difficulty of tissue engineering in general, or a poor match between the advantages of bioprinting and the necessary components for regenerating that tissue. Composite tissues are a major challenge facing regenerative medicine. No organ in the body is completely isolated, and many tissues such as tendons have specific and functional interfaces with other tissue types. Bioprinting is uniquely positioned to address this problem by spatially directing the placement of different cell types, growth factors, and biomaterials [110e112]. Fedorovich et al. extrusion bioprinted MSCs with hydroxyapatite, b-TCP, and biphasic calcium phosphate particles in alginate for one section of the scaffold and chondrocytes in alginate for the other [113]. Distinct tissue formation was found after 21 days in a mixture of osteogenic and chondrogenic media culture as well as after 6 weeks of subcutaneous incubation in vivo. Park et al. bioprinted osteoblasts in collagen I hydrogel and chondrocytes in HA hydrogel with good results after 14 days in vitro, in the process showing that the cells performed better on hydrogels made from their native ECM [37]. Finally, Merceron et al. targeted the MTU using PCL and 3T3 fibroblasts for the tendon zone and PU and myoblasts for the muscle zone [96]. Cells were viable after 7 days, and the scaffold showed appropriate trends in mechanical properties. Neural tissue has also been addressed by bioprinting. Owens et al. extrusion bioprinted a nerve graft containing bone marrow MSCs and Schwann cells [114]. The grafts were implanted for 10 months in a rat sciatic nerve injury model with autograft controls. The researchers concluded that that bioprinting was a promising approach to nerve grafting. Retinal ganglion cells and glia were piezoelectric inkjet bioprinted by Lorber et al., and the results showed good cell viability and growth-promoting properties in vitro [115]. Pateman et al. used a microsterolithographic technique to print PEG-based nerve guides for nerve repair [116]. In a 3-week, common fibular nerve injury mouse model, the nerve guide was capable of supporting reinnervation across a 3-mm injury with results similar to that of an autograft. Trachea is mainly composed of tightly stacked cartilage rings and respiratory mucosa in the luminal surface. Several synthetic implants have been used to reconstruct tissue defects [117e119]; however, these implants have been limited in their ability to mimic the tracheal functions biologically and biomechanically. We developed a biomimetic tracheal construct using a 3D bioprinting approach that could reconstruct a partial tracheal defect in a rabbit model (Fig. 47.5D) [120]. The printed tracheal PU constructs provided excellent structural characteristics. In the rabbit tracheal defect model, the printed PU constructs maintained the biomechanical function of the trachea, whereas the microscale porous architecture in the construct allowed cellular infiltration for the biological integration with host tracheal tissue. Moreover, the printed PU scaffold provided a proper microenvironment to facilitate the resurfacing of the ciliated respiratory epithelium and the ingrowth of connective tissue with microvasculature. Adipose tissue is not often targeted by regenerative medicine strategies, but Gruene et al. [121] laser bioprinted adipose-derived stem cells encapsulated in alginate. They proved that the cells maintained their viability, differentiation ability, and adipogenic gene expression after 10 days in vitro culture. Preliminary progress has also been made in several more complicated organs such as the intestine [122] and pancreas [123]. This section briefly examined the application of bioprinting to specific tissue types. Many studies were excluded because of space constraints; several tissue types that have been explored via bioprinting have not been covered here.

CONCLUSIONS AND FUTURE PERSPECTIVES The ultimate goal of tissue engineering and regenerative medicine is to reduce patient morbidity and mortality while improving quality of life by producing patient-tailored tissue constructs that induce tissue regeneration. 3D bioprinting technologies hold great promise for achieving this goal. The focus on replicating complex and heterogeneous tissue constructs continues to increase as 3D bioprinting technologies progress. Progression from single, simple tissues such as skin, bone, and cartilage, to organized contractile tissues such as skeletal muscle and cardiac tissue, to composite tissues such as osteochondral tissue and MTUs, and finally to robust organs such as the kidney and heart are under way. A novel bioink system needs to be developed to improve printability with high-resolution capability and structural maintenance. The availability is limited of biomaterials, including hydrogels and polymers, in 3D bioprinting that can serve as cell delivery bioinks and supporting structures, but which also provide biological properties and mechanical and structural support. Advances in biomaterials depending on 3D bioprinting mechanisms

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47. THREE-DIMENSIONAL TISSUE AND ORGAN PRINTING IN REGENERATIVE MEDICINE

are necessary for long-term success in tissue engineering applications. An approach that uses the decellularized ECM can provide a tissue-specific microenvironment to the cells. Decellularized ECM-based bioinks are still the closest biological microenvironment that mimics in vivo conditions; thus, tissue-specific ECM-based bioinks are capable of providing critical cues for targeted cell engraftment, survival, and tissue formation. There must be an increase in knowledge about biological, anatomical, and physiological aspects of complex tissues and organs. In particular, the development of in vitro 3D tissue models to study tissue-specific functions in the body will require a better understanding of morphological, structural, and functional units in tissues or organs. Moreover, the well-known limitation to building a large-scale tissue construct is vascularization in the construct. 3D bioprinting may have a unique capability among the various tissue engineering technologies to overcome this limitation. A few groups have made progress toward printing vascular structures; however, integrating functional microvascular structures into tissue or organ-like constructs has not been accomplished. Approaches to using high porosity, angiogenic factors, and highly organized patterns of vascular cells may encourage vasculogenesis. Bioprinting methods are able to construct 3D free-form shapes containing multiple cell types, biomaterials, and bioactive molecules, resulting in sophisticated tissue constructs that have the potential to replace damaged or diseased human tissues and organs. Although there is much work to be accomplished to advance these technologies toward successful clinical translation, our efforts will constantly contribute to producing clinically applicable tissue constructs until 3D bioprinting is able to improve the lives of patients.

List of Abbreviations 3D Three-dimensional b-TCP b-Tricalcium phosphate CAD Computer-aided design CAM Computer-aided manufacturing CT Computed tomography DNA Deoxyribonucleic acid ECM Extracellular matrix GAG Glycosaminoglycan GelMA Gelatin methacrylate HA Hyaluronic acid HUVECs Human umbilical vein endothelial cells MSCs Mesenchymal stem cells PCL Poly(ε-caprolactone) PEG Poly(ethylene glycol) PEGDA Poly(ethylene glycol) diacrylate PLGA Poly(D,L-lactic-co-glycolic acid) UV Ultraviolet light

Glossary Bioprinting The incorporation of biological materials into additive manufacturing techniques, either by directly depositing cells layer by layer or indirectly by three-dimensional printing biologically active materials for later use in cellular applications. Cross-linking A chemical bond between two polymer chains that changes the overall properties of the material. Electrospinning A method for producing fibers that uses electrical forces to draw out nanoscale threads of melted polymer material. Extracellular matrix The environment secreted by cells that biochemically and structurally supports a cellular network. Extrusion bioprinting Direct contact bioprinting mechanism that relies on pressure or displacement to force material through the syringes. Fused deposition modeling An additive manufacturing technology that extrudes heated material layer by layer to create three-dimensional structures. High-throughput screening Drug delivery process in which a lot of drugs or chemicals can be tested at a rapid pace. Hybrid bioprinting Using multiple bioprinting mechanisms in one system to overcome the limitations of each mechanism. Hydrogel A polymeric gel material in which the main component is water. In situ bioprinting Bioprinting directly in vivo such as onto a skin wound or into a bone defect, as opposed to bioprinting separately and then surgically placing a scaffold into a defect. Jetting bioprinting Originating from inkjet printers, this noncontact bioprinting mechanism uses pressure pulses to apply bioink in predetermined locations. Laser-assisted bioprinting Bioprinting mechanism that uses a focused laser to generate high-pressure bubbles that propel cell-containing material onto a substrate. Microfluidic device A device that is able to manipulate and control the flow of fluids on a microliter to picoliter scale. Scaffold The material which acts in place of the extracellular matrix , providing a physical, three-dimensional environment for cells to attach, migrate, and proliferate.

REFERENCES

849

Micropatterning Precisely controlling the cellular microenvironment on a substrate for the purposes of studying cell behavior. Perfusion The process of oxygen and other vital nutrients being delivered from the bloodstream to tissues and cells. Piezoelectric A ceramic crystal that creates an electric charge in response to an applied mechanical stress. Printability The ability and usefulness of a particular material to be applied as a bioink. Printing resolution The smallest dimension that can be controlled by a particular bioprinting system. Spheroid A three-dimensional conglomerate of cells, often organized into a spherical shape. Stereolithography Three-dimensional printing process in which liquid photopolymer is exposed above a perforated platform and then crosslinked by a UV laser forming the first layer. The platform then lowers exposing a new surface of liquid which the UV laser cross-links to form layer two. Structural stability The ability of a printed construct to maintain its shape.

References [1] Lee M, Wu BM, Dunn JC. Effect of scaffold architecture and pore size on smooth muscle cell growth. J Biomed Mater Res A Dec 15 2008;87(4): 1010e6. [2] Tsang VL, Bhatia SN. Three-dimensional tissue fabrication. Adv Drug Deliv Rev September 22, 2004;56(11):1635e47. [3] Xue W, Krishna BV, Bandyopadhyay A, Bose S. Processing and biocompatibility evaluation of laser processed porous titanium. Acta Biomater November 2007;3(6):1007e18. [4] Derby B. Printing and prototyping of tissues and scaffolds. Science November 16, 2012;338(6109):921e6. [5] Ozbolat IT, Hospodiuk M. Current advances and future perspectives in extrusion-based bioprinting. Biomaterials January 2016;76:321e43. [6] Hockaday LA, Kang KH, Colangelo NW, et al. Rapid 3D printing of anatomically accurate and mechanically heterogeneous aortic valve hydrogel scaffolds. Biofabrication September 2012;4(3):035005. [7] Landers R, Mu¨lhaupt R. 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C H A P T E R

48 Biomineralization and Bone Regeneration Kunal J. Rambhia, Peter X. Ma The University of Michigan, Ann Arbor, MI, United States

DEVELOPMENT AND FRACTURE OF BONE Bone Development Normal bone formation takes place by two different processes: endochondral ossification and intramembranous ossification. The endochondral ossification process features a sequential formation and degradation of cartilaginous tissue. By contrast, intramembranous ossification occurs by direct differentiation of precursor cells into mature osteoblasts [1]. Limbs of the body undergo endochondral ossification whereas the flat bones in the skull are formed by intramembranous ossification. Although these two processes are regulated by many of the same signaling molecules, the means by which cells differentiate and form bone are vastly different. In intramembranous ossification, precursor cranial neural crest cells differentiate into osteoblasts, which are bone-forming cells [2,3]. By contrast, mesenchymal stem cells (MSCs) establish a cartilage template, which is subsequently mineralized to form cortical bone during endochondral ossification [4]. During embryonic development, specific embryologic zones are defined for the formation of precisely defined structures of cartilage and bone. These early processes are regulated by several signals including soluble growth and differentiation factors, as well as cellecell and celleextracellular matrix (ECM) interactions. During an initial commitment phase, cells that will form bone are committed in a defined time and space. After initial commitment, cells are then differentiated into the terminal and mature cell phenotypes that are needed to construct bone.

Fracture Healing In some ways, bone fracture healing is similar to the process of initial bone development, as described earlier [5]. It involves the establishment of an environment that drives the differentiation of precursor cells to repair and replenish the tissue at the site of the injury. Fracture healing has several distinct features. After a fracture occurs, there is an immediate inflammatory response that recruits activated macrophages and polymorphonuclear neutrophils to the site of injury. The macrophages release multiple factors that stimulate the formation of a hematoma. Granulation tissue fibroblasts then proliferate to form a blastema. Osteoprogenitor cells are then recruited from the periosteum, surrounding soft tissue, and bone marrow at the site of the fracture. These cells differentiate into chondrocytes and osteoblasts to repair the fracture. Several soluble growth and differentiation factors influence this process. Mitogenic growth factors stimulate proliferation of precursor cells. They include fibroblast growth factors, insulin-like growth factors, and platelet-derived growth factors. Differentiation factors such as bone morphogenetic proteins (BMPs) induce the differentiation of precursors into osteoblasts and chondrocytes. Typically, differentiation follows a proliferative phase, so BMPs are often found in higher concentrations later in the healing process.

Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00048-5

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PRINCIPLES OF BONE TISSUE ENGINEERING For fractures to be healed successfully, a sufficient number of precursor cells must be present at the site of injury and they must be directed toward osteogenic differentiation. Cells can be recruited in vivo or implanted directly. Expression of signaling factors can be upregulated or they, too, can be applied directly. The signals must be given in appropriate time course and quantity to stimulate cell growth and differentiation in a controlled manner. Although fractures can heal naturally without intervention, larger defects require clinical interventions to repair or regenerate bone. These are referred to as critical-size defects. The reference standard treatment of these defects is autologous bone grafting, a process by which donor bone tissue removed from the patient’s hip is shaped according to the site of injury and surgically implanted into the defect site. These grafts contain viable cells, signaling factors, and a matrix that can support healing. However, the success of these grafts is variable, the surgical intervention has associated risks, and donor site morbidity has been reported as a complication of this treatment. In this chapter, we provide an updated overview of the state of bone tissue engineering as a means for treating bone defects and disease [6]. Bone tissue engineering provides an alternative method to regenerate bone while eliminating some limitations associated with grafts. In tissue engineering, a precisely engineered scaffold can be combined with an appropriate osteoprogenitor cell type and relevant growth and differentiation factors [7]. Damaged tissue sites that reach critical sizes can have limited self-healing potential in terms of the availability of precursor cells and concentration of growth and differentiation factors, as well as by scarring or inflammation. Thus, delivery of cells and growth factors can boost the efficacy of bioengineered scaffolds for bone regeneration in these cases. Nucleotide and gene delivery [8] and immune modulation [9] are also being pursued in bone and bone regeneration research.

STEM CELLS IN BONE TISSUE ENGINEERING Stem cells are defined cell populations with the capacity for self-renewal and the potential for differentiation to multiple cell fates [10]. In bone tissue engineering, adult MSCs, which are often derived from bone marrow, are the most common cell type used to regenerate bone. Other cell types frequently used in the field include embryonic stem cells (ESCs), induced pluripotent stem cells (iPSCs), and adipose-derived stem cells (ADSCs).

Mesenchymal Stem Cells MSCs are a common cell source for bone regeneration therapies because they are easily obtained and amplified. They are a multipotent cell type and have low immunogenicity. In vivo, MSCs develop into cells that make up bone, cartilage, and fat. MSCs have been isolated from fetal blood, liver, and umbilical cord, but they are most frequently derived from bone marrow. In culture, they exhibit fibroblastic morphology. They easily adhere to plastic dishes and flasks and can be expanded in culture for several passages without loss of phenotype. Osteogenic differentiation of MSCs can be stimulated in vitro by an induction media supplemented with dexamethasone, ascorbic acid (vitamin C), and b-glycerophosphate. Exposure to osteogenic media causes upregulation of alkaline phosphatase (ALP) activity and expression of two transcription factors that direct osteogenic differentiation, runt-related transcription factor 2 (Runx2) and osterix in MSCs during early stages of differentiation. During later stages, expression of matrix proteins including osteocalcin, osteopontin, and bone sialoprotein can be observed. Deposition of calcium within the ECM is also observed as osteogenic differentiation progresses (Table 48.1).

Adipose-Derived Stem Cells Like bone marrow, white adipose tissue is mesodermally derived and contains endothelial cells, smooth muscle cells, and MSCs. The MSCs can be isolated from adipose tissue and expanded in culture. These cells (ADSCs) behave similarly in culture to MSCs obtained from bone marrow. They are multipotent and highly proliferative and can be obtained in large numbers with minimal invasiveness. ADSCs can be directed toward osteogenic differentiation with the same induction media used for MSCs. Given their similarities to bone marrowederived MSCs, ADSCs are a clinically relevant cell source with potential for use in bone tissue engineering. However, compared with MSCs from other sources, the efficiency of osteogenic differentiation of ADSCs is low in both in vitro and in vivo

STEM CELLS IN BONE TISSUE ENGINEERING

TABLE 48.1

855

Markers of Osteogenic Differentiation

Marker of Osteogenic Differentiation

Highest Expression During Stages of Osteogenic Differentiation

Alkaline phosphatase (ALP)

Early

The function of ALP is not well-described; however, it is thought to have a role in promoting calcification of tissue [133].

Runt-related transcription factor 2 (Runx2)

Early

Runx2 is a transcription factor that activates osteogenic genes and has a role in cell fate determination of mesenchymal stem cell to osteoprogenitor cells [134].

Osterix

Early

Osterix is a transcription factor that acts downstream of Runx2 to direct osteogenic differentiation [135]

Osteocalcin

Late

Osteocalcin is the most abundant noncollagenous protein in bone extracellular matrix. It serves as a metabolic regulator of bone [136].

Bone sialoprotein (BSP)

Late

Extracellular matrix protein of bone that regulates the microenvironment of osteoblasts, hypertrophic chondrocytes, and osteoclasts. Regulates bone formation, mineralization, and turnover [137].

Osteopontin

Late

Similar to BSP, osteopontin is an extracellular matrix protein that regulates mineralization and bone matrix quality [138].

Function

studies [11e13]. Therefore, many researchers have designed studies that aim to increase osteogenic differentiation of ADSCs by treating cells with osteogenic growth factors or introducing osteoconductive scaffolds [13].

Embryonic Stem Cells ESCs are derived from fertilized embryos, typically donated from embryos that are not used after in vitro fertilization. These cells are grown in a plastic dish on a layer of feeder cells that maintain ESCs in their undifferentiated, pluripotent state. To differentiate ESCs, the cells are allowed to clump, forming embryoid bodies that can then be subjected to distinct culture conditions and develop into all three primordial germ layers. Direct differentiation of ESCs or differentiation of ESCs to an intermediate MSC stage can be used in tissue engineering applications [14]. Culture of committed cells is conducted in monolayer and supplemented with lineage-specific media conditions. Osteogenic differentiation media used to differentiate ESCs are the same combination of exogenous factors used for MSC differentiation. Although ESCs have greater capacity for self-renewal and can result in more tissue types compared with MSCs, ethical issues and complex culture requirements complicate their clinical and research use. Nonetheless, human ESCs (hESC) or hESC-derived mesenchymal progenitor cells have been seeded onto an osteoconductive nanofibrous scaffold and subsequently treated with osteogenic factors. These culture conditions guide the osteogenic differentiation of hESC-derived cells in three dimension (3D) and demonstrate the potential role of hESCs in bone tissue engineering [14,15].

Induced Pluripotent Stem Cells As an alternative to the controversial use of ESCs, iPSCs were developed to have the same capacity for selfrenewal and pluripotency as ESCs. These cells are adult cells that have been reprogrammed to behave like ESCs. When iPSCs were initially described in 2006, researchers exposed mouse embryonic or adult fibroblasts to four factors under culture conditions used to maintain ESCs. The resulting cells had the same morphology, growth properties, and gene expression as ESCs and formed tumors of all three germ layers [16]. These cells were referred to as iPSCs. The innovation of iPSCs was revolutionary to the field of stem cell biology, because it opened new areas of research into cell fate determination. It also opened the door to the use of iPSCs in tissue engineering. Several studies combined iPSCs or iPSC-derived cells with tissue engineering constructs to regenerate bone. In tissue engineering, one of the most appealing characteristics of iPSCs is that they originate from a patient with a tissue defect. The ability

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48. BIOMINERALIZATION AND BONE REGENERATION

to isolate and proliferate a patient-specific cell source would alleviate the risks in tissue engineering of allogenic graft rejection or disease transmission. Several methods and several adult cell types are being used to generate iPSCs. The most common of these methods is exposure of cells to exogenous factors, genetic modification by viruses, and other methods of genetic alteration [17,18]. Researchers are pursuing methods that increase the efficiency of reprogramming and reduce the risk for tumor formation [17,19e22]. As methods improve to reprogram adult somatic cells, so does the variety of cell types that can be reprogrammed to iPSCs. In humans, dermal fibroblasts are among the most common cells used for iPSC generation. Other cell types that have been reprogrammed include dental and oral cells, cord blood cells, and peripheral blood cells [23]. Human and murine iPSCs show promise for osteogenic differentiation [24,25] and in bone tissue engineering [26e31]. Like other cell types, human iPSCs and cells derived from iPSCs can be seeded onto scaffolds and differentiated bone-forming osteoblasts.

SCAFFOLDS FOR BONE TISSUE ENGINEERING Scaffolding Design Criteria Scaffolds engineered for bone regeneration are designed to support cell adhesion, migration, proliferation, and differentiation [7,32e34]. Key features of a supportive architecture for bone tissue engineering scaffold include a controlled porous microstructure, high interconnectivity between pores, mechanical stability, a controlled degradation rate, and osteoconductive interaction with cells. The porous microstructure facilitates the ingrowth of cells and enhances the regeneration of tissue. Interconnectivity of pores allows for uniform cell seeding and nutrient exchange. Scaffolds are designed to degrade at controlled rates to match the rate of formation of new tissue. Finally, the structure and microenvironment of scaffolds are conducive to differentiation of precursor cells into osteoblasts and to support functional activities of osteoblasts in bone regeneration. In addition, it was found that a fibrous nanostructure could enhance osteoconductive characteristics of scaffolds for bone tissue engineering [35]. Nanofibers mimic the native ECM of bone and nanofibrous scaffolds support enhanced osteogenic differentiation compared with non-nanofibrous scaffolds. Nanofibrous scaffolds were shown to enhance osteogenic differentiation in vitro and heal critical size calvarial defects better in a rat model compared with solid-walled scaffolds [36,37]. To fabricate porous 3D scaffolds and microspheres for bone regeneration, a variety of processing techniques have been developed. These include solvent casting/particulate leaching [38,39], gas foaming [40,41], emulsion freeze-drying [42], electrospinning [43,44], rapid prototyping [45,46], and thermally induced phase separation (TIPS) [47e49]. Each of these methods has advantages and disadvantages that are discussed at length in several review articles [7,35,50e52]. This chapter will not provide an exhaustive description of processing techniques. Instead, it will focus on how scaffolding design can be used to improve bone tissue engineering with respect to porosity, interconnectivity, mechanical strength, elastic modulus, morphology, and surface properties.

Porous and Highly Interconnected Scaffolds The porosity and interconnectivity of pores in scaffolds are important for uniform distribution of cells when seeded on a scaffold. A classical method to obtain porous scaffolds is the solvent casting/salt leaching method [38]. The process involves casting a mixture of polymer solution and salt (NaCl) into a mold. Subsequent drying of the mixture and leaching of the salt with water result in a porous structure. The pore size and porosity of the scaffolds fabricated using this method can be controlled by the particle size of the added salt and the salt-to-polymer ratio. This method produces scaffolds with limited interpore connectivity, which limits its value in tissue engineering. To create a scaffold with highly interconnected pores and a spherical pore shape, paraffin spheres and sugar spheres were used a porogen [53,54]. The interconnectivity of the scaffold was finely tuned by changing the heat treatment time to bond the paraffin spheres. Meanwhile, changing the concentration of polymer solution, size of paraffin spheres, and number of casting steps controlled the porosity and pore size. Together, this approach allowed for the creation of porous scaffolds with high interconnectivity, which is critical to the uniformity of cell seeding and tissue ingrowth on the scaffold.

SCAFFOLDS FOR BONE TISSUE ENGINEERING

857

Nanofibrous Scaffolds for Bone Tissue Engineering It is well-established that the ECM has a critical role in regulating cell behavior [55]. As previously noted, one of the key design features in bone tissue engineering has been the development of scaffolds with fibrous nanostructures that mimic native ECM and support cellular attachment, differentiation, and proliferation. Therefore, scaffolds were designed with nanofibrous structures that resembled the collagen fibers present in native ECM [56]. Collagen nanofibers were shown to have an integral role in cell adhesion, proliferation, and differentiation as early as the 1980s. To create poly-L-lactic acid (PLLA) scaffolds with nanofibers that resembled collagen fibers, a liquideliquid phase separation technique was developed by integrating sugar sphere template leaching with phase separation [57]. The resulting nanofibrous PLLA (NF-PLLA) matrix featured fibers with diameters ranging from 50 to 500 nm, or within the same range of collagen fibers [49,53,58,59]. Fig. 48.1. The combination of techniques to make porous materials and techniques to make nanofibrous materials produced scaffolds with biomimetic nanostructures and porous microstructures, which incorporate the advantages of both synthetic materials and natural structures for bone tissue engineering. Nanofibrous materials have been successfully fabricated using electrospinning for applications in bone tissue engineering. Electrospinning uses an electric field to overcome the surface tension of polymer solutions. This causes the polymer to be ejected out of a needle to a conductive collector, resulting in fibers on the nanoscale to submicroscale [35]. Early electrospun scaffolds for bone tissue engineering were made of poly-ε-caprolactone (PCL) as the biodegradable polymer of choice [60e62]. Today, electrospun scaffolds are made from a variety of synthetic and natural polymers and can also be combined into composite scaffolds, as described in the next section. Several different electrospun scaffolds for bone tissue engineering have been described, including one method that used PLLA nanofibrous scaffolds that were produced by both thermally induced phase separation and electrospinning, and enhanced by surface calcium phosphate deposition for bone tissue engineering [63,64].

(A)

(B)

(C)

(D)

FIGURE 48.1 (A, B) Scanning electron micrographs of scaffolds fabricated using a spherical porogen (paraffin sphere) leaching technique: (A) poly-L-lactic acid (PLLA), (B) poly(lactic-co-glycolic acid) (85/15), and (C, D) PLLA scaffolds fabricated using a thermally induced solid  liquid phase separation technique in dioxane (C) or benzene (D). (A, B) From Ma PX, Choi J-W. Biodegradable polymer scaffolds with well-defined interconnected spherical pore network. Tissue Eng 2001;7(1):23e33; (C, D) From Ma PX, Zhang R. Microtubular architecture of biodegradable polymer scaffolds. J Biomed Mater Res 2001;56(4):469e77; Ma PX, Zhang R, Xiao G, Franceschi R. Engineering new bone tissue in vitro on highly porous poly(alphahydroxyl acids)/hydroxyapatite composite scaffolds. J Biomed Mater Res (United States) February 2001;54(2):284e93.

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48. BIOMINERALIZATION AND BONE REGENERATION

Composite Materials for Bone Tissue Engineering Scaffolds The makeup of natural bone ECM is a combination of fibrous proteins, primarily collagen, upon which mineralized calcium is deposited [55]. In bone regeneration, researchers often attempt to replicate the natural ECM by synthesizing polymers from natural or synthetic building blocks. Because of intrinsic limitations of individual polymers, it is challenging to create a tissue engineering scaffold using a single material source that incorporates the fibrous microstructure of native ECM and supports sufficient mineralization. There are key limitations to each individual material, as no “perfect” single material exists for bone tissue engineering. Researchers therefore combine materials in various ways to form composite scaffolds that possess the advantageous qualities of each component material and better support the physical and biochemical requirements of bone regeneration. Integration of a mineral component into a fibrous scaffold can contribute to both the structural integrity and osteoconductive activity of the scaffold [35]. A variety of synthetic materials have been developed for bone tissue engineering, including PLLA, poly(lactic-coglycolic acid) (PLGA), and poly(glycolic acid). Similarly, collagen, hydroxyapatite (HAP), decalcified bone matrix, chitosan, and other natural materials have been used in bone regeneration applications. The goal of composite scaffolds is to have a biocompatible, osteoconductive matrix with suitable mechanical properties to support bone regeneration. One early composite matrix was made from combining synthetic polymer with ceramic through thermally induced phase separation. These scaffolds, made from highly porous poly(a-hydroxy acids)/HAP, had significantly improved physical properties (compressive modulus and yield strength) compared with pure polymer scaffolds. They also improved cellular ingrowth, increased osteoblast survival rate, and enhanced bone-specific gene expression [59,65,66]. Alternatively, coating a polymeric scaffold with apatite particles can produce a composite scaffold with the improved ability to deposit calcium in bone regeneration [66,67]. Other composites have been developed in an effort to combine the advantageous qualities of different synthetic and natural materials [68e70].

Injectable Scaffolds Concurrent with the development of implantable 3D scaffolds for tissue engineering, researchers have sought to develop injectable scaffolds in the form of hydrogels, microspheres, or nanospheres. Microspheres An effort to fabricate microspheres with the same porosity and interconnectivity of 3D scaffolds yielded two new tissue engineering scaffolds. These particles retain the favorable characteristics of their implantable counterparts while having the potential for noninvasive treatment. The nanofibrous hollow microsphere (single pore) and the nanofibrous spongy microsphere (multiple pores) (NF-SMS) were created to mimic the physical characteristics of a 3D scaffold in an injectable format. Nanofibrous microspheres were made from linear PLLA create solid nanofibrous microspheres. By substituting linear PLLA with star-shaped PLLA in a surfactant-free emulsification technique, a hollow microsphere, with a single pore was developed [71]. By changing the number of arms on the star-shaped PLLA and altering its hydroxyl density, additional pores could be introduced, forming the NF-SMS [72]. These two novel microsphere scaffolds have shown promise for cartilage and dental tissue regeneration [71,73]. Other microsphere-based systems are used as cell carriers and/or drug carriers and are composed of synthetic, natural, or composite polymers [74e78]. These can be porous or nonporous and are used independently as tissue engineering platforms, or in combination with a 3D scaffold or hydrogel [79,80]. Hydrogels Considerable research has been conducted into using hydrogels for bone regeneration. Hydrogels are polymer networks made of natural or synthetic components that are hydrophilic in nature and form viscous gels owing to 3D interactions between the polymer chains. These interactions can be either physical or chemical cross-linked between polymer chains. In bone tissue engineering, a number of hydrogels have served as a scaffold or used to deliver osteogenic drugs. An ideal hydrogel should be injectable so that it can be used to treat a bone defect noninvasively and fill abnormally shaped tissue defects. It should also be biocompatible and able to release osteogenic drugs over an extended time [81]. Controlled drug release remains a challenge for hydrogels because they are often prone to a burst release of drug within the first 24e48 h of treatment.

GROWTH AND DIFFERENTIATION FACTORS IN BONE TISSUE ENGINEERING

859

Hydrogels are often combined with other materials to form composite scaffolds for bone regeneration. PLLA/ PCL nanoyarns were suspended in a collagen hydrogel to make a composite scaffold that showed benefits for in vitro differentiation of human MSCs [82]. In another study, 45S5 Bioglass was introduced into alginate, forming a composite hydrogel that demonstrated enhanced osteogenic potential with MC3T3-E1 osteoblast precursor cells [83]. A third study used a nanohydroxyapatite-reinforced chitosan hydrogel to stimulate osteogenic differentiation in vitro and bone regeneration in vivo [84]. These represent a small sampling of the many composite hydrogels being evaluated for bone tissue engineering.

Surface Modification and Functionalization of Scaffolds for Bone Regeneration In an effort to increase the effectiveness of tissue engineering scaffolds, biologically active molecules can be attached to the surface or integrated within the scaffold. A specific molecule or peptide can be chosen to promote cell attachment and proliferation or direct cell differentiation [57,85,86]. For example, modification of poly(a-hydroxy acids) scaffold with gelatin enhanced osteoblast attachment, proliferation, and deposition of collagen [87]. Surface modification of a PLLA scaffold with calcium phosphate by electrodeposition enhanced osteogenic differentiation of MC3T3-E1 cells compared with cells attached to a control PLLA scaffold [64]. PEG hydrogels that are functionalized with integrin-specific peptides can enhance bone formation and vascularization when also treated with vascular endothelial growth factor [88]. Scaffolds can also be functionalized by the deposition of diamond nanoparticles. The addition of nanoparticles to a conventional scaffold can change the physical and chemical characteristics of the scaffold and enhance cell attachment and differentiation to support better bone formation [89]. Various additional methods and strategies for surface modification of scaffolds are being pursued [85,90].

Three-Dimensional Printed Scaffolds 3D printing technology is becoming increasingly capable of designing microenvironments and developing novel tissue engineering platforms. Many 3D printed scaffolds for bone regeneration have been developed that use many of the same natural, synthetic, and composite polymers discussed previously [91]. One approach to 3D printing scaffolds for bone tissue engineering is to print scaffolds that can be used to make vascularized bone grafts [92]. By varying the internal porosity of the scaffold or adding osteogenic and angiogenic factors such as zinc and silicon, researchers have attempted to increase the neovascularization of regenerated bone and simultaneously enhance the osteogenic differentiation of progenitor cells [93e95]. The successes and challenges of current strategies, descriptions of 3D printing methodologies, and clinical progress of 3D printing for bone tissue engineering are covered in depth in relevant reviews in the literature [96e98].

GROWTH AND DIFFERENTIATION FACTORS IN BONE TISSUE ENGINEERING Bioactive molecules, hormones, and nucleic acids can be used to enhance the growth and differentiation of cells. The effective combination of these factors with tissue engineering constructs and drug delivery systems can improve the regeneration of bone both in vitro and in vivo.

Bone Morphogenetic Proteins A group of proteins within the transforming growth factor-b superfamily, known as BMPs, have various roles in bone development, as well as the development of a multitude of tissues within the human body. Fifteen BMPs have been identified; of those, BMP-2 and BMP-7 show the most robust activity to induce bone formation. Both BMP-2 and BMP-7 have been used clinically in US Food and Drug Administration (FDA)-approved devices to treat a narrow range of severe bone defects. BMPs act on bone-forming osteoblasts, preosteoblasts, and other precursor cells, including MSCs, ADSCs, iPSCs, and ESCs. Typically, BMPs form homodimers and bind with known BMP receptors to activate a cascade of signaling pathways that initiate differentiation and mineralization of cells and tissue. Some evidence of BMP-2/7 heterodimers having a stronger osteogenic effect than the respective homodimers has been published [99,100]. BMP-2 and BMP-7 act to stimulate osteogenic differentiation through intracellular phosphorylation of the smad

860

48. BIOMINERALIZATION AND BONE REGENERATION

1/5/8 proteins. Upon phosphorylation, p-smad 1/5/8 proteins translocate to the nucleus and upregulates Runx2. BMPs can also activate other pathways in parallel, including Wnt signaling [101]. BMP-related signaling is generally divided into either smad-dependent or smad-independent, depending on which pathway is being described or studied. The clinical use of BMP-2 began after its approval by the FDA in 2002 for the treatment of anterior lumbar interbody fusion within a threaded titanium tapered cage. The scope of its use expanded in 2004 to include tibial nonunions, and in 2007 for oral maxillofacial reconstruction. After the expanded use of BMP-2 for these clinical and other off-label applications, a number of serious side effects were observed and reported [102]. Inflammatory complications, including swelling of the cervical spine [103], radiculopathy [104], ectopic bone formation [105,106], osteoclast activation [107,108], urogenital events [109], and wound complications were reported [102]. For in vitro bone regeneration, BMP-2 and BMP-7 are applied in concentrations ranging from 50 to 100 ng/mL. For in vivo applications, encapsulation of BMP-2 and BMP-7 in a drug delivery system can provide local, clinically relevant doses for short or long time. This is often accomplished by encapsulating BMPs in a delivery vehicle such as a biocompatible polymer microsphere or hydrogel. Controlled release of recombinant human BMP-7 in a microsphere delivery system enhanced a bone tissue engineering scaffold in a subcutaneous mouse model [110].

Parathyroid Hormone Delivery Parathyroid hormone (PTH) is secreted by the parathyroid glands and can have both catabolic and anabolic effects on bone [111]. It is well-known that a continuous and sustained treatment of PTH will lead to increased bone resorption or reduction in bone mass [112]. Conversely, an intermittent or pulsatile treatment of PTH can enhance osteogenesis or bone regeneration. PTH is used as an FDA-approved treatment of osteoporosis, which requires daily injections of a sufficient dose to be effective. A novel approach to PTH delivery uses a multilayered scaffold to achieve local pulsatile release of PTH to stimulate bone regeneration. Layers of PTH are separated by layers of polyanhydride (PA) and sealed with PCL. Surface erosion of PA allows the sequential release of PTH from the scaffold in controlled doses. Using this approach, a daily spike in PTH was observed and corresponded to superior healing of a critical-size defect in a calvarial mouse model [113] (Fig. 48.2).

Nucleotide Delivery and Gene Therapy Nucleic acids (DNA, RNA, small interfering RNA, and microRNA [miRNA]) regulate gene and protein expression in cells. There are several methods to enhance or suppress the expression of specific proteins or genes that are involved in healing or new tissue formation. The goal of gene therapy is to increase or sustain the local expression of factors related to healing, cell recruitment, cell proliferation, or cell differentiation. This is most frequently accomplished by delivering nonintegrating plasmid DNA or inserting additional copies of the desired genes into the chromosomal DNA. Transfer of genes to cells at the site of injury can be accomplished by viral and nonviral delivery methods [8,114]. Transcription and translation of these additional gene copies can enhance the expression of factors that promote healing. For bone regeneration, gene therapy often involves modulating the temporal and spatial expression of BMPs [115,116]. The use of adenovirus vectors allows researchers to transfect progenitor or differentiated cells with additional copies of genes that enhance regenerative processes. In one example, adenoviral gene transfer was used to elicit the expression of BMP-7 from muscle cells. These cells were then implanted into femoral defects to evaluate their regenerative potential [117]. Although gene transfer can be effective, improvements are needed in the efficiency, stability, and safety of these methods. In a nonviral approach, a gene vector that incorporates a plasmid encoding for BMP-2 is protected by a poly(D,Llactide) (PDLLA) copolymer and delivered to a mandibular defect [118]. In this strategy, the gene vector is delivered on a PDLLA scaffold to induce cells localized at the site of injury to express BMP-2 and improve regeneration of the bone defect. In a similar approach, a BMP-7 plasmid enhanced by a heparin-binding site is transfected into MSCs and applied to regenerate bone defects [119]. Local delivery of nucleic acid molecules is another promising novel approach to bone regeneration. In particular, applications that use miRNA are being developed to enhance pro-osteogenic factors or suppress antiosteogenic factors [120]. Controlled release of miRNA-26a demonstrated a significant regenerative capability in an in vivo osteoporotic mouse model. In this example, miRNA-26a increased bone regeneration and healing by targeting a known inhibitor of osteoblastic activity, glycogen synthase kinase-3b [121]. PolymeremiRNA complexes (polyplexes) were

IMMUNOMODULATION IN BONE REGENERATION

861

FIGURE 48.2 Experimental design of using a three-dimensional (3D) cell-free scaffold and a parathyroid hormone (PTH) deliver device (pulsatile or continuous) to repair calvarial bone defect in a mouse model. PA, polyanhydride.

created by mixing miRNA-26a and hyperbranched polyesters with affinity to miRNA-26a. These polyplexes were encapsulated in PLGA microspheres and were subsequently immobilized on a nanofibrous PLLA scaffold. The polyplexemicrosphereescaffold platform was surgically implanted in critical-size calvarial defects in osteoporotic mice. Efficient intracellular delivery of the miRNA was accomplished in a stepwise manner. As the PLGA microspheres were hydrolyzed, the miRNA-26aecontaining polyplexes were released from their interior. These polyplexes were subsequently engulfed by cells via endocytosis. Endosomal escape allowed release of the miRNA into the cytoplasm of the cell. Fig. 48.3 Intracellular release of miRNA-26a in osteoblasts increased bone mass and increased expression of early and late genetic markers of osteogenic differentiation [121]. The advent of Clustered Regularly Interspaced Short Palindromic Repeats (CRISPR)/CRISPR-9eassociated (Cas9) gene editing technology has increased the efficiency by which researchers can manipulate the genome of mammalian cells and establish new cell lines for use in culture. Although CRISPR/Cas9 has increased the ability of researchers to study bone function and regeneration [122e124], no CRISPR-based therapeutic approaches have been developed to date.

IMMUNOMODULATION IN BONE REGENERATION During the natural process of healing of a bone injury, the inflammatory and immune response activates several distinct cell types that can directly or indirectly interact with progenitor cells, osteoblasts, and osteoclasts [125e127]. Methods to modulate these interactions are being developed to enhance bone regeneration therapies.

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48. BIOMINERALIZATION AND BONE REGENERATION

FIGURE 48.3 Hyperbranched polymers (HPs) and microRNA (miRNA) formed polyplexes in water. The HPemiRNA polyplexes were encapsulated into the poly(lactic-co-glycolic acid) (PLGA microspheres) via the double-emulsion method, and the PLGA microspheres containing the HPemiRNA polyplexes were then attached on the poly-L-lactic acid (PLLA) nanofibrous scaffold. The PLGA microsphere-incorporated PLLA scaffolds were implanted into mice. The HPemiRNA polyplexes released from the PLGA microspheres (on the PLLA scaffold) could be taken into cells through endocytosis. Intracellular release of miRNA in the cytosol after enzymatic polymer degradation allows its regulation of gene expression.

Macrophages Macrophages are part of the innate immune response to injury. As they are recruited to the site of injury, macrophages act by phagocytizing bacteria and damaged tissue. They also release cytokines and growth factors that initiate and promote healing. One study demonstrated that coculture of MSCs with specific macrophage subtypes enhanced osteogenic differentiation of MSCs. Antiinflammatory macrophages (M2) promoted osteogenic differentiation of MSCs better than proinflammatory macrophages (M1) or immature macrophages (M0) [128]. In another study, M1 macrophages increased osteogenic differentiation of MSCs more dramatically than M2 macrophages [129]. These two studies are seemingly contradictory, although in both studies, coculture of MSCs with M1 and M2 macrophages enhanced osteogenic differentiation, but the importance of macrophage polarity [130] and the molecular signaling pathways involved have yet to be clearly elucidated. In practice, macrophages have been stimulated by the release of a macrophage-recruiting agent and platelet-rich plasma from a gelatin hydrogel scaffold to a defect site in the ulna bone in a rat model demonstrated enhanced bone healing [131]. Mechanistic studies that clarify the role of macrophages in bone healing are still limited.

T Cells The adaptive immune response also has a role in bone regeneration in the form of activated T cells. T lymphocytes regulate callus formation in bone healing by secretion of pro-osteogenic cytokines (tumor necrosis factor-a and interleukins [IL]-6 and -1b). Coculture of activated T cells with MSCs enhances osteogenic differentiation owing to the osteogenic effect of soluble factors released by CD4þ T lymphocytes [132]. Furthermore, there is a differential effect on osteoblast maturation from pro inflammatory and antiinflammatory CD4þ T-cell subsets. CD4þ T-helper 17e specific cytokines (IL-17A and IL-17F) increased ALP activity in MSCs, and IL-17A in particular works synergistically with BMP-2 to increase calcium deposition by MSCs.

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49 Hair Cell Regeneration in the Inner Ear and Lateral Line Matthew W. Kelley1, Jason R. Meyers2 1

National Institutes of Health, Bethesda, MD, United States; 2Colgate University, Hamilton, NY, United States

INTRODUCTION The ability to detect sounds and motion is mediated through mechanosensory hair cells located in the inner ears of all vertebrates. The basic cellular structures of hair cells are similar across vertebrate classes, but mammals are unique in their inability to regenerate hair cells that have been lost as a result of aging or trauma. As a result, mammals are the only vertebrates that incur lasting deficits in hearing or balance. In this chapter, we will describe the process of hair cell regeneration that occurs in nonmammalian species, explore the factors that prevent regeneration in mammal inner ears, and discuss efforts to induce regeneration using transgenic mouse models and gene therapy.

STRUCTURE OF THE INNER EAR The inner ear can be grossly divided into three structures: the membranous labyrinth, a series of epithelial-lined fluid filled chambers derived from the otocyst; the bony labyrinth, a dense bony structure that surrounds and protects the membranous labyrinth; and the eighth cranial nerve, also called the vestibulocochlear nerve, which provides afferent innervation to all sensory structures within the inner ear (Fig. 49.1). The membranous labyrinth can be further subdivided into a ventral auditory section that contains the cochlea and a dorsal vestibular section that contains the three semicircular canals that act to mediate balance, as well as two otolithic organs, the saccule and utricle, which have a role in perception of linear acceleration and gravitation forces. Each of these structures contains a sensory epithelium that is composed of mechanosensory hair cells and a population of surrounding epithelial cells referred to as supporting cells. Vestibular sensory epithelia appear as rounded or oblong patches containing thousands of mechanosensory hair cells. The overall structure of these sensory patches is grossly similar among all vertebrate classes. In contrast, auditory epithelia show a greater degree of diversity; some develop as narrower, elongated stripes of cells. The most extreme example of this is the mammalian auditory sensory epithelia, also called the organ of Corti, which features four or five rows of hair cells extending up to 70 mm [1], which are functionally and morphologically divided into one row of inner hair cells that detect auditory stimuli and three or four rows of outer hair cells that serve as a mechanical amplifier. Finally, aquatic vertebrates such as fishes and amphibians have an additional hair cell-based sensory system called the lateral line, which is composed of a series of canals and sensory patches arrayed along the outer surface of the head and body, and which enables the detection of flow and vibrations in the surrounding water. Mechanosensory hair cells are so named because of the presence of a modified bundle of microvilli, referred to as stereocilia, that project from the lumenal surface of each cell [2,3]. The morphology of the bundle is characterized by an asymmetric staircase pattern in which the lengths of the individual stereocilia increase toward one end of the bundle. Deflection of the bundle in the direction of the longest stereocilia leads to opening of transduction channels, the influx of positively charged ions, and subsequent increases in the rate of neurotransmitter release. Each hair cell Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00049-7

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Semicircular canals

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FIGURE 49.1 Overview of the inner ear. (A) Cut away view of the human ear. The pinna and external auditory meatus (EAM) transmit sound waves to the tympanic membrane located at the central end of the EAM. Vibration of the tympanic membrane is amplified through the middle ear ossicles (OC). The footplate of the stapes vibrates the footplate of the stapes, which creates a pressure wave that travels along the cochlear spiral. The three semicircular canals, along with the utricle and saccule, mediate sense of balance, acceleration, and gravity. Signals generated in the inner ear are carried to the central nervous system through separate branches of the eighth nerve. (B) View of the inner ear illustrating the three semicircular canals, cochlear spiral, and endolymphatic duct (dark gray). The branches of the eighth nerve are illustrated in yellow. (C) Schematic cross-sections of the sensory epithelia in the utricle (upper image) and cochlea (lower image). Hair cells are red and supporting cells are green. In the utricle, hair cells are arranged in a large patch whereas in the cochlea, hair cells are arranged in four or five rows that extend along the length of the cochlear spiral. (D) Schematic of a single utricular hair cell illustrating the staircase pattern of the stereocilia (orange) and the single kinocilium (gray).

forms synapses with peripheral axons from one or more afferent neurons that transmit changes in the rate of neurotransmitter release into brain stem nuclei and higher central nervous system structures, where these signals are interpreted as perceptions of sound, vibration, or movement.

HAIR CELL LOSS A number of factors have a role in the death of hair cells. Foremost among these is aging. According to the National Institute on Deafness and Other Communication Disorders, nearly 35% of all individuals will have a significant hearing loss by age 65 years, and this number will increase to over 50% by age 80 years. Similarly, virtually all individuals lose the ability to perceive frequencies above 17 kHz by age 25 years. The underlying cellular or genetic mechanisms that lead to this progressive hair cell loss remain largely unknown, although mutations in some genes, in particular Cdh23, have been shown to cause age-related hearing loss in mice [4]. Similar declines in hair cell numbers have also been reported in aged vestibular epithelia; however, a strict correlation between hair cell loss and loss of vestibular function has not been established [5].

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Environmental factors can significantly accelerate the rate of hair cell loss. Exposure to loud noises, even in the short term, can lead to hair cell death from both mechanical damage and neurotoxicity from calcium, resulting in hearing loss. Toxins such as copper and Gingko biloba have also been shown to be harmful to hair cells [6,7]. In addition, several medicines, most notably aminoglycoside antibiotics and the anticancer drug cisplatin, have been shown to kill hair cells [8]. As is the case for age-related hearing loss, the cellular basis for hair cell death after exposure to any of these stressors is not particularly well-understood, but metabolic stress and generation of reactive oxygen species, may be important mediators [9]. This is particularly true for outer hair cells in the organ of Corti, which are highly sensitive to all environmental stressors. Outer hair cells are motile cells that generate cellular contractions in response to stimulation. Because these contractions typically occur at the optimal frequency for each cell, some cells may need to contract repeatedly at rates as high as 20 kHz. This high level of activity may produce a chronic metabolic stress that makes the outer hair cells particularly sensitive to additional external challenges.

HISTORY OF HAIR CELL REGENERATION As discussed in the Introduction, in mammals, hair cells are primarily generated during the embryonic or early postnatal period, depending on the species. Therefore, in adults, hair cell loss is permanent. Before the 1980s, it was assumed that similar limitations existed in the inner ears of all other vertebrates. However, examinations of the inner ears of sharks and rays, which grow indeterminately, indicated an ongoing increase in the number of hair cells in the inner ear sensory patches of these animals. In several species of sharks, the increase in the number of hair cells is remarkable; more than 180,000 cells are added to just a single inner ear sensory patch, the macula neglecta, over the life of an individual animal [10]. Subsequent studies in amphibians used a mitotic tracer, tritiated-thymidine, to demonstrate that new hair cells in the inner ears of these animals were generated through cellular proliferation of surrounding cells [11]. These results motivated additional studies using chickens, homothermic vertebrates with determinant growth, and an elongated auditory structure, the basilar papilla (BP), which is functionally similar to the mammalian cochlea. After induction of hair cell damage using prolonged exposure to loud pure tones, initial experiments used scanning electron microscopy to image the BP [12]. Whereas BPs imaged immediately after noise exposure showed widespread hair cell loss and damage, analysis of similar BPs after a 2- to 4-week recovery period illustrated nearly complete recovery of the hair cell population. Moreover, recovery of auditory function tracked closely with morphological recovery, demonstrating that the new hair cells were functional and that the rest of the auditory system remained intact after exposure to noise. Finally, introduction of a mitotic tracer demonstrated that at least some of the regenerated hair cells arose from proliferation of the surrounding supporting cells [13,14]. To determine whether this regenerative ability is restricted to young birds, similar experiments were performed in aged quails [15]. Even quails that were near the end of their expected life span (3 years) were able to regenerate hair cells and recovery auditory function, which demonstrated that this ability is retained throughout the life of the animal. Subsequent studies at the cellular level identified two different mechanism for the formation of regenerated hair cells (Fig. 49.2). As discussed previously, some new hair cells arise from the re-entry and subsequent division of supporting cells, but in other cases, supporting cells are able to convert into hair cells directly, a process referred to as transdifferentiation [16,17]. These results demonstrate that in birds, supporting cells can act as hair cell progenitors. This suggests that these cells retain some stem or progenitor capacity throughout the life of the animal. Whether every supporting cell can act as a stem or progenitor cell or whether this ability is limited to a subset of cells within the epithelium has not been determined. Interestingly, the response of supporting cells, in terms of whether they transdifferentiate or reenter the cell cycle, is spatially segregated; cells located on one edge of the BP are more likely to transdifferentiate whereas those on the other edge are more likely to re-enter the cell cycle [18]. The molecular and/or functional bases for these differences have not been determined. A final consideration in the avian system is the unexpected observation that vestibular epithelia undergo constant turnover of hair cells; old cells are replaced with new ones such that the average life of a hair cell in a vestibular sensory epithelium is approximately 1 month [19]. This process may be a compromise of the ancestral trait of ongoing addition of hair cells, as seen in fish and amphibians, with the derived trait of determinant growth as occurs in birds and mammals. However, this also demonstrates that vestibular epithelia can efficiently maintain all of the neuronal connections required for normal function even as hair cells are being continually lost and regenerated.

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FIGURE 49.2 Modes of hair cell regeneration in the chick basilar papilla. (A) Cross-sectional view of an idealized basilar papilla. Hair cells (red) and supporting cells (green) are present in a pseudostratified epithelium. (B) In response to an insult, hair cells are ejected from the epithelium (small arrow), creating a gap in the mosaic of hair cells and supporting cells and eliciting a response from one or more of the remaining supporting cells (yellow). (C) Transdifferentiation. Under some circumstances one of the remaining supporting cells will directly transform into new hair cells (yellow). This results in a decrease in the number of supporting cells. (D) Proliferative regeneration. The second possible response is for a remaining supporting cell (yellow) to undergo mitotic proliferation to generate new progenitor cells. (E) These cells then go on to form new hair cells and supporting cells (yellow).

SPONTANEOUS HAIR CELL REGENERATION IN MAMMALIAN VESTIBULAR ORGANS The discovery of spontaneous hair cell regeneration in birds sparked a re-examination of the regenerative potential within adult mammalian vestibular and auditory epithelia. As a first step, work from Warchol et al. [20] and Forge et al. [21] demonstrated limited proliferation of supporting cells in cultured utricles from Guinea pigs and humans, as well as the presence of hair cells with bundles that appeared immature in Guinea pigs. Although these results were compelling, the number of potential new cells was small and a preponderance of clinical data clearly suggests that meaningful vestibular hair cell regeneration does not occur in adult humans. In contrast to the findings in the utricle, similar studies in the adult mammalian auditory system confirmed the absence of hair cell regeneration in this epithelium [22]. Despite the importance of the findings in the utricle, further examination of potential hair cell regeneration in this structure was slowed by difficulty in reliably killing hair cells in the vestibular system in vivo. In the auditory system, two different approaches have been developed to kill hair cells reliably and reproducibly: exposure to loud sounds and administration of aminoglycoside antibiotics such as neomycin or kanamycin, alone or combined with a loop-diuretic such as furosemide [23]. The effects of both of these treatments have been studied extensively and have led to precise protocols that yield consistent results. Moreover, because the auditory system is organized along a tonotopic gradient, damage can be mapped to particular regions of the auditory epithelium based on the results of tests for auditory sensitivity at different frequencies [24]. Because vestibular epithelia are not stimulated by auditory vibrations, exposure to loud sound does not cause hair cell loss, and whereas the sensitivity of vestibular hair cells to aminoglycosides is comparable to auditory hair cells in vitro, systemic treatments result in variable hair cell death. Moreover, many animals that exhibit significant vestibular loss of function after a chemical insult will show progressive and marked recovery over time as a result of functional compensation based on visual input [25]. As a result, between the late 1980s and the early 21st century, it was virtually impossible to assess the extent of hair cell regeneration accurately in vestibular epithelia in vivo because there was no way to kill the existing hair cells consistently. However, the situation changed with the development of several lines of transgenic mice beginning around 2010. First, lines were developed in which genes that are expressed specifically in supporting cells drive the expression of cre recombinase (cre). When one of these lines is combined with one of several reporter lines in which expression of cre leads to the permanent expression of a fluorescent molecule such as green fluorescent protein (GFP), it is possible to differentiate regenerated hair cells that will express GFP because they arise from supporting cells, from nonregenerated hair cells that are GFP-negative [26]. In addition, in 2013, Rubel and colleagues [27] developed a mouse model in which the promoter

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for the hair cellespecific gene Pou4f3 is used to drive expression of the human diphtheria toxin receptor (Dtr). Because the human Dtr is approximately 10,000 times more sensitive to diphtheria toxin [28] compared with the mouse Dtr, this line can be used to kill vestibular (and auditory) hair cells effectively and consistently by giving mice injections of diphtheria toxin. In fact, injection of diphtheria toxin in this line consistently kills approximately 94% of hair cells in the utricle by 14 days after treatment [27]. Animals killed at specific recovery times between 15 and 180 days showed a modest recovery of hair cells, up to approximately 17% of the original number, and the loss of supporting cells and lack of incorporation of mitotic markers demonstrated that these cells arise from the conversion of surrounding supporting cells into hair cells. In addition, a small percentage of regenerated hair cells arose as a result of mitotic division; the number of mitotically generated cells decreased rapidly as animals aged past the very early postnatal period [27,29]. These results provide definitive evidence that a limited amount of spontaneous hair cell regeneration can occur in adult mammalian vestibular epithelia. However, whether the regenerated hair cells lead to recovery of function remains to be determined.

ROAD BLOCKS TO REGENERATION An intriguing question which, if answered, might provide insights regarding the development of clinical strategies, is: Why do mammals have a greatly reduced capacity to regenerate hair cells, including a complete loss in the auditory system? At a systems level, one of the most appealing hypotheses is that the increased complexity of the organ of Corti relative to other auditory epithelia, and in particular the highly differentiated state of the supporting cells within the organ of Corti, has resulted in those cells losing the ability to de-differentiate, as might be required to change fate or re-enter the cell cycle. Although the basis for such a loss is not clear, one possibility would be alterations in the epigenetic landscape of the supporting cells, leading to an inability to reactivate important genes required for one or more aspects of regeneration [30]. There is ample evidence from other systems that is consistent with a negative correlation between cellular differentiation and plasticity, sometimes correlated with epigenetic changes, which supports this idea [31]. However, a noteworthy caveat is that this hypothesis does not explain why regenerative ability would also be greatly decreased in the vestibular system where supporting cells appear to be considerably less differentiated compared with similar cells in the organ of Corti and are comparable to those in nonmammalian vestibular organs. Still, it is possible that changes in the structural, transcriptional, or epigenetic state of auditory supporting cells might also lead to similar changes in other inner ear supporting cells. The reduced capacity for mammalian supporting cells to regenerate hair cells seems to be tied to their own maturation. Generally, loss of cochlear or vestibular hair cells during embryonic or perinatal periods results in supporting cell proliferation and hair cell regeneration [29,32], although the capacity for regeneration decreased rapidly with postnatal age. Corwin and colleagues provided intriguing, albeit correlative, data suggesting that a structural component of supporting cells, dense actin belts located just beneath the lumenal surfaces of supporting cells, may act to inhibit the ability of those cells to undergo a regenerative response [33,34]. At early postnatal time points, lumenal cortical actin in supporting cells creates a thin circumferential belt located close to the lateral cell membrane. As discussed, supporting cells in the utricle generate a significant regenerative response, including cellular proliferation, during this same period. However, as an animal ages, the width of supporting cell lumenal actin belts increases while regenerative ability decreases. Examination of actin belts in non-mammalian vertebrates, including birds and fish, revealed thin belts similar to those observed in newborn mammals, regardless of the age of the animal. Unfortunately, it has not yet been possible to disrupt these belts to demonstrate whether they actually prevent supporting cells from initiating a regenerative response. The ability of postnatal supporting cells to respond to growth factors and extracellular matrix components in culture also decreases rapidly [35,36]. Thus, the lack of mammalian regeneration appears to be a trait acquired by the maturation of supporting cells.

INSIGHTS FROM DEVELOPMENTAL BIOLOGY Because meaningful spontaneous regeneration does not occur in mammals, it seems likely that manipulation of the system, possibly through pharmacological and/or genetic approaches, will be necessary to induce a response. To develop an appropriate strategy, it will be necessary to identify the molecular and genetic pathways that regulate the key steps in a regenerative response. Based on the process that occurs during hair cell regeneration in nonmammalian vertebrates, hair cell regeneration can come from the nonmitotic conversion of supporting cells into

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hair cells or the proliferation of supporting cells followed by differentiation of some of the progeny into hair cells. Because neither proliferation of supporting cells nor differentiation into hair cells occurs at a high rate in adult mammalian hair cell epithelia, one possible way to identify the factors that regulate each of these events is to examine them during development. Significant progress has been made in understanding both proliferation and differentiation, but our understanding of the regulation of differentiation is more advanced and therefore will be discussed first. One of the initial findings was actually related to a genetic signal that inhibits rather than promotes hair cell formation (Fig. 49.3). The Notch signaling pathway is an ancient developmental process that regulates the number of cells that assume a particular cell fate through cell-cell based lateral inhibition [37]. Briefly, transmembrane Notch receptors are activated through binding similarly membrane-bound ligands called Deltas or Jaggeds. Because both ligands and receptors are membrane bound, cellecell contact is required for receptor activation. After binding, Notch molecules are cleaved to generate a Notch intracellular domain (NICD) molecule that is then translocated to the nucleus of the cell, where it forms a transcriptional complex with other factors including Rbpj. The results of Notch activation vary, depending on the biological system, but a common effect is the inhibition of cellular differentiation through the induction of expression of a class of inhibitory transcription factors that includes Hes1, Hes5, and Hey2. In the developing cochlea, Notch1 is expressed broadly in the progenitor cells that will develop as both hair cells and supporting cells [38,39]. As cells begin to develop as hair cells, they upregulate expression of two Notch ligands, Jagged2 and Delta1. Binding of these ligands leads to the generation of NICDs and expression of Hes genes in surrounding cells [38,40,41]. Deletion of any component of this signaling pathway or prevention of Notch cleavage by pharmacological g-secretase inhibitors results in an increase in the number of cells that develop as hair cells, which demonstrates that the Notch signaling pathway acts as an inhibitor of hair cell development [38,42]. Whether this pathway is still active in the adult auditory epithelium and could therefore have a role in inhibiting hair cell regeneration, is a matter of debate [43e45]. However, some initial results to be discussed subsequently suggest that this is possible. The demonstration of a role for Notch signaling also provided clues regarding some of the genes that might have a positive role in inducing hair cell formation. As discussed, Notch is part of an ancient signaling pathway that includes Hes and Hey transcription factors. These transcription factors are part of a large family that share common structural motifs including a basic DNA binding domain (b) and a helixeloopehelix (HLH) dimerization domain. Other members of the basic HLH family include molecules that promote cellular differentiation [46]. Screening for basic HLHs expressed in developing hair cells indicated that Atoh1 might be a strong candidate. Results of both in situ hybridization and mouse reporter line studies demonstrated that Atoh1 turns on in a large number of cochlear progenitor cells before hair cell development, but by embryonic day (E)16 to E17 in the mouse, expression is restricted to the developing hair cells [47,48]. More important, genetic deletion of Atoh1 leads to a complete absence of all inner ear hair cells whereas forced expression of Atoh1 in the embryonic inner ear can induce cells to adopt a hair cell fate [47,49,50]. Finally, lineage tracing experiments demonstrated that the number of cells that initially express Atoh1 is greater than the number that ultimately develop as hair cells and that activation of Notch signaling has a critical role in determining which cells will maintain expression of Atoh1 [51]. Based on these results, Atoh1 has been established as a strong inducer of hair cell formation. Following from these experiments, both expression of Atoh1 and modulation of Notch signaling have been examined as possible mechanisms to induce hair cell regeneration in the mature inner ear. Although promising in some respects, the results have unfortunately, been limited. To assess the ability of cells in different regions of the inner ear to develop as hair cells, transgenic mice in which broad expression of Atoh1 can be induced pharmacologically were generated by two laboratories. The results indicated that whereas forced expression of Atoh1 induces ectopic hair cell formation in embryonic or neonatal cochlea, this ability is completely lost in adult inner ear cells [52,53]. Furthermore, even in neonatal tissue, hair cells that were formed in response to forced expression of Atoh1 failed to mature completely. The reasons for both lack of hair cell maturation and loss of responsiveness to Atoh1 are unclear. The most likely explanation for the defects in hair cell maturation are that other factors (either co-factors for Atoh1 or possibly additional transcription factors) are required to drive hair cell maturation. This hypothesis was supported by a study showing that the efficacy and maturity of hair cells derived from embryonic stem cells could be significantly increased if the cells were forced to express three relatively inner earespecific transcription factors: Atoh1, Gfi1, and Pou4f3 [54]. Gfi1 and Pou4f3 are expressed in hair cells soon after Atoh1, but both loss-of-function and gain-of-function experiments indicated that neither Gfi1 nor Pou4f3 is necessary or sufficient for initial hair cell formation [55,56]. Regardless, these results suggest that the combination of these three factors may have the ability to induce more mature hair cells in an adult inner ear epithelium, although this possibility has not yet been directly tested.

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FIGURE 49.3 Regulation of cell fates during inner ear development. Three stages of molecular signaling interactions are illustrated. (A) An initial period of Wnt signaling leads to stabilization and subsequent nuclear-translocation of b-catenin (b-cat). In the nucleus, b-catenin induces expression of TCF/Lef transcription factors that bind directly to an Atoh1-30 enhancer leading to the initial expression of Atoh1. (B) Atoh1 protein binds to multiple DNA recognition sites, including one in the Atoh1-30 enhancer and others in uncharacterized hair cell (HC) genes. (C) In some cells, expression of two Notch ligands, Delta1 and Jagged2, is initiated. Delta1 and Jagged2 bind to Notch1 receptors on neighboring cells, causing the upregulation of inhibitory transcription factors such as HES5. HES5 inhibits the expression of Atoh1, resulting in a downregulation of Atoh1 target genes, including Atoh1. Cells that maintain expression of Atoh1 go on to develop as hair cells whereas cells in which Atoh1 is downregulated develop as supporting cells.

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Regarding the loss of ability of Atoh1 to induce hair cells in more mature inner ear cells, the most likely explanation for this change is progressive epigenetic changes that prevent Atoh1 from binding to and activating essential downstream target genes. Although this type of epigenetic modulation has been reported in other cell types, no definitive studies have been performed in the inner ear. Other possibilities include downregulation of essential co-factors or required post-translational modifications that do not occur in more mature cells. The results described earlier focused on the cochlear epithelium; however, an experiment demonstrated consistent, although not entirely similar, results for the vestibular epithelia. As is true for the cochlea, forced expression of Atoh1 induced hair cell formation in utricular epithelia at early postnatal ages and the efficacy of this effect decreased over time with 75% fewer hair cells formed in response to Atoh1 expression at postnatal day 21 (P21) versus the same level of induction at P0 [57]. However, the formation of new hair cells in the utricle at P21 contrasted with results from the cochlea in which no new hair cells were formed at that age. Unfortunately, the ability of Atoh1 to induce new hair cells in the utricle was not tested at ages older than P21, so it is not clear whether the level of induction observed at P21 represents a plateau or whether further loss of efficiency would occur at older ages. Similar experiments in which Atoh1 expression was induced using viral vectors in more mature animals reported the formation of a limited number of hair cellelike cells in different regions of the inner ear, which is consistent with the possibility that some adult inner ear cells located outside the cochlear duct retain the ability to develop as hair cells [58]. In fact, recovery of some degree of vestibular function was reported in animals in which hair cells were killed using systemic applications of aminoglycosides and loop diuretics [23]. However, as discussed previously, it was difficult to determine the relative contributions of hair cell regeneration and functional compensation in those experiments. During development, activation of the Notch pathway serves to inhibit hair cell formation. This observation raised the question of whether continued activation of this pathway, in particular in response to injury, might have a role in inhibiting hair cell regeneration. Consistent with this idea, examination of the expression of Notch pathway genes in the chicken basilar papilla indicated that both Notch and its receptors are re-expressed during a regenerative response, although these factors were subsequently downregulated once the epithelium completed the regenerative process [59]. Neither Notch1 nor the ligands Delta1 and Jagged2 were observed in the mature intact organ of Corti. However, examination of epithelia after noise trauma yielded contrasting results; one laboratory reported an increase in the expression of Notch pathway genes, whereas a second observed the opposite result [43]. More intriguing are the results of a study in which partial recovery of hearing sensitivity was demonstrated after noise trauma in animals treated with a systemic g-secretase inhibitor [45]. g-Secretase is required for the cleavage of Notch [60]. Although the recovery was modest, the results suggested a potential role for Notch signaling in inhibiting hair cell regeneration. Finally, the Wnt signaling pathway has been shown to have diverse roles in hair cell development. Wnts are secreted glycoproteins that have been implicated in myriad biological processes including cell fate, cellular polarity, and the regulation of cancer [61]. Studies in which a highly conserved target of the canonical Wnt signaling pathway, b-catenin (b-cat), was modulated both in vitro and in vivo demonstrated important roles for b-cat, and presumably Wnt signaling, in cochlear proliferation and subsequent hair cell formation [62,63]. In particular, inhibition or deletion of b-cat causes a significant disruption in hair cell development whereas increased activation of b-cat leads to increased proliferation of cochlear prosensory cells and subsequent formation of extra hair cells. Unfortunately, it has been difficult to identify the specific Wnts and Wnt receptors that may mediate these effects. A study used a combination of microarray analysis and in situ hybridization to identify and profile the expression of Wnts and Wnt receptors in the developing cochlea, but because of redundant expression and function, it is not yet possible to determine which factors might act to activate Wnt signaling within the cochlea. As discussed earlier, under some circumstances, cellular proliferation may be required to facilitate hair cell regeneration. For instance, the loss of structural differentiation that typically occurs before cellular mitosis may increase the ability of an individual cell to undergo a phenotypic switch. In some epithelia, such as the organ of Cori, the limited number and high degree of specialization of supporting cells may necessitate the generation of new cells to maintain a functional epithelium after the conversion of some supporting cells into new hair cells. Cellular proliferation is minimal to non-existent in adult mammalian inner ear sensory epithelia [64]; however, before hair cell differentiation, the progenitor cells that will develop as hair cells are highly proliferative. Understanding how cellular proliferation is promoted or inhibited in these cells could provide valuable insights regarding possible approaches to inducing mature supporting cells to reenter the cell cycle. Extensive previous studies have identified a large number of factors that modulate cell cycle progress and exit. The first of these to be examined in the developing cochlea was the cell cycle inhibitor Cdkn1b (formerly p27kip1). Cdkn1b is specifically expressed in all cochlear prosensory cells before terminal mitosis; expression is maintained in supporting cells through

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adulthood [65,66]. Deletion of Cdkn1b results in approximately one additional round of mitosis within the prosensory population, leading to the generation of supernumerary hair cells and supporting cells and to auditory dysfunction. Subsequent experiments demonstrated important roles for a number of other cell cycle regulators including cyclin D1, N-myc, retinoblastoma, p130, p19, and p21 [67]. Deletion of factors that promote proliferation such as N-myc leads to reduced cells within the inner ear and significant defects in cellular patterning, whereas deletions of cell cycle inhibitors leads to supernumerary cells and some level of cell cycle reentry in adults, although in most cases those cells undergo subsequent cell death. Overall, studies on the development of the inner ear sensory epithelia have provided valuable insights regarding the genes and signaling pathways that regulate processes relevant to hair cell regeneration. In particular, identification of the transcription factor Atoh1 as a strong inducer of a hair cell fate and the demonstration that the Notch signaling pathway has a critical role in inhibiting cells from differentiating as hair cells suggest two possible approaches to enhancing hair cell regeneration in an adult epithelium. Similarly, the discovery that the Wnt signaling pathway and cell cycle regulators act to modulate proliferation of inner ear progenitor cells may provide insights regarding methods to induce cell cycle reentry in an adult inner ear. Results described in the next section will discuss initial attempts to induce regeneration by modulating these pathways.

INDUCTION OF HAIR CELL REGENERATION USING TRANSGENIC MICE The development of inducible transgenic mouse lines in which specific genes can be activated or inactivated in specific cell types at specific time points has revolutionized the ability to study inner ear regeneration. Based on the results of the developmental studies described earlier, four genes (Cdkn1b, Atoh1, Notch1, and b-cat) have been targeted. In most cases, two general concepts were addressed individually or together. The first of these is that forced activation of Atoh1 alone or combined with the deletion of Notch1 can enhance hair cell differentiation by expressing a hair cell inducer (Atoh1) and removing a hair cell inhibitor (Notch1). The second is that removal of the cell cycle inhibitor Cdkn1b might allow supporting cells to reenter the cell cycle. Finally, activation of the Wnt pathway has been implicated in both proliferation and differentiation, and so it might be able to mediate both events. For the bulk of these projects, the basic experimental design has been to use supporting cell-specific inducible cre lines combined with floxed deletion or activator lines to modulate one or more of these factors after injury [68,69]. These approaches have yielded encouraging results after induced hair cell loss in neonatal cochleae; unfortunately, similar results have not been obtained in adult inner ears. Because newborn mouse pups do not begin to hear until approximately 14 days postpartum, it seems possible that the results obtained in neonates may be a remnant of the embryonic developmental program or immaturity of the supporting cells. Whether the inability of these factors to induce new hair cells in adult cochleae is a result of changes in post-transcriptional or post-translational processing, loss of obligate co-receptors, epigenetic changes, or a progressive loss of stem cells within the epithelium remains to be determined. The results of two studies shed some light but also provide some confusion regarding the mechanism that might act to prevent hair cell regeneration in the cochlea. As discussed, activation of Notch1 requires a g-secretasee dependent cleavage of the Notch1 protein leading to the formation of NICD, which is then translocated to the nucleus to initiate signaling. Several pharmacological inhibitors of g-secretase have been shown to block Notch1 activation effectively in vitro. Mizutari and colleagues induced hair cell damage in adult mice using noise or driving caspase expression in hair cells to promote apoptosis and then injected LY411575, a new, highly potent g-secretase inhibitor into the inner ears of these animals for several days [45]. The mice in these studies also carried a lineage marker that allowed the authors to mark all of the supporting cells in the cochlea permanently. After 3 months of recovery, the LY411575-treated animals showed some improvement in auditory function and evidence of replacement hair cells in the cochlea. Moreover, the new hair cells expressed the lineage marker, which suggested that they had developed from existing supporting cells, although it is not clear whether they were exclusively from nonmitotic conversion or whether any proliferative regeneration was induced. This result was particularly exciting in that the use of a pharmacological agent is well-suited for development as a clinical application. From a biological standpoint, this result suggested that the Notch pathway remains active or is re-activated after injury in adult tissue. In fact, both this study and work from a separate laboratory indicated re-expression of some components of the Notch pathway in response to injury [44]. However, a subsequent study by a different group of researchers observed a different result. In that case, explant cultures of neonatal cochleae were established, treated with g-secretase inhibitors and then assayed for the development of new hair cells. The results indicated a marked decrease in the number of new hair

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cells that developed even in cochleae as young as P3. In addition, polymerase chain reaction analysis of the expression of Notch pathway genes after noise damage in adult animals indicated no reactivation of the Notch pathway [43]. Thus, these two studies would seem to present two highly disparate results for which the bases for the differences cannot easily be determined. One possibility is that the mechanisms of hair cell damage were different, as were the time scales. Another possibility is that although Notch genes are a main target of g-secretase, they are not the only target, which raises the possibility that the regeneration observed in vivo is mediated through a different pathway. Therefore, although the results are intriguing, additional studies are clearly required.

STUDIES OF HAIR CELL REGENERATION USING THE LATERAL LINE As discussed, in mammals and other terrestrial vertebrates, hair cell sensory epithelia are confined to the inner ear. In contrast, aquatic vertebrates, such as fishes and amphibians, possess an additional hair cell sensory structure, the lateral line system (Fig. 49.4). This sensory system is composed of a series of organs, termed neuromasts, arrayed along the surface of the skin or in bony dermal canals that connect to the skin surface (Fig. 49.4A). The amphibian lateral line is indeed the first place that hair cell regeneration was observed. After amputation and subsequent

FIGURE 49.4 Overview of the lateral line. (A) Schematic diagram of the lateral line in a 4-day-old zebrafish embryo. The green circles represent the position of neuromasts (NM) and the green lines represent the underlying nerve. The neuromasts in the head make up the anterior lateral line, whereas those along the body make up the posterior lateral line. (B) Schematic of a surface view of an individual neuromast in a 3- to 4-day-old zebrafish. The mantle cells make up the outer edge of the neuromast and the interneuromast cells are a line of cells connecting adjacent neuromasts (dark green). The supporting cells (light green) surround the sensory hair cells (red). (C) Cross-section of a neuromast in a 3- to 4-day-old zebrafish. The apical side of the neuromast is constricted and the hair cells (red) extend their hair bundles into an overlying gelatinous cupula (yellow). Mantle cells and interneuromast cells (dark green) lie at the edge of the neuromast, with supporting cells (light green) extending the width of the epithelium and interdigitating between the hair cells. (D) Schematic of the development of the posterior lateral line. The leading edge of the primordium (arrow) has migratory mesenchymal cells that crawl forward, driving the primordium along the horizontal myoseptum of the fish from head to tail. Cells in more posterior positions within the primordium begin to organize into rosettes, which are deposited as the primordium continues to migrate. In the deposited neuromast, centrally positioned cells become the Atoh1-expressing hair cell precursor (yellow) and organize the neuromast.

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regrowth of larval salamander tails, Stone reported in the 1930s that replacement neuromasts, which contain new sensory hair cells, were produced from cells located in the most distal remaining neuromast [70,71]. As a result, the amphibian lateral line was used as a model for both sensory organ regeneration and hair cell regeneration for many years. The development of zebrafish as a genetically manipulatable model system has allowed more extensive investigation of the molecular control of these regenerative phenomena. The mechanosensory lateral line is used to detect the flow of water and water pressure differences across the fish’s or amphibian’s body and has important roles in schooling behavior, predator and prey detection, rheotaxis, obstacle avoidance, and, potentially, communication between individuals [72e74]. Each neuromast is composed of a number of sensory hair cells and supporting cells and, at the outermost edge, a layer of mantle cells (Fig. 49.4B,C). The lateral line sensory hair cells are morphologically and physiologically similar to those in the inner ear, and although there are fewer markers or identifying features, the supporting cells also appear to be similar to those in the inner ear [75e77]. During development, the lateral line system is established by several different primordia that migrate from the region of the otic placode along the head and trunk, progressively depositing clusters of cells that organize into neuromasts (Fig. 49.4D). Whereas the migrating primordia establish an initial series of neuromasts, additional, new neuromasts arise from latent, multipotent interneuromast cells that are also deposited by the initial primordia as well as from budding of new neuromasts from existing ones, ultimately to form stitches, or linear arrays of neuromasts, along the body [70,71,78e80]. Once a neuromast begins to develop, the initial formation of hair cells is coordinated by atoh1 and Notch signaling, similar to the development of hair cells in the vertebrate inner ear (see earlier discussion). Cells in the center of each new neuromast acquire atoh1 expression, becoming hair cell precursors, and those cells use Notch and fibroblast growth factor (FGF) signaling to suppress hair cell precursor fate in neighboring cells [80e82]. The hair cell precursors then divide using planar polarity cues oriented during the migration of the primordium, such as vangl2, to produce pairs of hair cells oriented in the opposite direction; the hair bundles and kinocilia are oriented 180 degrees away from each other, with the kinocilia of both located closest to the center of the plane of division [83,84]. The similarity of the genes, signals, and patterns used to develop sensory hair cells within the lateral line and the vertebrate inner ear has suggested that studying the robust regeneration that occurs in the lateral line will provide valuable insights into the mechanisms that control (and limit) hair cell regeneration in all systems.

FORMATION OF NEW NEUROMASTS FROM MULTIPOTENT PROGENITORS As mentioned previously, the first evidence that hair cells could be regenerated from supporting cells came from early experiments in salamander tail regeneration, in which it was noted that replacement neuromasts were derived from the outermost (mantle) cells on the posterior edge of the distal-most remaining neuromast after tail amputation [70,71]. Experimentation revealed that the anterior mantle cells also had the capacity to regenerate lost neuromasts if they were rotated to be proximal to the amputation site, and that these mantle cells were responsible for the budding of new neuromasts during stitch formation. These results suggested that the mantle cells retained an intrinsic latent multipotency that could be stimulated to proliferate, forming new or replacement neuromasts [70,71,78,85e87]. It has also been suggested that macrophages responding to the injury may have an important role in initiating regeneration [87], and that the regenerative primordium may reexpress markers used in the developmental primordia [88], but little is known about the signals that initiate either neuromast budding or regeneration from the mantle cells (Fig. 49.5). In addition to the mantle cells, it has been suggested that interneuromast cells, which are deposited by the developmental primordia between primary neuromasts, may also serve as a pool of multipotent progenitors capable of forming new or replacement neuromasts. In fact, several studies suggested that the glia ensheathing the lateral line nerve, which runs beneath the lateral line, suppress the interneuromast cells from forming new neuromasts, and that interstitial growth may therefore come from de-repressed interneuromast cells that escape glial inhibition [89e92]. In addition, in response to localized destruction of an entire neuromast, interneuromast cells are capable of replacing the missing neuromast, and this regeneration is enhanced by blocking the development of lateral line nerve glia [93]. Thus, there appear to be at least two populations of latent multipotent progenitors capable of producing new sense organs in the lateral line: the mantle cells and the interneuromast cells, with both populations held in check until injury or growth requirements stimulate them (Fig. 49.5A). Notably, although full molecular characterization of mantle and interneuromast cells is incomplete [94], many common markers are expressed in both populations, which raises the possibility that these are highly similar cells, even though the mantle cells are epithelial whereas interneuromast cells appear to be more mesenchymal.

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FIGURE 49.5 Models of progenitor cells in the lateral line. (A) Schematic diagram of a neuromast showing the areas where multipotent

progenitors (interneuromast cells and mantle cells) can develop into new neuromasts. Interneuromast cells can divide and form a new neuromast between two existing neuromasts. Mantle cells can proliferate on the dorsoventral sides of the neuromast to bud off additional neuromasts during stitch formation. In response to amputation of the tail, the posterior neuromasts divide to become a migratory regenerative primordium that forms and deposits replacement neuromasts as it migrates into the regenerating tail. (B) Model of hair cell regeneration in the lateral line. When hair cells are ablated (Xs), they are rapidly extruded from the epithelium. Adjacent supporting cells divide symmetrically to develop into two replacement hair cells, and more peripheral supporting cells divide to give rise to replacement supporting cells.

HAIR CELL REGENERATION IN THE LATERAL LINE The ability to regenerate entire sensory organs within the lateral line demonstrates a robust regenerative capacity; moreover, the ability also to regenerate just damaged or missing hair cells within the sensory organs makes this an important system for working out the mechanisms underlying hair cell regeneration. Like inner ear hair cells, lateral line hair cells are susceptible to damage from aminoglycoside antibiotics and chemotherapeutics such as cisplatin [95e97], and the presence of sensory structures on the surface of the fish means that hair cells can be killed simply by placing a fish in water containing ototoxins for 15e60 min. The result is rapid hair cell death and extrusion from the neuromast. The ease of applying such compounds to the lateral line hair cells, particularly in zebrafish larvae, has been leveraged to develop several screens for the discovery both of new potential ototoxins, such as copper ions and other clinically relevant compounds [98,99], as well as compounds that serve protective functions [100e103]. In addition, the surface location of the cells allows for the direct ablation of individual cells, such as by laser irradiation, and tracking of the regenerative response [104,105].

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Notably, even without the addition of ototoxic compounds or other insults, there is a significant and ongoing turnover of hair cells in lateral line neuromasts [106,107], similar to that seen in avian vestibular organs [108e110]. The correlation between levels of cell death and levels of proliferative regeneration both at baseline and after the induction of damage suggests that each neuromast uses a feedback mechanism to maintain the number of hair cells within a certain range [106]. After an acute insult, loss and regeneration of hair cells occur rapidly; hair cell loss and extrusion occur within 3 h and the first regenerated hair cells appear within 12 h. Baseline numbers of hair cells are typically restored within 72 h [96,111]. Similar to results from nonaquatic vertebrates, studies in both amphibian and zebrafish lateral line have clearly identified the supporting cells as the source of regenerated hair cells [83,96,97,104e106,111,112]. Most studies suggest that regenerated hair cells are formed as a pair after the single division of a supporting cell [83,111e113]; indeed, treatment with mitotic inhibitors inhibits hair cell regeneration [97,112,114]. However, there are reports that a few regenerated hair cells may be derived via nonmitotic conversion of a supporting cell into a hair cell [105,111,115]. Because there is constant turnover within neuromasts, this nonproliferative conversion may reflect the presence of some recently produced, undifferentiated progeny at the time of lesion, or it may be tied to the specific type and level of damage induced. However, this type of regeneration is composed of only a minor percentage of new hair cells; the vast majority originate from the proliferation of nearby supporting cells. The ability of the supporting cells to regenerate hair cells extends from larval stages through to adulthood with little loss in capacity and only slight changes in the kinetics, and there is little evidence that the population of progenitors can be exhausted, because repeated rounds of neomycin still lead to robust regeneration [107,116]. Because regenerated hair cells appear to be produced as a pair from the division of one supporting cell [83], repeated rounds of regeneration would lead to a depletion of supporting cells (and loss of regenerative capacity) without proliferative replacement of the supporting cell population. However, the fates of supporting cell divisions appear to be compartmentalized based on their position within the neuromast. More central supporting cells divide to produce hair cells, whereas more peripheral supporting cells divide to renew the supporting cell population. Differences have been observed among mantle cells, with more anterior supporting cells representing a slower dividing or quiescent pool [107,117]. Mantle cells rarely appear to contribute to regeneration after limited damage, but they can contribute to the production of new supporting cells after more extensive damage such as the forced depletion of the supporting cell pool [118]. Thus, although mantle cells may be a multipotent progenitor for new sensory organ formation, they may not have regular roles in maintaining and regenerating hair cells and supporting cells, although the relationship between mantle cells and peripheral support cells, such as whether they have nichelike interactions, needs to be clarified.

PATHWAYS COORDINATING HAIR CELL REGENERATION IN THE LATERAL LINE The ability to use both forward and reverse genetics techniques in zebrafish, the strength of various gene expression assays, as well as the ease of adding pharmacological compounds to the water has allowed extensive investigation of the pathways and genes coordinating the regeneration of hair cells after damage and comparison of those with both the developmental pathways and regeneration in other vertebrates. Transcriptomic analysis of supporting cells and mantle cells has begun to characterize genes that are active during different windows of regeneration [94,118]. Shortly after hair cell loss, Wnt, Notch, and FGF signaling pathways are each inhibited; Wnt and Notch become active in later stages of regeneration [118]. In particular, Wnt10a and the Frizzled 7b and 8a receptors are regulated during the first 5 h after hair cell ablation [118]. Consistent with these results, several studies found that the inhibition of Wnt signaling, either genetically or pharmacologically, blocks supporting cell proliferation and hair cell regeneration, whereas the activation of Wnt signaling promotes supporting cell proliferation [117,119e121]. Similarly, the inhibition of Notch signaling promotes increased numbers of supporting cells returning to the cell cycle and biases the progeny to differentiate into hair cells, whereas the activation of Notch blocks supporting cell proliferation and hair cell regeneration [111,112,117]. These pathways appear to be hierarchically arranged, because Notch inhibition activates Wnt signaling in supporting cells, likely via the loss of Notchmediated expression of the Wnt-inhibitor dkk2 [117,120]. This relationship is similar to that reported for Wnt/Notch control of the proliferation of supporting cells in the mouse utricle [122]. Although FGF signaling is modulated after hair cell death and is known to be important in the initial development of hair cells within neuromasts [81,82,118], the specific role of FGF signaling in hair cell regeneration is less clear. During regeneration, hair cell progenitors appear to have active FGF [123], similar to the progenitors during development, and inhibition of FGF signaling or ablation of FGF receptor 1-expressing support cells reduces hair cell regeneration [124]. Retinoic acid

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(RA) signaling also seems to be an important inducer of supporting cell proliferation by altering p27kip, and although RA and FGF signaling often interact, they seem to be independent during hair cell regeneration in the lateral line [123]. In addition to studies finding signals required for hair cell regeneration, negative regulators have been found. An insertional mutagenesis screen identified N-glycosylation by mgat5a as an important negative regulator of hair cell regeneration, because mutations in this gene had increased regeneration [125]. Although the specific pathway(s) modulated by mgat5a is not known, transforming growth factor-b (TGFb) has been suggested as a candidate. Consistent with this idea, inhibition of Stat3, which is typically downstream of TGFb signaling, enhances hair cell regeneration [126], although the early upregulation of Stat3 suggests that it may also have a role in initiating the regenerative response [118,126]. Studies have also begun to examine how epigenetic control of these signaling pathways may mediate regeneration, in particular focusing on histone methylation and its control of Wnt and FGF signaling [127,128]. Many of the most heavily studied pathways seem to mediate proliferation and/or differentiation of the hair cell progenitors, but it remains unclear what signals initiate the regenerative response, because none of the pathways that have been studied appear to be sufficient to stimulate supporting cell proliferation. Transcriptomic studies have identified other cellecell signaling pathways including insulin, mitogen-activated protein kinase, tumor necrosis factor-a, nitric oxide, reactive oxygen species, Fat, and integrins, that are activated during regeneration, but specific analysis of the roles of these pathways during hair cell regeneration are still required [94,118]. Several of these candidate pathways, such as nitric oxide and reactive oxygen species, could be tied to pathways active in and released by dying hair cells. It has also been proposed that hair cell death may stimulate regeneration via recruitment of immune cells that secrete cytokines, because macrophages migrate to the sites of hair cell damage before the initiation of proliferation [104,105] and ablation of recruited macrophages delays hair cell regeneration [129]. Although there have been similar proposals for macrophage involvement in initiating hair cell regeneration in the avian inner ear [130e133], ablation of macrophages does not appear to limit hair cell regeneration in that system [134]. In contrast to the extensive study of hair cell regeneration in the fish and amphibian lateral line, less is known about the pathways underlying regeneration within the inner ears of fish and amphibians. Hair cells lost from the saccule, utricle, and cristae after aminoglycosides, laser ablation, or acoustic overstimulation are regenerated within 1e7 days from sox2-expressing supporting cells [123,126,135e137]. RA and FGF signaling appear to regulate inner ear and lateral line regeneration similarly, [123]. Interestingly, although proliferative regeneration was found in cristae and saccule [123,135], rapid recovery of hair cells in the utricle after laser ablation occurred without proliferation of the supporting cells [136], which suggests a potential capacity for direct phenotypic conversion.

OPEN QUESTIONS ABOUT LATERAL LINE REGENERATION The lateral line system provides a unique opportunity to study hair cell regeneration at multiple levels. As described earlier, it has provided valuable insights regarding the regeneration process. One of the central questions still to be answered is whether there are unique populations of supporting cells responsible for regenerating hair cells, or whether all supporting cells have this capacity and instead environmental factors regulate which cells respond to injury. In either case, the specific gene interactions that distinguish the subset of supporting cells responsible for hair cell regeneration (in contrast to quiescent or self-renewing supporting cells) remain to be clarified. Similarly, whether there are distinctions between subpopulations of mantle or interneuromast cells in the lateral line is unclear. The ability of mantle cells and supporting cells to generate hair cells, directly in the case of supporting cells and secondarily in the case of mantle cells, suggests some common stem cellelike properties, but also important differences in their control. The separation of hair cell production and self-renewal of supporting cells into separate populations of symmetrically dividing cells is different from the mixture of symmetric and asymmetric (i.e., producing both hair cells and supporting cells) divisions reported in hair cell regeneration in the chick [18]. Whether this represents a unique adaptation of the neuromast or is a general property of hair cell regeneration in fish will require more careful examination of regenerative proliferation in the fish inner ear. Similarly, open questions still surround the potential role of nonproliferative regeneration among the lateral line, the utricle, and the saccule: in particular, whether the differences are more tied to the levels of damage that is induced, relate to different temporal effects, or reflect differences among these sensory organs.

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CONCLUSIONS All vertebrates use sensory structures containing mechanosensory hair cells to perceive sound, motion, and gravity. Although the overall structure of both the hair cells and their surrounding sensory epithelia appear to be largely comparable among orders, only mammals have lost almost all ability to regenerate hair cells. As a result, humans and all other mammals have the potential to develop permanent deficits in auditory and/or vestibular function. The biological basis for the loss of regenerative ability in mammals remains unclear. In an effort to understand the molecular bases for hair cell formation and regeneration, two general strategies have been developed: an examination of the development of hair cells in mammals and studies of regeneration in other vertebrates, in particular zebrafish. Results from both of these studies have identified general concepts that appear to apply to hair cell formation and regeneration in all vertebrates. Hair cells typically arise from surrounding supporting cells that become activated in response to injury. The nature of that response can vary based on the extent of the injury and the cells that respond. In some cases, supporting cells directly convert into hair cells, whereas in more severe situations, different classes of supporting cells may reenter the cell cycle to generate new progenitor cells before differentiation. In both developing and regenerating epithelia, the number of cells that develop as hair cells is mediated through the Notch signaling pathway, with developing hair cells expressing ligands that bind to and activate Notch in neighboring cells preventing those cells from developing as hair cells. Similarly, the transcription factor Atoh1, which is inhibited by the Notch pathway, acts as a positive regulator of hair cell fate for all vertebrate hair cells. Based on these results, efforts have been made to manipulate both Atoh1 expression and/ or the Notch pathway to induce hair cell regeneration in a mature mammal auditory organ. The results of those experiments are equivocal. Some studies suggest possible regeneration in animal models whereas others have concluded the opposite. Although subsequent work will be required to determine whether these or other factors or pathways hold the key to inducing hair cell regeneration in a mature mammalian epithelium, progress that has been achieved has been remarkable and suggests that a clinical therapy for hearing loss may be developed in the not too distant future.

CLINICAL TRIAL Based on the results of some of these experiments, a multicenter Phase I clinical trial to examine the safety of introducing Atoh1 into the inner ears of patients was initiated in October, 2014 (NCT02132130) by Novartis. The general concept of the trial was to use an adenoviral vector to drive expression of Atoh1 in the ears of patients with profound hearing loss. As a Phase I trial, the primary end points of this study were concern regarding patient safety and the ability to tolerate the adenoviral injections. However, some patients may have been examined for possible recovery of auditory function. Although the results of this trial will not be available for some time, its approval marks the first step toward the application of gene therapy to the inner ear. As discussed, the ability of Atoh1 to induce hair cell formation in cells of an adult ear is probably limited, but if the results indicate that gene therapy can be applied safely to the inner ear, these results have the potential to usher in a series of clinical trials based on manipulations of the signaling pathways discussed in this chapter. Which of these pathways, if any, have the potential to induce a meaningful recovery of function remains to be determined. Nevertheless, that a trial of this nature has been initiated is a remarkable event in and of itself.

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Mechanosensory signaling as a potential mode of communication during social interactions in fishes. J Exp Biol 2016;219(18):2781e9. ˚ , Jørgensen JM. The ultrastructure of lateral line sense organs in the juvenile salamander Ambystoma mexicanum. Cell Tissue Res 1974; [75] Flock A 152(3):283e92. ˚ . The ultrastructure of lateral line sense organs in the adult salamander Ambystoma mexicanum. J Neurocytol 1973;2(2): [76] Jørgensen JM, Flock A 133e42. [77] Nicolson T. The genetics of hearing and balance in zebrafish. Annu Rev Genet 2005;39:9e22. [78] Ledent V. Postembryonic development of the posterior lateral line in zebrafish. Development 2002;129(3):597e604. [79] Ghysen A, Dambly-Chaudie`re C. The lateral line microcosmos. Genes Dev 2007;21(17):2118e30. [80] Chitnis AB, Nogare DD, Matsuda M. Building the posterior lateral line system in zebrafish. Dev Neurobiol 2012;72(3):234e55. [81] Nechiporuk A, Raible DW. FGF-dependent mechanosensory organ patterning in zebrafish. 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Replacement of lateral line sensory organs during tail regeneration in salamanders: identification of progenitor cells and analysis of leukocyte activity. J Neurosci 1993;13(3):1022e34. [88] Dufourcq P, et al. Mechano-sensory organ regeneration in adults: the zebrafish lateral line as a model. Mol Cell Neurosci 2006;33(2):180e7. [89] Grant KA, Raible DW, Piotrowski T. Regulation of latent sensory hair cell precursors by glia in the zebrafish lateral line. Neuron 2005;45(1): 69e80.

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[90] Lo´pez-Schier H, Hudspeth AJ. Supernumerary neuromasts in the posterior lateral line of zebrafish lacking peripheral glia. Proc Natl Acad Sci USA 2005;102(5):1496e501. [91] Nun˜ez VA, et al. Postembryonic development of the posterior lateral line in the zebrafish. Evol Dev 2009;11(4):391e404. [92] Lush ME, Piotrowski T. ErbB expressing Schwann cells control lateral line progenitor cells via non-cell-autonomous regulation of Wnt/ b-catenin. eLife 2014;3(3):e01832. [93] Sa´nchez M, et al. Mechanosensory organ regeneration in zebrafish depends on a population of multipotent progenitor cells kept latent by Schwann cells. BMC Biol 2016;14(1):27. [94] Steiner AB, et al. Dynamic gene expression by putative hair-cell progenitors during regeneration in the zebrafish lateral line. Proc Natl Acad Sci USA 2014;111(14):E1393e401. [95] Song J, Yan HY, Popper AN. Damage and recovery of hair cells in fish canal (but not superficial) neuromasts after gentamicin exposure. Hear Res 1995;91(1e2):63e71. [96] Harris JA, et al. Neomycin-induced hair cell death and rapid regeneration in the lateral line of zebrafish (Danio rerio). J Assoc Res Otolaryngol 2003;4(2):219e34. [97] Mackenzie SM, Raible DW. Proliferative regeneration of zebrafish lateral line hair cells after different ototoxic insults. PLoS One 2012;7(10): 1e8. [98] Herna´ndez PP, et al. Sub-lethal concentrations of waterborne copper are toxic to lateral line neuromasts in zebrafish (Danio rerio). Hear Res 2006;213(1e2):1e10. [99] Chiu LL, et al. Using the zebrafish lateral line to screen for ototoxicity. J Assoc Res Otolaryngol 2008;9(2):178e90. [100] Owens KN, et al. Identification of genetic and chemical modulators of zebrafish mechanosensory hair cell death. PLoS Genet 2008;4(2): e1000020. [101] Coffin AB, et al. Chemical screening for hair cell loss and protection in the zebrafish lateral line. Zebrafish 2010;7(1):3e11. [102] Ou HC, et al. Quinoline ring derivatives protect against aminoglycoside-induced hair cell death in the zebrafish lateral line. J Assoc Res Otolaryngol 2012;13(6):759e70. [103] Thomas AJ, et al. Identification of small molecule inhibitors of cisplatin-induced hair cell death: results of a 10,000 compound screen in the zebrafish lateral line. Otol Neurotol 2015;36(3):519e25. [104] Balak KJ, Corwin JT, Jones JE. Regenerated hair cells can originate from supporting cell progeny: evidence from phototoxicity and laser ablation experiments in the lateral line system. J Neurosci 1990;10(8):2502e12. [105] Jones JE, Corwin JT. Regeneration of sensory cells after laser ablation in the lateral line system: hair cell lineage and macrophage behavior revealed by time-lapse video microscopy. J Neurosci 1996;16(2):649e62. [106] Williams JA, Holder N. Cell turnover in neuromasts of zebrafish larvae. Hear Res 2000;143(1e2):171e81. [107] Cruz IA, et al. Robust regeneration of adult zebrafish lateral line hair cells reflects continued precursor pool maintenance. Dev Biol 2015; 402(2):229e38. [108] Jørgensen JM, Mathiesen C. The avian inner ear. Continuous production of hair cells in vestibular sensory organs, but not in the auditory papilla. Naturwissenschaften 1988;75(6):319e20. [109] Roberson DF, et al. Ongoing production of sensory cells in the vestibular epithelium of the chick. Hear Res 1992;57(2):166e74. [110] Kil J, Warchol ME, Corwin JT. Cell death, cell proliferation, and estimates of hair cell life spans in the vestibular organs of chicks. Hear Res 1997;114(1e2):117e26. [111] Ma EY, Rubel EW, Raible DW. Notch signaling regulates the extent of hair cell regeneration in the zebrafish lateral line. J Neurosci 2008;28(9): 2261e73. [112] Wibowo I, et al. Compartmentalized Notch signaling sustains epithelial mirror symmetry. Development 2011;138(6):1143e52. [113] Mirkovic I, Pylawka S, Hudspeth AJ. Rearrangements between differentiating hair cells coordinate planar polarity and the establishment of mirror symmetry in lateral-line neuromasts. Biol Open 2012;1(5):498e505. [114] Namdaran P, et al. Identification of modulators of hair cell regeneration in the zebrafish lateral line. J Neurosci 2012;32(10):3516e28. [115] Herna´ndez PP, et al. Regeneration in zebrafish lateral line neuromasts: expression of the neural progenitor cell marker sox2 and proliferation-dependent and-independent mechanisms of hair cell renewal. Dev Neurobiol 2007;67(5):637e54. [116] Pinto-Teixeira F, et al. Inexhaustible hair-cell regeneration in young and aged zebrafish. Biol Open 2015;4(7):903e9. [117] Romero-Carvajal A, et al. Regeneration of sensory hair cells requires localized interactions between the notch and wnt pathways. Dev Cell 2015;34(3):267e82. [118] Jiang L, et al. Gene-expression analysis of hair cell regeneration in the zebrafish lateral line. Proc Natl Acad Sci USA 2014;111(14):E1383e92. [119] Head JR, et al. Activation of canonical Wnt/b-catenin signaling stimulates proliferation in neuromasts in the zebrafish posterior lateral line. Dev Dynam 2013;242(7):832e46. [120] Wada H, et al. Wnt/Dkk negative feedback regulates sensory organ size in zebrafish. Curr Biol 2013;23(16):1559e65. [121] Jacques BE, et al. The role of Wnt/beta-catenin signaling in proliferation and regeneration of the developing basilar papilla and lateral line. Dev Neurobiol 2014;74(4):438e56. [122] Wu J, et al. Co-regulation of the Notch and Wnt signaling pathways promotes supporting cell proliferation and hair cell regeneration in mouse utricles. Sci Rep October 2015;6:29418. [123] Rubbini D, et al. Retinoic acid signaling mediates hair cell regeneration by repressing p27kip and sox2 in supporting cells. J Neurosci 2015; 35(47):15752e66. [124] Lee SG, et al. Myc and Fgf are required for zebrafish neuromast hair cell regeneration. PLoS One 2016;11(6):1e21. [125] Pei W, et al. Loss of Mgat5a-mediated N-glycosylation stimulates regeneration in zebrafish. Cell Regen 2016:1e12. [126] Liang J, et al. The stat3/socs3a pathway is a key regulator of hair cell regeneration in zebrafish. [corrected]. J Neurosci 2012;32(31):10662e73. [127] He Y, et al. LSD1 is required for hair cell regeneration in zebrafish. Mol Neurobiol 2016;53(4):2421e34. [128] Tang D, et al. Inhibition of H3K9me2 reduces hair cell regeneration after hair cell loss in the zebrafish lateral line by down-regulating the wnt and Fgf signaling pathways. Front Mol Neurosci 2016;9(May):1e12. [129] Carrillo SA, et al. Macrophage recruitment contributes to regeneration of mechanosensory hair cells in the zebrafish lateral line. J Cell Biochem 2016;117(8):1880e9.

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C H A P T E R

50 Craniofacial Regenerative Medicine Brandon T. Smith1,a, Emma Watson1,a, Issa A. Hanna2, James C. Melville2, Antonios G. Mikos1, Mark E. Wong2 1

Rice University, Houston, TX, United States; 2University of Texas Health Science Center at Houston, Houston, TX, United States

INTRODUCTION The craniofacial region is composed of several tissue types and a number of niche environments. As such, the need to regenerate injured or diseased tissues requires strategies for engineering both hard and soft tissue, taking into account their surroundings. Tissue engineers have made great strides developing techniques that support and encourage new tissue growth. However, it is imperative that one first appreciates the unique characteristics associated with the defect. With this knowledge, health care providers can select specific approaches that optimize aspects of tissue engineering technology. Additive manufacturing (AM) techniques have become more popular in the clinical setting, and attention has turned to how patient-specific models can be used to plan and execute surgical care whereas patient specific implants can be leveraged to enhance tissue regeneration further within the craniofacial region. This chapter will discuss the craniofacial environment, review current clinical reconstructive practices, and highlight bone tissue engineering strategies with applications in craniofacial reconstruction.

UNDERSTANDING THE CRANIOFACIAL REGENERATIVE ENVIRONMENT Several types of craniofacial defects are associated with different environments; the characteristics of each must be considered if predictable regenerative medicine results are expected. Although comprehensive descriptions of each environment are incomplete, some of the most important features have been identified through a detailed study of the normal development of tissue types and regenerative technologies based on a reproduction of embryological and remodeling biology. Using bone engineering as an example, defect characteristics exist such as [1]: 1. the resident population of pluripotent or multipotent stem cells available for differentiation; 2. vascularity of the defect and the ability of the newly formed tissue to undergo neovascularization; 3. the activity of critical genes, growth factors, and signal transduction agents that mediate tissue formation and remodeling; 4. physical features of the defect that promote tissue formation, including available space and naturally occurring scaffolds; 5. mechanical influences on the defect, including types and magnitude of loads; and 6. interactions between epithelial and mesenchymal elements. In addition to these features, the ability of different craniofacial defects to undergo successful reconstruction is affected by the cause and the presence of infection. We will examine several common defects with different a

These authors contributed equally.

Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00050-3

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environmental characteristics to illustrate the need for customized regenerative strategies and why techniques for regeneration work for some defects but not for others. One of the smallest and most challenging defects to regenerate is the periodontal apparatus that surrounds an erupted tooth and is responsible for its support. The periodontium is composed of an epithelial gingival cuff supported by mesenchymal connective tissue covering the alveolar bone, which forms the tooth socket encasing the root(s). The connective tissue fibers that attach the cementum lining of the tooth root to the socket walls are known as the periodontal ligament; within these fibers lies a network of vascular channels and neural elements with associated cells in various stages of differentiation. Loss of the periodontium typically follows chronic inflammation caused by the accumulation of bacteria- or virus-laden biofilms on the surface of a tooth and root (Fig. 50.1) [2]. The resulting infection results in osteolysis creating a pocket between the alveolar bone and overlying gingival soft tissue and in loss of the periodontal ligament attachments between the bone and root surface cementum (Fig. 50.1). This defect can have single or multiple bony walls. Defects with multiple bony walls constitute more protected environments and regenerative technologies are more successful in these circumstances. Complete restoration of the composite structures within periodontal defects remains elusive, but certain treatment strategies are required to achieve any measure of regenerative success. An appreciation of these factors provides general guidelines for the successful reconstruction of craniofacial defects that can be applied broadly. To begin with, physical removal of the biofilm covering all surfaces of the defect is necessary. Systemic and local delivery of broad-spectrum antibiotics against periodontal pathogens are also beneficial [3,4]. The effect of infection and inflammation on bone formation is a complex topic that acknowledges the important role of proinflammatory mediators in initiating the coordinated processes responsible for bone regeneration. Some important mediators of early inflammation include interleukin (IL)-1, IL-6, tumor necrosis factor-a, and eicosanoids such as prostaglandin (PGE) (e.g., PGE2). Evidence of the role of these mediators is provided by knockout animal studies or the observed effects of anti-PGE2 medications such as nonsteroidal antiinflammatory drugs, which result in compromised bone formation. After these initial events, the activity of proinflammatory mediators abates and a rise in local levels of antiinflammatory mediators such as resolvins, protectins, and lipoxins (autacoids) is responsible for countering inflammation and coincides with the start of the reparative process. Both eicosanoids and autacoids are derivatives of arachidonic acid, and the mechanism for a change in synthesis from inflammatory to antiinflammatory mediators has been characterized as “class switching” and occurs by activating enzymes such as 15-lipooxygenase. If inflammation in a defect site persists as a result of chronic infection, osseous regeneration is diminished. An excellent review of the topic of inflammation and bone regeneration may be found in an article by Thomas and Puleo [5]. Once regeneration begins, the kinetics of the different reparative tissues becomes important. More specialized components of the periodontium, such as the bony walls, ligament, and bone-cementum attachments, take longer

FIGURE 50.1

Periodontal pocket (marked by arrow) produced by loss of alveolar bone under the gingival cuff.

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to form than do epithelium or connective tissue. Unless a physical void is preserved by preventing the ingrowth of epithelium and connective tissue (a technique known as osteopromotion by barrier techniques), periodontal regeneration is compromised. Scaffolds in combination with barrier membrane technology have been shown to be effective in restoring bone volume, and the addition of exogenous growth factors and gene therapy for the local production of growth factors are additional approaches that have been studied in preclinical studies [6]. Larger defects of the craniofacial skeleton extending beyond the periodontal defect differ from each other in a number of physical and biological ways. Defects can be intrabony and surrounded with multiple bony walls (Fig. 50.2A). In this case, the rigidity of the walls facilitates bone regeneration by protecting biological scaffolds such as blood clots, healing tissue, and neovasculature. In addition, periosteum lining the bone surfaces and the underlying endosteal surfaces are an excellent source of cells capable of differentiation under the influence of the proper growth factors. When a bony defect is segmental (Fig. 50.2B), only two bony walls remain, leading to greater instability of the skeletal structure, reduced apposition of osteogenic cell beds, and the potential for adjacent soft tissue to prolapse into the defect, later reducing bone formation. Immobilization of the bone ends (with bone plates) is important to protect early reparative activities such as the secretion of extracellular matrix and neovascularization from mechanical disruption by external loads. Filling such defects with rigid graft material (autologous, allogeneic, or alloplastic) also enhances regeneration by providing a source of living cells or tissue with inductive or osteogenic

(A) 1 2

(B) 1

(C)

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FIGURE 50.2 (A) A five-wall bone defect covered with periosteum (1) provides a protected space that contains a reparative scaffold and cells (2). (B) Segmental defects typically have a reduced number of bony walls with soft tissue (e.g., muscle [1]) adjacent to the defect. (C) Segmental defects with compromised vasculature typically have fewer vessels (1) in the soft tissue envelope, areas of fibrosis (2), and a reduced number of reparative cells (3).

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FIGURE 50.3 Histological cross-section of tissue generated within a bioreactor (initially filled with morcellized autologous graft) after implantation against the sheep periosteum for 9 weeks. High-magnification images taken adjacent periosteum and distal to the periosteum can be found on the right (magnification, 10x). Courtesy Alexander Tatara, MS, PhD; doctoral thesis.

capabilities. The graft can also serve as a scaffold and a barrier to the ingrowth of nonosteogenic tissue. Techniques for reconstructing segmental bone defects using tissue engineering principles are described later in this chapter. As described in the strategies for promoting periodontal regeneration, inflammation and infection have significant roles in the regenerative capabilities of craniofacial defects. Infection of a graft site is a major factor accounting for a complication rate of 48% in a review of the literature on nonvascularized bone grafting [7]. Strategies for reducing infection through implantation of antimicrobial agents will be described in a later section. The paradoxical relationship between inflammation and bone formation raises theoretical questions about the significance of inflammatory states, but the negative effect of prolonged inflammation on bone grafting and fracture repair is well-documented [8]. As such, it is essential to treat infected craniofacial defects by physical debridement and antimicrobial agents for healing or regeneration to occur. Newer strategies to deliver antibiotics locally using devices fabricated through tissue engineering will be described later in this chapter. Other important conditions that can reduce the regeneration of bone include those that compromise the vascularity of the tissue surrounding a bone defect. The role of the vasculature as a source of inducible cells (pericytes), conduit for inflammatory cells (platelets, macrophages, and monocytes) and essential foundation for supporting metabolism in all living tissue is well-known [9]. When the blood supply to a defect site is compromised by therapeutic measures such as radiation therapy, comorbid conditions (e.g., diabetes), and even healing processes (e.g., scarring and fibrosis), the ability of a defect to undergo osteogenic healing is reduced (Fig. 50.2C). Efforts to reconstruct craniofacial defects afflicted by such conditions include adjuvant strategies to mitigate against the compromised vascularity such as the transfer of well-vascularized tissue beds or chemical modulators of angiogenesis. An example of the critical role of the vasculature is provided in this histological section. Fig. 50.3 illustrates bone formation within a chamber filled with autogenous bone particles placed against a vascularized periosteal membrane. New bone formation associated with neovascularization from the periosteum is seen as osteogenesis progresses from the vascularized margin upward.

CURRENT METHODS OF MAXILLOFACIAL RECONSTRUCTION Maxillofacial reconstruction evolved from the use of nonvascularized grafting to the addition of vascularized free flaps in the late 1980s [10]. Nonvascular autogenous bone grafts were considered the reference standard for the repair of bone defects. However, successful osteogenesis from nonvascularized grafts depended on adequate soft

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tissue coverage to isolate the graft site from infection, serve as a source of osteoprogenitor cells, and provide a vasculature supply. Common donor sites for bone harvest included the ilium, tibia, and calvarium [11,12]. Free tissue transfer techniques added the use of vascularized flaps anastomosed to local vessels when defect sites were contaminated or associated with compromised vascularity [13]. The impact of free flap reconstructions has been so profound that it is considered to be one of the most influential advances in head and neck surgery [14]. These flaps are harvested from areas with an axial blood supply that are isolated and anastomosed to vessels close to the defect site to produce an immediately “living” transplant. The most commonly used sites are chosen for their ease in harvest, modest donor site morbidity, and ability to include a large volume and variety of tissue types nourished from a single vascular pedicle [15]. When soft tissue is missing or when a vascularized soft tissue bed is required, the anterior lateral thigh flap or radial forearm flap can be used. For defects requiring composite soft tissue and bony reconstruction, options include the fibula, scapula, and deep circumflex iliac arteryenourished ilium flaps. Technical concerns with these techniques include the morbidity associated with graft harvest, the patency and length of vessels, the duration of surgery, and the recovery time [16]. However, the quality of reconstruction judged by the amount of bone stock is a matter of the patient’s anatomy and may or may not be adequate for functional purposes.

TISSUE ENGINEERING TECHNOLOGIES CURRENTLY USED The field of tissue engineering emerged in the early 1990s as a new way to combine the principles of biology and engineering for the development of functional tissues [17]. Since the inception of tissue engineering, several technologies have entered the marketplace offering hope for patients who have a range of conditions. In the following sections we will review advances in the field of tissue engineering that show promise for the future of craniofacial reconstruction.

Implantable Scaffolds To allow damaged bone to be replaced with functional tissue, engineers have developed a wide range of materials that serve to stimulate the adherence and proliferation of osteogenic cells. Some of these technologies mimic the biomechanical and/or biochemical properties of native bone, whereas others try to recapitulate the anatomy [18e20]. In 1881, Sir William MacEwen of Rothesay used tibial bone wedges from three donors to reconstruct a humeral defect in a 3-year-old child, which represented the first published account of interhuman bone grafting [21]. The procedure was unsuccessful; however, later studies identified the factors affecting graft acceptance and rejection and established the parameters of allogeneic grafting. Nevertheless, despite advancements in developing allogeneic, xenogeneic, and artificial substitutes, autogenous bone grafts remain the reference standard for reconstructing segmental bone defects [22]. By definition, bone grafts can be classified as an autograft, allograft, or xenograft, depending on the source [23]. In contrast, alloplastic implants are synthetically manufactured, inorganic, and biocompatible [24]. Whereas autologous bone grafts produce the most predictable results, alloplastic materials offer several advantages over biologically derived materials. They can be fabricated to fit a patient-specific defect and the resorption rate can be controlled by adjusting material properties and compared with autologous grafts, to provide more material than can be typically harvested from a patient while avoiding a second surgical donor site. For oral and maxillofacial surgery procedures, the US Food and Drug Administration has approved a number of alloplastic materials that can be broadly divided into ceramics and polymers. Ceramics Numerous ceramics are available; calcium phosphate (CaP)-based ceramics represent the most widely used bioactive ceramic. CaPs offer excellent biocompatibility, possess remarkable osteoactivity, and have a chemical and crystalline structure close to those of native bone mineral [25]. Since becoming available for clinical use in 1992, CaP ceramics have been used with success within craniofacial surgery [26,27]. There are two commonly used CaP ceramics: slow-resorbing hydroxyapatite (Ca10[PO4]6[OH]2) and resorbable tricalcium phosphate (TCP) (Ca3[PO4]2) [28]. Both of these formulations can be produced as a paste, which enables the surgeon to inject and mold the cement before final setting occurs, and which makes CaPs an attractive option for dental and orthopedic applications. Depending on the formulation of the CaP, the pore and particle size, and the metabolic activity of the

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recipient site, CaP cement can take 3e36 months to be completely replaced with bone [25]. To accelerate the degradation of CaP-based scaffolds, several studies have investigated the introduction of macropores to increase the surface area of the implant and accelerate tissue integration [29e31]. The use of macroporosity has gained a lot of interest owing to improved degradation kinetics and improved tissue infiltration. Potential applications of CaP-based scaffolds have improved as several studies have investigated techniques to manufacture CaP-based scaffolds through AM techniques. This allows the engineer to improve degradation kinetics further, reproduce tissue architecture more precisely, and improve tissue infiltration [32e34]. Klammert et al. explored the ability to powder-print CaP implants to repair craniofacial defects [34]. In that study, the authors fabricated custom implants by using three-dimensional powder printing to repair defects in human cadaver skulls. The physical and mechanical characteristics of the implants were tested [34]. Whereas the four-point bending strength was in the range of 3.9e5.2 MPa, which is significantly lower than the bending strength of native compact bone (115e209 MPa), these implants could be used to reconstruct the craniofacial skeleton in noneload bearing areas [35]. A porosity of approximately 28e35% could be achieved, increasing the degradation rate and with sufficient fidelity to size and shape so that a printed implant would be able to fit the defect accurately [34]. Polymers Polymeric-based biomaterials first emerged within the field of surgery as a material for suture and fixation devices [36]. Although numerous synthetic polymers have been tried in craniofacial surgery, only a few are extensively used in the clinical setting. Unlike natural polymers, in which degradation depends on enzymes present within the host, synthetic polymers degrade in a uniform and predictable manner by simple hydrolysis [37]. Altering the reaction environment during synthesis can precisely control the observed degradation rate, and even other properties such as mechanical properties, by altering the average molecular weight and size distribution [38]. Specifically within the field of craniofacial surgery poly(glycolic acid), poly(lactic acid), and their copolymer poly(lactic-co-glycolic acid) (PLGA) have been successfully used clinically [39]. Landes et al. examined the efficacy of using PLGA-based implants for maxillary and mandibular osteosynthesis [40]. In that study, five patients (6%) had an apparent foreign body whereas 75 (94%) had no observable reaction. At 24 months, histology revealed few macrophages, giant cells, and minimal bleeding residuals [40]. This demonstrated that PLGA has excellent biocompatibility but there is need for further investigation into the efficacy of bone integration. PLGA has been used as a delivery vehicle for bioactive factors. Several studies investigated the effect of using PLGA microparticles to produce a sustained release of growth factors, signaling molecules, and/or antibiotics [41e44]. These studies demonstrate the applications PLGA microparticles may have in the clinical setting. In addition to PLGA, poly(ε-caprolactone) (PCL) has been investigated extensively within craniofacial tissue engineering owing to some of its unique properties. For example, PCL can form a wide array of biocompatible composites, blends, and copolymers [38]. If one combines PCL with other lactones, the degradation time dramatically decreases from that of the homopolymer of PCL, which is roughly 2e3 years [45]. Hollister et al. showed that the concentration of copolymer can be varied so that it has the ability to support bone formation [46]. Finally, the unsaturated linear polyester, poly(propylene fumarate), can be fabricated into scaffolds by UV-initiated photocross-linking. Although this polymer has been extensively characterized, it has been mainly used within bone tissue engineering applications [47e49].

Bioactive Molecules Just as materials can be leveraged to promote bone growth, bioactive molecules can be introduced to craniofacial defect sites to encourage osteogenesis or angiogenesis. Growth factors are peptides that bind to receptors on a cell, leading to a cellular response. These responses vary by growth factor, cell type, the cell receptor bound, and the time course of growth factor exposure [50]. Clinically, growth factors are used in several applications. They can be employed to augment bone grafting systems, such as with the bone marrow aspirate concentrate (BMAC) technique [51,52]. Growth factors can also be used to stimulate bone formation in heterotopic sites [53e56], an important concept discussed in depth in the section on Bioreactors. Finally, growth factors can be used to form bone without using grafts, employing resorbable scaffolds such as collagen sponges, as seen with sinus augmentation procedures [57e60]. The following section focuses on bone morphogenetic proteins (BMPs) and platelet-derived growth factors (PDGFs), early animal studies, and clinical applications. Many other growth factors (i.e., transforming growth factor-b family, fibroblast growth factor family, insulin-like growth factor, and vascular endothelial growth factor) are important in tissue engineering and regenerative medicine; however, evaluation of these factors in human craniofacial defects is limited and will not be discussed here.

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Bone Morphogenetic Proteins BMPs are a large family of peptides that have a crucial role in bone and cartilage development, as well as some other functions of embryogenesis such as heart and kidney formation [61]. The family acts via two distinct type II and type I serine/threonine kinase receptors; both types are required for signal transduction. BMPs act by stimulating osteoblasts. BMP-2, BMP-4, and BMP-7 show efficacy with bone formation in vivo including in critical-sized defects [62]. The US Food and Drug Administration approved recombinant human BMP-2 and BMP-7 in 2002 for spinal fusion surgeries. In addition, BMP-14 is being tested clinically as a growth factor alternative [63]. There is controversy regarding the use of BMPs in spinal fusion surgeries; a metaanalysis showed efficacy similar to autograft but with more complications, especially when used in the cervical spine [63]. Despite controversy about the use of BMPs in spinal fusion [63], they show great potential for use in the craniofacial area [64e68]. In a rodent study, BMPs have been shown to aid in the healing of critical-sized cranial defects [64]. After loading BMP-2 into PLGA microparticles within a scaffold in a critical-sized, 1-cm defect cranial defect, total union was found in three of eight animals, and the bone volume generated significantly increased over the blank group [65,66]. Larger animals show similar trends. In baboons, a 2.5-cm, critical-sized cranial defect was formed [65]. A collagenous bone scaffold loaded with BMP-7 was introduced to the defect. Groups receiving high doses of BMP-7 (2.5 mg/g scaffold) had extensive osteogenesis, whereas those with low doses (0.1-mg/g scaffold) showed significant increases in bone growth over control animals. A box-type dehiscence defect in the edentulous region of dogs was employed to compare the use of synthetic bone substitute (SBS) and BMP-2eloaded hydrogel [66]. After 8 weeks, the animals were killed and the defects examined. Significantly increased bone volume was seen in the groups containing BMP-2 (14.7 mm3 for the BMP-2eloaded hydrogel and only 2.01 mm3 for the SBS-loaded hydrogel). Several clinical studies were conducted to evaluate the efficacy of recombinant human BMPs (rhBMPs) in human craniofacial defects [67]. In 47 identified case studies in the systemic review, rhBMP-2 was shown to be effective in enhancing bone formation in socket healing and in sinus lift procedures, and BMP-14 was shown to be effective in sinus lift procedures [67]. In one maxillary floor sinus augmentation study, 160 human subjects were enrolled and treated with rhBMP-2 on a collagen sponge or autograft [57]. At 6 months, there was no difference in the change of height between the two groups (7.83  3.52 and 9.46  4.11 mm for the rhBMP-2 and autograft groups respectively), but the density of the bone in the rhBMP-2etreated group was significantly higher. Another clinical study analyzed the effects of rhBMP-2 on the repair of buccal wall defects [68]. Groups received rhBMP-2 (1.5 or 0.75 mg/mL) on an absorbable collagen sponge, just the collagen sponge, or no treatment. The assessment of the alveolar bone showed significantly increased adequacy (as calculated by height and width measured computed tomography [CT] scans and number of repeat procedures needed) over the control or untreated groups. Platelet-Derived Growth Factor PDGF is a dimer composed of PDGF A, B, C, or D joined by sulfide bonds [69]. Several homodimers (PDGF-AA, PDGF-BB, PDGF-CC, and PDGF-DD) and one heterodimer (PDGF-AB) exist. Two isotypes of the receptor exist, a and b, with different binding affinities. PDGF-BB can bind to all receptor isotypes and is often considered the universal PDGF [69]. In vitro studies have shown PDGF is produced by osteoblasts and inhibits osteoclastogenesis [70]. In vivo, PDGF is important in new bone formation and in fracture healing [70]. In humans, PDGF is expressed naturally in the course of fracture healing [71]. PDGF-A chains were found expressed by a variety of cell types throughout the healing process. PDGF-B chains were expressed in a more selective pattern: by osteoblasts at the time of bone formation [71]. Several clinical studies investigated the use of PDGF on the healing of craniofacial defects [72,73]. One study investigated intrabony periodontal defects of greater than 4 mm depth treated with b-TCP alone or b-TCP with recombinant human (rhPDGF)-BB [72]. In the group treated with b-TCP and 0.3 mg rhPDGF-BB, the clinical level of attachment was found to be greater and gingival recession was found to be less at 3 months than that of control patients who received only b-TCP. The clinical level of attachment was not significant at 6 months. Given the significant difference at the early 3-month evaluation with loss of significant differences at the 6-month appointment, rhPDGF-A-BB increased the rate of attachment in the treated groups, providing an advantage for treated patients soon after the surgery. A similar defect treated with rhPDGF-BB was analyzed for the rate of bone turnover [73]. Pyridinoline cross-linked carboxyterminal telopeptide of type I collagen, a well-known biomarker of bone turnover, was shown to be increased at the early study time points of patients treated with rhPDGF-BB, which again indicated advantages for patients with decreased recovery times. rhPDGF has also been evaluated in humans for the treatment of sinus augmentation and ridge preservation [59,60]. Both studies involved using inorganic bovine bone material with rhPDGF-BB as the treatment method, and both studies exhibited good bone growth.

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Bone Marrow Aspirate Concentrate Technique As described in the preceding sections, the principal efforts in tissue engineering have focused on stimulating cell multiplication and differentiation and the development of artificial scaffolds with tunable properties to enhance cell adhesion, vascular ingrowth, and scaffold degradation in coordination with new tissue formation. To meet the goal of reconstructing bony defects, another approach was devised to combine several of the more important contributions in tissue engineering methodology in vivo to take advantage of the body’s established vascular system. In keeping with the classical tissue engineering paradigm, the three essential components are provided with clinically available materials. This technique uses allogeneic bone to maintain space and support the ingrowth of vessels and adhesion of bone-forming cells. Differentiation and regenerative signals are provided by inductive bone proteins such as rhBMP-2, which promotes the differentiation of stem cells and migrating osteogenic cells [74]. Finally, a concentrated regenerative cell population composed of autogenous BMAC is introduced into the defect site. The aspirate contains a highly concentrated population of mesenchymal stem cells (MSCs) (Figs. 50.4 and 50.5) and migrating osteogenic cells that serve as a source of cells for differentiation into osteoblasts. By combining these agents, bone regeneration is promoted, which can lead to successful reconstruction of a bony defect (Fig. 50.6). Several studies reported the efficacy, safety, and ability of rhBMP-2 when combined with osteoconductive grafts to accomplish mandibular reconstruction [51,52]. The use of osteoprogenitor or stem cells from a bone marrow aspirate has improved results and offered another technique to reconstruct craniofacial bone defects. Bone marrow aspirate serves as a rich and readily available source of bone-forming cells (MSCs), which otherwise would not be present in sufficient quantities in a traditional bone harvest. It is easily harvested with simple aspiration through large-bore needles and concentrated with centrifuge devices without significant donor site morbidity [75,76]. Bone marrow transplant/aspirate was first used to treat hemopoietic and oncologic diseases, but it has found additional uses as a cell source in the regeneration of other tissues in the body. Several studies have shown bone marrowederived stem and progenitor cells to be capable of regenerating bone in both animal and human models [77,78]. BMAC has also been used with spine, long bone, and myocardial regeneration [77,79e81]. Of the mixed population of cells collected in an aspirate, Marx and Harrell suggested that CD34þ, CD44þ, CD90þ, and CD105þ cells are the main types of osteoprogenitor cells collected in the concentrate [82]. In a comparison of osteogenic activity, Gimbel et al. studied patients undergoing alveolar cleft repair with different graft materials. A total of 69 patients were divided into three groups. Group 1 underwent grafting with bone marrow aspirate seeded onto a resorbable collagen matrix (n ¼ 21); group 2 received autogenous cortical and cancellous bone harvested from the ilium through a traditional open approach (n ¼ 25); and group 3 received autogenous cancellous bone that was collected from the ilium with a cannula (n ¼ 23) [83]. Whereas alveolar bone formation was comparable in all three groups, the BMAC plus scaffold group experienced significantly less morbidity, operative time, duration of hospital stay, and cost. A study conducted by Hendrich et al. described 101 patients treated with BMAC injections for various bone healing disturbances of the femur including necrosis of the head of the femur (n ¼ 37), avascular necrosis (n ¼ 32), nonunion of fractures (n ¼ 12), and other problems (n ¼ 20) [81]. After an average of 14 months (2e24 months),

FIGURE 50.4 Aspiration of bone marrow from ilium.

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Aspirate after centrifuging, demonstrating separation of cells from plasma and red blood cells.

FIGURE 50.6 Tissue engineered graft for mandibular reconstruction: mixture of bone marrow aspirate concentrate, allogeneic bone, and recombinant human bone morphogenetic protein-2.

the patients were reexamined clinically and radiologically and interviewed. Of the 101 patients, only 2 required additional surgery to correct a nonunion or inadequate bone formation; no additional complications such as infections, excessive bone formation, or harvest site morbidity were observed. The researchers’ conclusion was that BMAC therapy was a suitable alternative to open treatment but required additional studies to determine the full benefits of this treatment modality.

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One of the most exciting applications of the BMAC technique in maxillofacial reconstruction lies in treating large segmental defects of the mandible. Marx and Harrell reported on a series of 40 patients who underwent mandibular reconstruction using this method [75]. All 40 patients achieved successful mandibular continuity that was functionally useful. Melville et al. conducted another study with five patients who underwent immediate intraoral reconstruction with the BMAC technique after resection of benign mandibular tumors [84] (Figs. 50.7 and 50.8). All patients achieved excellent bone quality both clinically and radiographically and were successful candidates for endosseous dental implant placement (Figs. 50.9 and 50.10). This case series demonstrated that composite allogeneic

FIGURE 50.7 Orthopantomogram radiograph of benign tumor (ameloblastoma) of the left mandible.

FIGURE 50.8 Orthopantomogram radiograph 1 year after resection and reconstruction with bone marrow aspirate concentrate, allogeneic bone, and recombinant human bone morphogenetic protein-2.

FIGURE 50.9 Histology of tissue engineered mandibular reconstruction bone (hematoxylin-eosin stain., 10 magnification). Normal reactive bone with regular trabecular pattern with fibrosis. No remnants of cadaver bone were seen after 8 months.

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FIGURE 50.10 Placement of dental implants in tissue engineered reconstructed bone. Alveolar ridge demonstrating excellent height and width for placement of four 5  13-mm implants.

bone, BMAC, and rhBMP grafts constitute an effective and predictable technique for immediate reconstruction of mandibular continuity defects. In addition, there was no donor site morbidity, intraoperative time was reduced, hospital stay was shorter, and total costs were lower compared with more traditional methods of mandibular reconstruction.

Bioreactors If the desire is to use living biological tissue and avoid donor site morbidity, another strategy employs bioreactors to grow tissues in geometries matching the defect in either in vitro or in vivo environments [85]. In vitro bioreactors have been used to engineer tissues such as the temporomandibular joint (TMJ) or ear perichondrium [86,87]. Using human MSCs obtained from fresh bone marrow aspirate, researchers were able to observe confluent layers of lamellar bone and osteoids on the decellularized bone scaffold [86]. This required the use of a bioreactor of the same size and dimensions as the TMJ and a constant flow of media across the scaffold. Other types of in vitro bioreactors are also capable of generating viable tissue. A rotating wall bioreactor has been successfully used to cultivate elastic cartilage [87]. Tissue obtained from patients were differentiated into cartilage progenitor cells and injected into a porous scaffold of collagen, hydroxyapatite, and chondroitin sulfate. After culture for 6 weeks, differentiated chondrocytes and elastic fibers were observed in histology. Although the formation of tissue engineered grafts is possible in vitro, it is important to consider how these tissues will obtain nutrients in vivo. For small tissues, the transport of nutrients and byproducts may be adequate; however, for large tissues, the diffusion rates may not be adequate for the tissue to survive [88]. By introducing vasculature into the construct, cells deep within the tissue can receive appropriate nutrition and gas exchange. In vivo bioreactors have been used to generate vascularized flaps that match the geometry of the defect [85,89e91]. Using strategies previously discussed, in vivo bioreactors can be designed to encourage bone growth in a chamber of desired dimensions. As with osteopromotion strategies, growth factor incorporation or cell seeding can be used directly within the defect to encourage bone growth. These techniques can be used to grow bone in bioreactors in animal models [89e91]. Using MASTERGRAFT, a clinically available scaffold consisting of hydroxyapatite and b-TCP, bone can be grown adjacent to rib periosteum in a large-animal model [89]. Because of the close proximity to the intercostal arteries and veins running below the ribs, tissue from the bioreactor chambers could be removed with the nearby vasculature and transferred to a mandibular defect as a vascularized flap. Whereas some chambers were filled with autograft (requiring harvest from elsewhere in the body), those filled with synthetic graft allowed for the formation of mineralized tissue and reconstruction of large mandibular defects. Incorporation of growth factors to in vivo bioreactors can also be used to promote advantageous ectopic bone growth [90]. BMP-7 was added to Bio-Oss (a scaffold derived from bovine bone) and placed in a pouch created within the latissimus dorsi muscle. After 6, 12, and 24 weeks, tissue was harvested and analyzed. Analysis showed the formation of new bone with vascular supply that could be used for mandibular reconstruction. As an alternative to loading scaffolds with growth factors, cells can be seeded onto scaffolds to promote tissue growth in in vivo bioreactors [91]. Porous, degradable PCL scaffolds were seeded with MSCs from neonatal rats. After 4 weeks of in vitro culture, the cell-seeded scaffolds were implanted in the omenta of rats. Four weeks later, the scaffolds were harvested

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and mineralization was observed throughout the scaffold. However, the formation of vasculature allowing for flap transfer was not investigated. With strong support for the efficacy of in vivo bioreactors in animal models, several bioreactor systems have been used clinically [53e56,92]. The earliest reported case involved angle-to-angle mandibular resection owing to a recurring ameloblastoma [53]. The mandibular-shaped chamber was loaded with autologous bone harvested from the iliac crest and BMP before implantation within the fascia above the scapula. After 4 months, the flap was harvested and transferred with vasculature to the mandible. Competence was restored for liquids and solids, but the patient did not regain the ability to swallow solids. Another clinical case study involved Bio-Oss, bone marrow, and rhBMP-7 within a titanium mesh [54]. The titanium mesh was shaped using CT scans to ensure a close match with the native patient architecture. After 7 weeks within the latissimus dorsi muscle, the mineralized tissue, titanium mesh, and adjacent vasculature were transferred to the mandible. The flap showed remodeling after transfer and the patient had an improved level of mastication after reconstruction completed. Unfortunately, after 13 months, the mesh fractured and the gingiva became disrupted, resulting in infection and necrosis of the mandible replacement [92]. The patient died soon thereafter of cardiac arrest. Many of the human in vivo bioreactor studies require the need for morcellized autograft or bone marrow. Hydroxyapatite with rhBMP-7 was capable growing mineralized tissue in the pectoralis muscle [55]. This tissue was then transferred to the mandible. The patient could talk and eat food; however, 5 months after transfer, the flap became infected. Another clinical study used CT images to shape a titanium mesh cage [56]. Bone blocks, rhBMP-2, and bone marrow aspirate were placed within the cage before implantation in the gastric omentum. After 3 months within the gastric omentum, the free tissue flap was harvested and used to reconstruct the mandible. Three months after transfer, the patient could talk and eat normally.

Adjuvant Therapies Antibiotics When repairing craniomaxillofacial tissue defects, proximity to bacterial flora is an important consideration. The oral cavity hosts over 500 bacterial species, including many species of Streptococcus [93], and the nasal cavity has been shown to be colonized by Staphylococcus aureus in 21% of patients at admission [94]. Although native bacteria generally do not cause problems, trauma or the following reparative surgery can shift the bacterial balance in favor of pathogenic bacteria or permit the passage of bacteria to previously aseptic sites. In facial fracture repair, infection rates can be as high as 42% if no antibiotics are received before surgery [95]. The use of cefazolin sodium has been shown to reduce the rate of infection to 9% during the surgical repair of facial fractures [95]. Although antibiotics are commonly provided systemically before many craniofacial surgeries, a systemic review and metaanalysis of endoscopic sinus surgeries showed no statistical difference in infection rate between patients who receive antibiotics and those who do not [96]. The authors noted that the amount of available data may have limited the statistical power of their analysis, however. For clean-contaminated wounds such as those commonly seen in patients with head and neck cancers, antibiotic prophylaxis becomes mandatory [97]. In cases in which the upper aerodigestive track is entered, the infection rate is much greater. In addition, these infections are often polymicrobial, and the chosen antibiotic should cover aerobic, anaerobic, and gram-negative flora [97]. From a tissue engineering perspective, local antibiotic release is an exciting avenue for exploration. A scaffold designed to promote bone growth, by material properties or the incorporation of bioactive molecules, could potentially deliver antibiotics to a local area, thereby alleviating some of the systemic effects of the antibiotic. Using an infected composite defect in a rabbit mandible, clindamycin-loaded poly(methyl methacrylate) space maintainers were shown to clear infection [98]. The study investigated both burst and extended release kinetics, but the inoculated bacteria were not recovered from any group. Several clinical products exist for local antibiotic release [99]. These clinically available products include gels, chips, fibers, and polymers consisting of a range of antimicrobials such as tetracycline, metronidazole, and doxycycline. After several days of controlled drug release to the site of the periodontal disease, degradation occurs for many of these products whereas others must be surgically removed [99]. Although the species distribution of craniofacial infection has not changed much over the years, the antibiotic resistance of these organisms is growing [100]. The production of b-lactamases has limited the efficacy of penicillin against some gram-negative species, and others are increasingly resistant to clindamycin. This growing resistance has led researchers to explore other potential methods of combatting infection [101,102]. Human saliva contains antimicrobial peptides that can cause bacterial cell death [102]. Although over 45 such peptides have been found,

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efficacy in human trials has been limited. Antimicrobial peptides can be designed with different functional groups and can be modified to bind hydroxyapatite selectively [101]. A modest minimum inhibitory concentration (125 mg/mL) was seen against Streptococcus mutans and Lactobacillus acidophilus when bound to hydroxyapatite. Although this particular antimicrobial peptide was designed to bind to tooth surfaces to combat biofilm formation, the in vitro studies were conducted on hydroxyapatite scaffolds, a major component of the CaP cements discussed earlier. Patient-Specific Technology Tissue engineering scaffolds have classically been fabricated from techniques such as electrospinning, phase separation, gas foaming, and particulate leaching [103]. Although researchers have worked to optimize and refine these techniques further, there are intrinsic limitations to the architecture and topography that can be achieved with these conventional techniques. Furthermore, techniques such as electrospinning require the use of organic carcinogenic solvents [104]. The field of tissue engineering has embraced AM or rapid prototyping techniques to bypass these limitations and produce scaffolds with superior reproducibility, more sophisticated structural features, and even patient-specific constructs [105]. Tissue engineers have leveraged AM to control the internal architecture of scaffolds, whereas surgeons have adopted AM to produce solid biomodels that reproduce the patient’s anatomy. The field of craniomaxillofacial surgery first adopted AM techniques to produce patient-specific models for surgical planning [106]. This proved to be an invaluable tool for surgeons, because an operation could be planned beforehand. In addition, it allowed the surgeon to predict physical outcomes of the procedure [106]. AM techniques have been extended to produce patient-specific implants. These custom implants precisely fit within a defect site, reducing surgical time and improving aesthetics [107,108]. To generate a patient-specific implant, a fine-cut CT study must be obtained that is acquired according to a special protocol. The radiographic data are first processed with software such as 3D Doctor, to creates a three-dimensional model of the defect. This model is then transferred to design software that allows the engineer to create the implant design. The implant is manufactured from this dataset by subtractive or AM techniques before sterilizing, packaging, and delivery to the surgeon. Initially, the fabrication of patient-specific implants was hindered by limitations within the fields of design and manufacturing. However, technological advances have led to customized cranial, dental, and facial implants in addition to space holders for grafts [58,109e111]. Chacon-Moya et al. published a case report of a 63-year-old female patient diagnosed with esthesioneuroblastoma [111]. The patient underwent an anterior craniofacial resection and radiotherapy. After treatment, the patient developed secondary complications of osteomyelitis and osteoradionecrosis that resulted in loss of the anterior cranial vault [111]. After resolution of the infection, the patient underwent frontal bone reconstruction using a computer-generated poly(ether ether ketone) (PEEK) implant. The patient was observed for 5 months after the PEEK implantation and no complications were observed. This case serves to demonstrate that reconstruction using PEEK implants is an excellent option in patients with large bone defects. Alternatively, computer-aided designebased technology can be leveraged to fabricate customized jigs and cutting guides (Fig. 50.11A and B). This allows the surgeon to create osteotomies in the fibula, resulting in an autogenous graft that approximates the defect contour (Fig. 50.12). Hou et al. used three-dimensional model simulation to contour vascularized fibular osteomyocutaneous flaps to repair mandibular defects in 15 patients [112] (Fig. 50.13). In this report, all patients experienced uneventful healing and were satisfied with the functionality and esthetics when questioned 6 months after surgery. These examples further illustrate the beneficial role AM has within the field of craniofacial surgery.

CONCLUSION The field of tissue engineering continues to make important strides toward the goal of producing biological tissue of high functional and esthetic fidelity to replace anatomical structures lost to trauma or disease. The field has developed sophisticated biomaterials that leverage manufacturing techniques, characterized and manufactured cell signaling proteins, and developed strategies to deliver therapeutic doses of antibiotics locally to improve tissue regeneration further. In addition, clinicians trying to apply this technology have gained new appreciation for the underlying cause of a defect, which has a significant role in the success of tissue regeneration. Although there have been significant advancements in the field, new directions of investigation into regenerating composite, vascularized, and innervated constructs are required to account for all the tissue types present within the craniofacial region.

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FIGURE 50.11 (A) Virtual surgical planning (VSP) matches the geometry of the fibula to the planned mandibular defect for reconstruction. (B) VSP plan shows final mandibular reconstruction using the fibula, which has been contoured to fit the defect through osteotomies and ostectomies.

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FIGURE 50.13

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Custom-fabricated plastic jigs applied to the fibula to guide the creation of osteotomies to contour the bone.

Fibula contoured with custom jigs is plated to maintain its shape before transplantation to the mandibular defect.

List of Abbreviations AM Additive manufacturing BMAC Bone marrow aspirate concentrate BMPs Bone morphogenetic proteins CaP Calcium phosphate CT Computed tomography MSC Mesenchymal stem cells PCL Poly(ε-caprolactone) PDGF Platelet-derived growth factor PLGA Poly(lactic-co-glycolic acid) rhBMP-2 Recombinant human bone morphogenetic protein-2 rhPDGF Recombinant human platelet-derived growth factor TCP Tricalcium phosphate TMJ Temporomandibular joint

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Acknowledgments We acknowledge support toward the development of new technologies for craniofacial regenerative medicine by the National Institutes of Health (NIH) (R34 DE025593) and the Army, Navy, NIH, Air Force, Veterans Administration, and Health Affairs to support the AFIRM II effort, under Award No. W81XWH-14-2-0004.

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C H A P T E R

51 Dental Tissue Engineering Nelson Monteiro1, Pamela C. Yelick2 1

University of Connecticut Health, Farmington, CT, United States; 2Tufts University School of Dental Medicine, Boston, MA, United States

INTRODUCTION Dental caries, trauma, genetic disorders, and periodontal diseases may cause damage and loss of dental tissues, and ultimately loss of the whole tooth [1]. Despite advancements in dental therapies, loss of dental tissues and teeth is a concern both to individuals and to their professional health care providers and remains a highly prominent public health issue [2,3]. It was reported that the public health burden associated with dental treatment of root canals represents a significant fraction of all expenses related to dental care, which totaled $100 billion in 2009 in the US alone [3]. Approximately 150 million adults experience tooth loss, and it is expected that over 10 million new cases of edentulism will have arisen in one decade [2]. Another highly significant population and of particular concern to the field of regenerative medicine and the US Department of Defense is the military. In 2000, a cost of $1.9 billion was estimated for active duty personnel and $203 million for recruits; periodontal disease accounted for the greatest proportion of active duty treatment costs (47%) and oral surgery for the greatest proportion of recruit treatment costs (32%) [4]. Therefore, the health of both military and civilian populations would significantly benefit from the development of new, improved, and alternative functional dental tissue replacement therapies. Treatments for deep carious lesions include pulp capping or partial pulp amputation to preserve the pulp tissue [5,6]. However, in the case of irreversible pulpitis, root canal treatment or extraction of the tooth is necessary. Root canal therapy involves complete removal of the infected pulp tissue and replacement with inert material [7]. There are some limitations associated with root canal treatment, including pulp tissue devitalization and excessive enlargement of the pulp chamber; in addition, debridement of infected root canals and the preparation of a post to support a synthetic tooth crown can increase the possibility of tooth root fracture and tooth loss. Dental prosthetic procedures such as dental implant placement have been used as tooth replacement therapies for many centuries [8]. Dental implants function through osseointegration, which is direct integration of the implant with the surrounding alveolar bone [9]. Dental implants have improved significantly over the years, based on innovations from basic and translational research, material sciences, and clinical techniques [8]. However, they are not equivalent to natural teeth in either function or aesthetics. They lack periodontal and cementum tissues, which function to cushion and modulate the mechanical stresses of mastication [9]. These disadvantages have driven an ongoing search for alternative strategies to overcome both the need for root canal and dental implant treatments. This search has led to the development of dental tissue engineering (TE) technologies, which consist of combined TE and clinical approaches to bioengineer dental tissues and whole teeth. Dental TE has demonstrated the potential use of dental stem cells (DSCs), biodegradable scaffolds, and bioactive agents such as growth factors (GFs) to control the spatial and temporal organization of regenerated dental tissues [8,10]. It is expected that dental TE will soon emerge as the preferred solution to root canal and dental implants therapies. Here, we summarize advances in dental TE to regenerate dental tissues including dental pulp, dentin, periodontal tissues, alveolar bone, and whole teeth.

Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00051-5

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TOOTH DEVELOPMENT Tooth development is a tightly regulated process mediated by dental epithelialemesenchymal cell interactions, Fig. 51.1. The dental mesenchyme and dental epithelium are derived from the neural crest and the ectoderm, respectively [11]. Tooth organogenesis is regulated by a complex and intricate network of cellecell signaling, gene expression, and GF signaling pathways [12]. Progress in elucidating the roles of molecular signals in natural tooth development has shed light on new approaches to control better the cell migration, growth, and differentiation of regenerated tissues [8,12,13]. It is known that at least 12 transcription factors are expressed in odontogenic mesenchyme [12], and more than 200 genes have been identified that are expressed in the oral epithelium, dental epithelium, and dental mesenchyme during the initiation of tooth development [14]. Tooth organogenesis is initiated by the formation of the dental lamina within the dental epithelium. The dental epithelial placode exhibits localized proliferative activity leading to dental epithelial outgrowths into the ectomesenchyme. The developing tooth organ proceeds through bud, cap, and bell stages [11,12]. In the bell stage, species-specific cusp patterns emerge, forming either a single or a multicusped tooth. The bell stage is followed by the differentiation stage, in which final growth and matrix secretion occur as the inner enamel epithelium differentiates into enamel-producing ameloblasts, whereas adjacent dental mesenchymal cells differentiate into dentin-producing odontoblasts [11,13]. The dental pulp is the soft connective tissue in the center of the tooth, which is enclosed by dentin [12]. Dental pulp has a variety of functions: (1) to support nerves that provide sensitivity to the tooth, (2) to nourish the avascular dentin, and (3) to produce the dentin that surrounds it. Dental pulp differs from other craniofacial structures in that it is extensively vascularized and innervated. Pulp vascularization is established by vasculogenesis during embryonic development of the tooth, and angiogenesis occurs during regeneration and therapeutic processes. Pulp innervation occurs at a relatively late postnatal stage and is innervated primarily by nociceptors [15]. Dentin is a resilient and elastic tissue that forms the bulk of the tooth, supports the enamel, and compensates for the brittleness of the highly mineralized enamel tissue. Dentin is a sensitive tissue that is capable of limited self-repair [12]. Odontoblasts and dental mesenchymal cells present in the tooth pulp can be stimulated to deposit new dentin in response to mechanical injury, termed reparative dentin. Enamel is the most highly mineralized tissue in the body; it consists of greater than 96% hydroxyapatite (HA) and exhibits a complex crystalline lattice organization [11,12]. Ameloblasts,

FIGURE 51.1 Principal stages of tooth formation [1].

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FIGURE 51.2 Adult human tooth morphology.

which secrete the enamel, largely undergo apoptosis before the tooth emerges into the oral cavity, with only a few dental epithelial cell remnants remaining in the periodontal ligament (PDL), termed the epithelial rests of Malassez. The tooth is attached to the jaw by specialized supporting tissues that consist of the PDL, the cementum, and the alveolar bone, all of which are derived from the neural crest cellederived dental mesenchyme and which are protected by the gingiva (Fig. 51.2). To date, vital vascularized and innervated human adult teeth cannot be regenerated or regrown once damaged or lost. Therefore, the development of effective therapies to regenerate and repair lost or damaged teeth is a major goal of dental TE and regenerative medicine.

DENTAL STEM CELLS DSCs can be collected from a variety of embryonic and postnatal or adult dental tissues. Embryonic stem cells (ESCs) are pluripotent stem cells that have the capacity to form all tissues of the body, including dental tissues [1,10]. ESCs can be obtained from the inner cell mass of a blastocyst in the 4- to 5-day-old embryo and from the embryonic germ in the 10- to 15-day-old embryo [10,16,17]. In contrast to ESCs, postnatal DSCs are multipotent stem cells that can develop into a restricted number of differentiated dental cell types [18]. The main functions of DSCs in natural teeth are to repair damaged dental tissues and maintain normal dental tissue turnover. Five different mesenchymal stem cell (MSC) populations have been identified in dental tissues, including dental pulp stem cells (DPSCs), stem cells from human exfoliated deciduous teeth (SHEDs), PDL stem cells (PDLSCs), dental follicle precursor cells (DFPCs), and stem cells from apical papilla (SCAPs). Each type of DSC population has specific characteristics and advantages for application in regenerative medicine and dentistry. In 2000, Gronthos et al. were the first to report the isolation and identification of DPSCs in adult human dental pulp [19]. They showed that DPSCs can produce densely calcified nodules in in vitro tissue culture, but did not exhibit the capacity to form adipocytes, compared with bone marrow mesenchymal stem cells (BMSCs), which do. However, in 2002, Gronthos et al. reported the capacity of DPSCs to differentiate into both adipocytes and neural-like cells [20]. Periodontal ligament stem cells (PDLSCs) are found in PDL tissues located between the tooth and the alveolar bone and can be isolated from the roots of extracted teeth. PDLSCs have the capacity to differentiate into cementoblast-like cells, adipocytes, osteoblasts and collagen-forming cells [21]. Moreover, they were shown to regenerate a cementumePDL-like structure and contribute to periodontal tissue repair [21]. SHEDs are isolated from an accessible tissue resource: autologous baby teeth [22]. They were identified as a population of highly proliferative, clonogenic cells capable of differentiating into a variety of cell types including neural cells,

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adipocytes, and odontoblasts [22]. SHEDs showed the capacity to induce bone formation, generate dentin, and survive in mouse brain after in vivo transplantation. DFPCs can be isolated from the dental follicle, which surrounds the developing tooth and supports the formation of cementum, PDL, and alveolar bone [23]. DFSCs have the capacity to differentiate into osteoblasts, adipocytes, and nerve-like cells in in vitro culture [24e26]. DFPCs expressed higher amounts of insulin-like growth factor-2 transcripts than human BMSCs, and after in vivo transplantation in immunocompromised mice, DFPCs expressed osteocalcin (OCN) and bone sialoprotein (BS) but did not exhibit cementum or bone formation [23]. However, other studies showed that DFSCs transplanted into mice contributed to the formation of a new PDL and cementum tissues [27,28]. SCAPs can be isolated from the dental papilla, which normally develop into the tooth dentin and pulp. It was reported that SCAPs proliferate two- to threefold more rapidly than stem cells obtained from the pulp organ [29]. Also, SCAPs were demonstrated to differentiate into osteoblasts, odondoblasts and adipocytes in in vitro culture, and when implanted in vivo, they were shown to differentiate into osteoblasts and odontoblasts [30,31]. Immunophenotypically, SCAPs are similar to DPSCs with respect to osteogenic or dentinogenic and GF receptor gene profiles, and they also express the stem cell marker STRO-1 and dentinogenic markers including bone sialophosphoprotein, OCN, and the GFs fibroblast growth factor (FGF) recombinant 1 (FGFR1) and transforming growth factor (TGF)-b [29]. Moreover, upon stimulation with a neurogenic medium, SCAPs express a wide variety of neurogenic markers including nestin and neurofilament M. Based on the promising characteristics of DSCs, dental TE efforts are focused on using DSC populations to bioengineer dental tissues, tooth supporting structures, and bioengineered whole teeth. Moreover, DSCs are considered to be a promising treatment method for several clinical conditions including Alzheimer disease, Parkinson disease, and spinal cord injury [32]. Here, we describe advances in dental TE to regenerate whole tooth and dental tissue structures, including dental pulp, dentin, PDL, and alveolar bone.

DENTAL TISSUE ENGINEERING The basic principle of TE employs the use of cells, scaffolds, and bioactive agents to regenerate tissues similar to native human tissues (Fig. 51.3). In vivo delivery exposes cells to a host of survival challenges, including immune issues related to inflammation and autoimmunity [33]. To survive the initial onslaught of the immune response

FIGURE 51.3

Dental tissue engineering approach [1].

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better and avoid death, in vivoedelivered cells can be seeded in or onto a biodegradable scaffold before implantation when they are used to repopulate a tissue defect and/or restore function [34]. Ideal TE scaffolds mimic the properties and functions of the natural extracellular matrix (ECM), providing appropriate physical characteristics to support the development of new functional tissues. Therefore, it is important to consider the mechanical integrity and surface functionality of any given biomaterial scaffold, to ensure appropriate cell adhesion, proliferation, and differentiation [35,36]. Both natural and synthetic biomaterial scaffolds have been proposed for dental tissue regeneration [1,35,37e40]. The advantage of using natural scaffolds (i.e., collagen, alginate, fibrin, chitosan, gelatin, silk) is the ability to fine-tune their degradation rates by varying the concentration of the polymer and/or cross-linking agents. Synthetic scaffolds (polylactic acid, polyglycolic acid (PGA), PLLA, PLGA, PCL, etc.) have advantages including processing flexibility and the ability to be manufactured in any shape or size and with the desired predefined architecture and structural parameters [35]. Hydrogels are a specific class of scaffolds that exhibit huge potential for applications in dental tissue regeneration, owing to their versatility and adaptability [41,42]. Hydrogels offer several advantages including injectability, easy incorporation of therapeutic agents and cells under mild conditions, minimally invasive local delivery, and high contourability [42,43]. Bioactive agents such as GFs and nucleic acids have an important role in odontogenesis [12] and can be used to functionalize scaffolds to initiate the formation of new tissues [1]. As such, scaffolds can provide a multitude of advantages for bioactive agents, including safe delivery profiles, protection from biodegradation, and the ability to deliver the bioactive agents locally to where they are needed [13,44e46]. GF are critical to the development, maturation, maintenance, and repair of craniofacial and dental tissues [12]. It is therefore important to understand which of the GFs resident in the DSC niche provide appropriate cues to control their fate. For instance, GFs produced by DSCs during tooth development are responsible for the regenerative capacity of the dentin, which maintain cell proliferation and differentiation potential [12]. GFs function by binding to the extracellular domain of an appropriate target GF receptor, which in turn activates intracellular signal transduction pathways [47]. Several GFs (i.e., bone morphogenetic proteins [BMPs], sonic hedgehog [SHH], TGFs, FGFs, brain-derived neurotrophic factor [BDNF], and VEGF) are expressed during tooth formation and repair [1,15]. Therefore, dental tissue regeneration may be facilitated by incorporating GFs into scaffolds to promote dental cell differentiation. Also, because GF protein expression is the result of their gene expression [12], incorporating nucleic acid expression constructs into scaffolds has been proposed as a means to overcome certain limitations of GF delivery, including their short half-lives, denaturation during encapsulation processes, the time-consuming and expensive problem of GF production, the requirement for supraphysiological doses and GF combination for the most effective approach, the extended times required for cell differentiation, and difficulties in differentiating cells toward specific lineages [47e49]. Gene delivery of transcription factors can be used to ensure proper expression of particular splice variants in a coordinated time and sequence, and the ability to regulate a cascade of multiple genes, all from a single delivered construct [1]. Moreover, interference RNA (RNAi), a gene silencing mechanism, can be used to induce DSC differentiation [50]. Several approaches have been developed to incorporate a variety of bioactive agents into scaffolds and control their release profiles [45,46]. For instance, covalent immobilization of bioactive agents offers additional control over the spatiotemporal distribution of a particular agent compared with physical adsorption [1]. Immobilized bioactive agents can be released upon degradation of the matrix or by the hydrolysis of degradable links. Another way to entrap bioactive agents into scaffolds is by incorporating microparticles and nanoparticles [51e54]. The nanoscale properties of nanoparticles enable the ability to fine-tune release kinetics for improved transport properties, diffusivity, solubility, regulated biodistribution, minimization of toxic side effects, and the enhanced therapeutic index of bioactive agents. The combination of DSCs, biomaterial scaffolds, and bioactive agent delivery systems allows for the creation of multifunctionalized systems that can be used to facilitate dental tissue regeneration [1]. Notably, decellularized scaffolds and scaffold free approaches have created new areas of research in addition to the previously used natural and synthetic scaffold approaches [55e58]. Decellularized scaffolds, created by gently removing immunogenic cells from natural tissues such as heart, lung, bone, liver, and tooth buds, preserve the structure, shape compatibility, mechanical integrity, and bioactive molecule gradients that facilitate cellecell interactions, cell adhesion, and ECM formation [55,59]. As such, detailed characterizations of ECM composition and organization in natural dental tissues could also facilitate dental TE efforts [56,60]. Finally, cell sheet technologies have been proposed to facilitate the regeneration of dental tissues including tooth root, pulp, dentin, and periodontal tissues [57,61e63]. This technique enables the creation of intact sheets of cells that can be harvested without trypsinization and preserves the ECM formed by the cells, including adhesive proteins such as fibronectin. Application of these TE approaches in dental tissue regeneration are discussed subsequently.

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Whole Tooth Engineering Several approaches to engineering entire biological teeth have been proposed, including dental TE, stimulation of third dentition formation, celletissue recombination, chimeric tooth TE, and gene-manipulated tooth regeneration [1]. The two major approaches used for tooth regeneration are celletissue recombination and dental TE approaches. Dental celletissue recombination approaches rely on replicating the natural processes of tooth development, in which cultured progenitor stem celletissue constructs are directly implanted in the defect site (Fig. 51.4). Many studies have reported the bioengineering of functional teeth from embryonic stem cells cultured in vitro and/or implanted in vivo [16,17,64e66]. Ohazama et al. showed that embryonic day (E)10 oral epithelium stimulated an odontogenic response in cultured neural and BMSCs [16]. Moreover, BMSC-derived recombinants were demonstrated to form tooth crown structures composed of enamel, dentin, and pulp. In contrast, tissue recombinants derived from ESCs and neural stem cells did not form teeth but expressed odontogenesis-related genes [16]. In another study, high celledensity suspensions generated from E14.5 incisor epithelial and mesenchymal tissues cultured in vitro or implanted into the subrenal kidney capsule subsequently formed bioengineered teeth when implanted in the extraction socket of a rat mandibular incisor [17]. These studies demonstrate the potential for using ESCs for tooth regeneration. However, the potential tumorigenic nature of transplanted ESCs, ethical issues regarding the use of human embryos, and the potential for allogeneic immune rejection are all major concerns of ESC-based therapies [10,67]. Adult DSCs, BMSCs, umbilical cordederived MSCs, and induced pluripotent stem (iPS) all exhibit potential for use in dental TE [10,68]. Young et al. demonstrated the first successful use of adult DSCs seeded onto PGAePLLA scaffolds to generate tooth crowns containing dentin and enamel; this demonstrated the presence of adult dental epithelial and MSC populations in pig third-molar tooth tissues [69]. Another study demonstrated that bioengineered pig tooth crowns containing dentin, pulp, and enamel formed in 25e30 weeks, whereas rat tooth crowns formed in just 12 weeks [70]. Honda et al. harvested dental cells from canine first-molar tooth buds, seeded them onto PGA scaffolds, and implanted them into tooth sockets of extracted teeth [71]. Subsequent analyses of harvested implants showed the formation of dentin and bone, but no enamel tissue or dental-root formation was observed. The same authors reported that dental epithelial and mesenchymal cells harvested from porcine third molar teeth seeded onto collagen scaffolds and implanted in vivo formed a single tooth in each scaffold that was morphologically similar to natural teeth [72]. In another interesting study, comparison of bioengineered dental tissues grown in the mandible versus the omentum revealed that both implant sites supported the formation of bioengineered dentin, enamel, pulp, and periodontal tissues [73]. However, omental implant dental tissues appeared to be more organized than those grown in the mandible. Yang et al. reported the formation of a complete tooth, consisting of a tooth crown, root, pulp, enamel, dentin, odontoblasts, cementum, blood vessels, and PDLs when tooth bud cells were suspended in fibrin glue and autografted back into the original alveolar sockets of a pig [74]. Smith et al. demonstrated the use of gelatin methacryloyl (GelMA) hydrogels, combined with postnatal porcine dental epithelial and dental mesenchymal progenitor cells, to support the formation of mineralized and functionally vascularized tissues of specified size and shape [75]. Also, the same model was improved by incorporating dental epithelial and dental mesenchymal cell sheets that expressed appropriate tooth marker expression patterns including SHH, BMP-2, Runt-related transcription factor 2, tenascin, and syndecan in vitro and in vivo [57]. Yang et al. reported whole tooth regeneration using DPSCs combined with epithelial cells isolated from gingival epithelium implanted into the mandibular alveolar socket of a pig for 13.5 months [66]. Seven of eight pigs developed two teeth containing crown, root, and pulp structures. Subsequent histological analyses demonstrated the formation of enamel-like tissues, dentin, cementum, odontoblasts, and periodontal tissues. All of the pig hosts formed regenerated molar teeth regardless

FIGURE 51.4 Dental cell tissue recombination approach [1].

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of the original tooth type used to procure the DPSCs, whereas pig hosts that had tooth germs removed or received acellular scaffold implants did not develop new teeth. The authors also reported that the implant location may have influenced the morphology of the regenerated tooth. More detailed characterizations of tooth expressed ECM molecules, including their respective developmental and spatial organization, may facilitate the design of effective scaffolds for tooth regeneration. One approach to achieving this goal included devising methods to decellularize and demineralize porcine molar tooth buds effectively while preserving natural ECM protein gradients [56]. This report showed that the collagen I, fibronectin, collagen IV, and laminin gradients present in natural tooth tissues were retained in decellularized tooth bud samples. Second harmonicegeneration image analysis and three-dimensional (3D) reconstructions showed that natural tooth tissue exhibited higher collagen fiber density and more organized collagen fibers compared with decellularized tooth tissue. That report showed that dental cells seeded back into the decellularized tooth bud scaffolds were able to establish residence within the scaffold and to elaborate and remodel the matrix. Ongoing research has focused on detailed 3D characterizations of tooth pulp tissue to examine collagen fiber destruction and remodeling as a consequence of the decellularization and reseeding processes.

Dental Pulp and Dentin Regeneration Dental pulp consists of DPSCs, odontoblasts, endothelial cells, neurons, immune system cells, GFs and the ECM, all of which are crucial for maintaining the functions of healthy teeth. Dentin, which encases the dental pulp, is a mineralized form of the collagen-based predentin matrix; its crystalline structure primarily consists of HA and water [76]. Creating an in vitro tissue model that takes into account all of the aspects of a pulpedentin microenvironment is challenging. Nonetheless, adopting a deconstruction strategy to reduce the pulpedentin ecosystem to a few of the main components anticipated to be involved in maintaining functional pulpedentin biology may be sufficient to establish in vitro models capable of pulpedentin regeneration. Because interactions between tissue vasculature and nerves is critical to maintaining dental pulp homeostasis, dental TE strategies for pulp regeneration must recreate a microenvironment that supports the cellular cross-talk needed to maintain the healthy, functional tissue heterogeneity of natural tooth pulp [76]. Advances in the field of TE have made dental pulp regeneration a realistic, attractive, and alternative therapy for endodontic treatments to restore damaged teeth [5,6,77e81]. The goal is to replace the damaged pulpedentin complex with a bioengineered, functional biological tissue surrogate that can integrate with remaining healthy host tissues. The first evidence of novo pulp regeneration was demonstrated using human tooth slices or root segments (RSs) injected with DSCs combined with scaffolds [79]. In that report, Cordeiro et al. showed that SHED cell-seeded PLLA scaffolds placed within human tooth slices and transplanted into immunodeficient mice created bioengineered tissue architecture and cellularity closely resembling those of physiologic dental pulp [79,81], including SHED cell differentiation into odontoblast and endothelial-like cells. When supplemented with VEGF, SHED cells expressed VEGF receptor2, CD31, and vascular endothelial cadherin and organized into capillary-like sprouts [81]. These studies demonstrated that the VEGF signaling pathway is an important regulator of endothelial cellecontrolled differentiation of DSCs, and therefore that scaffold biomaterials could be used to control the delivery of GFs such as VEGF for dental TE [13]. Another study consisted of SCAPs and DPSCs seeded onto PLG scaffolds, inserted into tooth RSs and transplanted in mice. Analyses of harvested implants showed that the root canal space was filled entirely by a pulp-like tissue exhibiting well-established vascularity, and the formation of a continuous layer of dentin-like tissue deposited along the dental wall of the tooth root canal. The newly formed dentin like tissue was elaborated by a layer of newly formed odontoblast-like cells expressing dentin sialophosphoprotein, BS, alkaline phosphatase, and CD105. For optimized clinical translation, dental pulp regeneration will require the use of injectable scaffolds. Rosa et al. tested the hypothesis whether SHED encapsulated in PuraMatrix hydrogel or collagen I, could regenerate a functional dental pulp when injected into full-length root canals and implanted in vivo [77]. The authors reported the formation of pulp-like tissues, including the presence of odontoblasts capable of generating new tubular dentin throughout the root canals. The engineered pulp tissue exhibited cellularity and vascularization similar to those of natural human dental pulps. Therefore, this strategy might successfully facilitate the completion of tooth root formation in damaged, necrotic, immature permanent teeth. In another study, autologous DPSCs encapsulated in collagen were transplanted into a root canal along with stromal cellederived factor-1 (SDF-1) after pulpectomy of mature teeth with complete apical closure in dogs [82]. The results showed that by day 14, the root canal was successfully filled with regenerated pulp tissue, including nerves and vasculature, followed by new dentin formation along the dentinal wall. Khayat et al. also defined a reliable method to regenerate pulp-like tissues within tooth RSs using 5% GelMA hydrogel-encapsulated human DPSCsehuman umbilical vein endothelial cells

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(HUVECs) implanted subcutaneously in nude rats for 4 and 8 weeks [6]. Analyses of harvested implants showed that hDPSCeHUVEC-encapsulated GelMA constructs formed pulp-like tissue that adhered to the inner dentin surface of the RS, including odontoblast-like cells exhibiting cell extensions into the dentin tubules. Moreover, that report showed that GelMA hydrogels supported hDPSCeHUVEC cell attachment, proliferation, and infiltration of host cells and promoted the establishment of well-organized neovasculature formation. This study identified GelMA hydrogels combined with DSCs as a promising, clinically relevant pulpal revascularization treatment to regenerate human dental pulp tissues. Another study showed that SHEDs and DPSCs cultured in peptidee amphiphile hydrogel scaffolds exhibited cell proliferation and differentiation within the scaffolds, and that SHEDs formed a soft tissue whereas DPSCs deposited mineral [43]. These hydrogel scaffolds are desirable for clinical applications for pulp regeneration, because they are easy to handle and can be introduced into small defects or root canal. Hyaluronic acid sponge scaffolds also exhibit an appropriate structure, biocompatibility, and biodegradation for dental pulp regeneration [83]. In vivo studies using hyaluronic acidebased scaffolds in an amputated dental pulp of rat molar showed dental pulp proliferation and blood vessel invasion. Prescott et al. investigated the role of DPSCs, collagen scaffold, and dentin matrix protein 1 (DMP1) in a simulated furcal perforation, in vivo mouse model [84]. The successful formation of organized pulp tissue was observed in the group containing the triad of DPSCs, a collagen scaffold, and DMP1. In another report, a chitosan bilayer membrane containing TGFb1 releasing microspheres was developed to promote reparative dentin formation in a pulp-capping dog model [85]. Analyses of in vivo implanted constructs showed that TGF-b1ereleasing chitosan membranes generated reparative dentin three to six times thicker than those generated by chitosan bilayer membrane alone. In another study, decellularized human dental pulp ECM supported the proliferation and differentiation of SCAP into odontoblastlike cells near pulp chamber dentinal walls [86]. That study demonstrated that human dental pulp from healthy extracted teeth can be successfully decellularized and subsequently used to differentiate reseeded DSCs, potentially improving clinical outcomes and ultimately promoting the survival and function of injured teeth. Ravindran et al. also demonstrated the odontogenic differentiation of both human DPSCs and human PDLSCs when cultured on a decellularized 3D pulp ECM scaffold, without the need for the exogenous addition of GFs. Subcutaneous implantation of ECM scaffolds containing DPSCs showed the formation of dental pulp-like tissue containing cells expressing DSP and dentin phosphophorin . These results showed that a decellularized dental pulp ECM scaffold can be used as a biomimetic scaffold for dental tissue regeneration and as tool to study the extracellular function of multifunctional proteins [60]. Kim et al. reported the regeneration of dental pulp-like tissue in human teeth by cell homing and without cell transplantation [80]. Human teeth implanted in the mouse dorsum and delivering basic FGF (bFGF) and/or VEGF formed recellularized and revascularized connective tissue that integrated with native dentinal walls of tooth root canals. Also, the combined delivery of bFGF and VEGF, or platelet-derived growth factor (PDGF) combined with a basal set of nerve growth factor and BMP-7, generated cellularized and vascularized tissues and neodentin formation over the surface of native dentinal walls. In another study, scaffold-free 3D tissues were engineered from DPSC sheets, placed into the canal space of human tooth RSs, and implanted subcutaneously into mice for pulp regeneration [87]. Histological results indicated that after 3e5 months, implanted tooth roots containing 3D scaffold-free engineered tissues exhibited vascularized fibrous tissue formation throughout, whereas empty tooth roots remained predominantly empty.

Periodontal Regeneration Periodontal tissue regeneration strategies aim to restore the supporting periodontal tissues of teeth, including the formation of new cementum, PDL, and alveolar bone [88]. Clinical studies show that transplantation of autologous PDL progenitor (PDLP) cells may be a useful therapy to repair periodontal defects [89]. It was shown that PDLPs were similar to PDLSCs with respect to their ability to exhibit high proliferation rates, express mesenchymal surface molecules, and regenerate in vivo tissue; this provides clinical and experimental evidence supporting the efficacy and safety of using autologous PDL cells in treating human periodontitis. In another study, human PDLSCs were transplanted into immunocompromised mice and rats to assess their capacity for tissue regeneration and periodontal repair [21]. Using defined in vitro culture conditions, PDLSCs were shown to differentiate into cementoblast-like cells, adipocytes, and collagen-forming cells. When transplanted into immunocompromised rodents, PDLSCs exhibited the capacity to generate cementumePDL-like tissues and contribute to periodontal tissue repair. Therefore, transplantation of PDLSCs may hold promise as a therapeutic approach for reconstructing tissues destroyed by periodontal diseases. It was also reported that autologous PDLSCs, combined

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with HAeb-TCP scaffolds and transplanted into surgically created periodontal defects in pigs, were capable of regenerating periodontal tissues, leading to a favorable treatment for periodontitis [90]. Using a similar approach, it was shown that transplanted human SCAPs and PDLSCs, combined with HAeTCP ceramic particles and implanted in an in vivo swine model, regenerated a functional tooth root and periodontal complex capable of supporting a porcelain crown [91]. As such, this approach led to the recovery of both tooth strength and appearance. Zhou et al. reported a PDL cell-sheet delivery system to promote periodontal tissue healing in a canine model [62]. After root canal treatment, the tooth roots were wrapped with PDL fibroblast-derived cell sheets and replanted back into the same tooth socket for 8 weeks. Subsequent analyses of the in vivo implanted teeth showed that multilayered PDL cell sheet constructs attached to the tooth root and that most cells of the PDL sheet-tooth constructs were viable after replantation. The PDL cell-sheet group showed a significantly higher occurrence of healing (88.4%) than the control group without cell sheets (5.3%). PDL and cementum tissue regeneration was observed in the experimental group and the regenerated tissues expressed high levels of collagen type III, collagen type I, and fibronectin expression. In a similar study, three-layered PDL cell sheets supported by PGA nanofibers were transplanted to dental root surfaces exhibiting a three-wall periodontal defect while filling existing bone defects with porous b-TCP. PDL cell sheet transplantation resulted in regenerated new bone formation as well as the formation of cementum containing embedded, properly oriented collagen fibers. These results suggested that PDL cells exhibit the ability to differentiate into periodontal tissues composed of both hard and soft PDL tissues, and that PDL cell sheet transplantation may be useful for periodontal regeneration in clinical settings [92]. In another study, Lei et al. characterized the cell properties of DPSCs and PDLSCs after in vivo implantation [93], showing that DPSCs and PDLSCs can maintain MSC-like characteristics after in vivo implantation compared with PDLSCs, DPSCs appear to be much more stable under in vivo conditions. These findings provide additional cellular evidence supporting and expanding the use of dental tissueederived stem cells in dental TE. In another study, intrabony defects were created in rats to evaluate the regenerative potential of an injectable macroporous calcium phosphate cement (CaP) combined with BMP-2 or FGF-2. Results from that study showed that the combined topical application of FGF-2 with an injectable CaP may prove to be a promising treatment for periodontal regeneration [94]. Another study, by Kim et al., tested the hypothesis that anatomically correct teeth can be regenerated in scaffolds without the need for cell transplantation. Anatomically shaped human molar and rat incisor scaffolds were fabricated by 3D bioprinting of hybrid PCLeHA with 200-mm-diameter interconnecting microchannels. SDF1 and BMP-7 were delivered into the microchannels of the scaffold, which were then implanted orthotopically after mandibular incisor extraction in the case of rats, or by using human molar-shaped scaffolds implanted ectopically into the rat dorsum. Analyses of explanted constructs showed the regeneration of PDL and new bone, which formed at the interface of the rat incisor scaffold and the native alveolar bone. This study showed that SDF1 and BMP-7 delivery recruited significantly more endogenous cells and also induced a greater angiogenesis response compared with growth factorefree control scaffolds [95]. Porous chitosanecoral composites combined with plasmid-encoding PDGF-B gene were also tested for periodontal regeneration, showing that subcutaneous implantation of gene-activated scaffolds supported greater cell proliferation than pure coral scaffolds, and that PDGFB and type-I collagen expression was upregulated in gene-activated scaffold. Therefore, coral scaffolds combined with PDGF-B gene delivery may also serve as a suitable approach for periodontal tissue regeneration [96]. In another study, a mesoporous bioglassesilk fibrin scaffold combined with BMP-7 and/or PDGF-B adenovirus synergistically promoted up to twofold greater regeneration of PDL, alveolar bone, and cementum compared with each adenovirus used alone [97]. Ex vivo BMP-7 gene transfer was used to stimulate the repair of large mandibular alveolar bone defects in a rat wound repair model [98], consisting of syngeneic dermal fibroblasts transduced ex vivo with adenoviruses encoding green fluorescent protein, BMP-7 (Ad-BMP-7), or an antagonist of BMP bioactivity, noggin (Ad-noggin). These studies demonstrated that Ad-noggin treatment inhibited osteogenesis compared with the control- and Ad-BMP-7etreated specimens. Furthermore, the osseous lesions treated with Ad-BMP-7 gene delivery demonstrated rapid chondrogenesis and subsequent osteogenesis, cementogenesis, and predictable bridging of the periodontal bone defects [98]. In another study, a thermosensitive chitosan hydrogel was used as a small interfering RNA (siRNA) reservoir to silence receptor activator of nuclear factor-kB signaling and promote PDL regeneration [50]. The cumulative in vitro release of siRNA from the hydrogel was 50% over 14 days, and high PDL cell viability was observed for cells seeded on the siRNA-loaded scaffold. In vivo studies showed that the fluorescent signal from siRNA within hydrogel was maintained for up to 14 days when subcutaneously implanted in mice [50].

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Alveolar Bone Regeneration Jaw reconstruction can be challenging because both teeth and bone must be restored [99]. Autologous bone graft techniques followed by the placement of dental implants is one approaches being used to repair jaw defects. However, undesirable limitations to this approach include associated donor-site morbidity for harvested bone, insufficient quantities of available bone, and difficulties in dental implant placement owing to inadequate size, shape, and quality of the reconstructed alveolar ridge [100,101]. Because of high rates of progressive periodontitis, which can alter alveolar bone morphology and destroy surrounding tooth-supporting tissues, alveolar bone is highly susceptible to inflammation, which may result in necessary tooth extraction. The alveolar ridge may continue to resorb even after dental implant placement, and dysregulated bone remodeling in response to mechanical loading (mastication forces) may occur owing to uneven strain distributions caused by resorbed alveolar bone tissues [101]. Therefore, reliable tooth and alveolar bone regeneration strategies are needed to repair jaw defects effectively. Inorganic materials such as CaP ceramics, bioactive glass, and ceramicepolymer composites have been especially developed for bone TE applications [39,101]. HA ceramics, b-TCP cements, and biphasic calcium phosphates are examples of synthetic CaP bone substitutes [101]. Although they lack the mechanical properties of naturally formed bone, they exhibit osteoinductive or osteogenic abilities, and these ceramics gradually acquire mechanical strength similar to that of cancellous bone [101,102]. For alveolar bone regeneration, biodegradable granular forms of ceramics are preferred, such as b-TCP, because they are easy to shape and can easily adopt the 3D size and shape of the bony defect, an important consideration for proper aesthetics [102]. Advantages of combining ceramics with polymers include improved biodegradability, biocompatibility, and an ability to readily bind GFs critical for osteoinduction. PCL scaffolds coated with CaP or HA exhibit improved osteoblast adhesion, spreading, and proliferation and promote alveolar bone formation in periodontal defects [101]. Goh et al. used a monkey model to examine periimplant bone regeneration and implant stability after immediate implant placement into tooth sockets containing facial wall defects [103]. In the control group, the bony defect was reconstructed with autogenous particulate bone, whereas in the test group, PCLeTCP scaffolds were used. Histological analyses showed better maintenance of facial bone contour in the test group; however, bone regeneration was observed only in areas adjacent to the bony wall of the defect. Implant survival was 100% 6 months for both groups, but the use of a PCLeTCP scaffold showed better maintenance of the alveolar contour compared with autogenous particulate bone at 6 months, although there was minimal bone regeneration within the defect [103]. Aquino et al. reported the use of autologous human DSPCs combined with a collagen sponge scaffold for oral maxillofacial bone tissue repair in patients requiring third-molar extraction [104]. Three months after autologous DSPC grafting, they reported that patient alveolar bone showed optimal vertical repair and complete restoration of periodontal tissue. Moreover, optimal bone regeneration was still evident 1 year after grafting. This clinical study demonstrated the effectiveness of DSCs combined with scaffolds to restore dental tissues such as alveolar bone and periodontal tissues [104]. In another study, bone implants created from pig iliac crestederived MSCs seeded onto lattice scaffolds and grown in a rotational oxygen-permeable bioreactor system (ROBS) reactor were then combined with dental epithelial and dental mesenchymal cellederived tooth constructs created from unerupted tooth buds harvested from the same pig [105]. Analyses of harvested implants revealed the formation of bioengineered mineralized tooth tissues, including primary and reparative dentin and enamel in the tooth portion of the hybrid toothe bone implants, and the formation of OCN and BS-positive mineralized bone in the bone portion of the hybrid toothe bone constructs. Collagen type IIIepositive connective tissue resembling PDL and tooth root structures were also present at the interface of the bioengineered tooth and bone tissues. These results demonstrate the potential use of hybrid toothealveolar bone constructs for clinical treatment of tooth loss accompanied by alveolar bone resorption [105]. In a similar study, toothebone constructs were prepared from third-molar tooth tissue and iliac crest bone marrowederived osteoblasts isolated from, and implanted back into, the same pig as an autologous reconstruction [100]. Subsequent analyses showed the formation of small tooth structures consisting of organized dentin, enamel, pulp, cementum, and PDL surrounded by newly formed bone, indicating the possibility of tooth regeneration and associated alveolar bone in a single procedure [106]. Cai et al. studied which differentiation approach (e.g., maintenance of stemness, osteogenic or chondrogenic induction) was most suitable for periodontal regeneration, using in vivo implanted, rat BMSC-seeded PLGAePCL electrospun scaffolds. Their results showed that BMSCs exhibited chondrogenic differentiation, followed by regeneration of alveolar bone and ligament tissues [107]. Successful regeneration of alveolar bone and surrounding

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periodontal tissues using gene therapy vectors such as adenoviral BMP-7 (Ad-BMP-7) was also achieved in vivo [98]. Mesoporous Bioglassesilk fibrin scaffolds combined with BMP-7 and/or PDGF-B adenovirus synergistically promoted periodontal regeneration by enabling up to twofold greater regeneration of PDL, alveolar bone, and cementum tissues compared with each adenovirus used alone [97].

CONCLUSIONS Root canal therapy and prosthetic dental procedures such as dental implants are the only clinical treatments available to treat necrotic dental pulp tissue defects and tooth replacement. These treatments present numerous disadvantages, including loss of tooth vitality and the inability to mimic properties and functions of natural teeth. Based on the promising characteristics of DSCs for regenerative tissue applications, dental TE efforts are focused on using these stem cell populations to bioengineer tooth supporting tissues and whole tooth. Although a variety of studies support the potential use of ESCs for tooth regeneration, the possibility of malignant tumorigenesis, ethical issues regarding the use of embryos, and the potential for allogeneic immune rejection make this approach problematic and unrealistic at present. In contrast, adult DSCs and iPS cells both exhibit significant potential for utility in dental TE. Successful use of DSCs for applications in dental tissue regeneration requires the combination of appropriate scaffolds and GF signals to induce dental cell differentiation and dental tissue formation. In addition, the concepts of scaffold-free approaches and cell homing have also been developed, suggesting alternative approaches using dental cell sheets and the recruitment of endogenous MSCs for dental tissue by GFs. These reports demonstrate that it is possible to regenerate correct teeth anatomically using a combination of dental cells, scaffolds, and signals. Indeed, the tooth organ is a highly complex biological organ whose formation requires the intricate regulation of a cascade of molecular signals and gene expression. As such, a better understanding of the biological processes and interactions of important GFs and gene expression patterns regulating natural tooth development is critical for successful dental tissue regeneration. Methods to differentiate DSCs successfully into tooth specific lineages, to create all of the tissues that comprise a functional living tooth, is a main challenge in DSC research. New and effective methods for delivering bioactive agents such as GFs and nucleotides (i.e., plasmid DNA and RNA interference) combined with optimized biomaterials scaffolds, are anticipated to provide the means for successful alveolar bone and tooth regeneration in the foreseeable future, to meet the challenge of generating functional, living, bioengineered dental tissues.

List of Abbreviations BMP Bone morphogenetic protein BS Bone sialoprotein CaP Calcium phosphate cement DFPC Dental follicle precursor cells DPSC Dental pulp stem cells ECM Extracellular matrix ESC Embryonic stem cells FGF Fibroblast growth factor GF Growth factors HA Hydroxyapatite OCN Osteocalcin PCL Polycaprolactone PDGF Platelet-derived growth factor PDL Periodontal ligament PDLSC Periodontal ligament stem cells PGA Polyglycolic acid PLGA Poly(lactic-co-glycolic acid) PLLA Poly-L-lactide SCAP Stem cells from apical papilla SHED Stem cells from human exfoliated deciduous teeth TCP Tricalcium phosphate TGF Transforming growth factors VEGF Vascular endothelial growth factor

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Acknowledgments This study was supported by the National Institutes of Health/National Institute of Dental and Craniofacial Research (R01DE016132, PCY and AFIRM2 CF-04, PCY).

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C H A P T E R

52 Cell Therapy for Blood Substitutes Shi-Jiang Lu1, Robert Lanza2 1

Vcanbio Center for Translational Biotechnology, Natick, MA, United States; 2Astellas Institute of Regenerative Medicine, Marlborough, MA, United States

INTRODUCTION RBCs, the oxygen-carrying component of the blood, are transfused in over half of all anemic patients admitted to intensive care units in the United States [1e3]. It is estimated that nearly 5 million patients receive approximately 14 million units of RBCs per year in the United States alone [4,5]. Limitations in the supply of RBCs can have potentially life-threatening consequences for patients, specifically for those who have rare or unusual blood types with massive blood loss caused by trauma or other emergency situations. Unfortunately, the supply of transfusable RBCs, especially the “universal” donor type (O)Rh-negative, is often insufficient, particularly in the battlefield environment and/or major natural disasters owing to the lack of blood type information and the limited time required for lifesaving transfusion. Moreover, the low prevalence of (O)Rh-negative blood type in the general population ( 433 mmol/L2/h . week1 (90th percentile), or a composite with the HYPO score greater than 423 (75th percentile) and LI greater than 329 (75th percentile) [51]. Because patients with poor diabetes compliance or an inadequate baseline insulin regimen are likely to benefit from an improved design of insulin dosing regimens, patients selected for transplant should have a plasma HbA1C less than 10%. In an effort to reduce the risk of serious procedural and immunosuppressive drug-related complications, the patient’s cardiac and renal function should be carefully assessed. Selected recipients should have adequate cardiac function including blood pressure less than 160/100 mmHg, no evidence of myocardial infarction in the 6 months before assessment, no angiographic evidence of noncorrectable coronary artery disease and left ventricular ejection fraction greater than 30% as measured by echocardiogram. To eliminate patients who are better candidates for simultaneous kidneye pancreas transplantation or those who may experience adverse renal function as a result of tacrolimus, selected recipients should have no evidence of macroscopic proteinuria (70 in females) mL/min/1.73 m2. Proliferative retinopathy should be stabilized before transplantation, because acute correction of glycemic control may lead to accelerated retinopathy. Finally, to reduce

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the risk of antibody-mediated graft rejection, potential recipients should be screened for panel reactive antibody (PRA) assays. It remains unknown whether patients with elevated initial PRA will have worse long-term islet function if they have a negative cross-match to the specific donor cells.

Islet Transplantation Procedure Although several locations have been tested as potential implantation sites for islet grafts, the high level of graft function and ease of delivery associated with infusion into the portal circulation of the liver have led to this being the transplantation site of choice in clinical protocols [27]. There are two accepted approaches for implanting purified islets into the liver by way of the portal vein. Whereas surgical laparotomy and cannulation of the portal vein was most often used in the early islet transplant programs, current protocols routinely employ the percutaneous transhepatic approach to implant donor islets in cadaveric islet transplantation (Fig. 56.2A) [51]. Compared with surgical laparotomy, this procedure is minimally invasive and thus can be performed using local anesthesia combined with opiate analgesia and hypnotics given as premedication. Access to the portal vein is achieved by percutaneous transhepatic approach using a combination of ultrasound and fluoroscopy to guide the radiologist. A branch of the right portal vein is cannulated and a catheter is positioned proximal to the confluence of the portal vein, which is confirmed with a portal venogram [52]. The risk for portal vein thrombosis is reduced by inclusion of unfractionated heparin (70 U/kg) in the islet preparation. Islets are then infused aseptically into the main portal vein under gravity, with regular monitoring of portal venous pressure (by an indirect pressure transducer) before, during, and after the infusion. An ultrasound examination should be performed at 1 day and 1 week after transplant to rule out intraperitoneal hemorrhage and confirm that the portal vein is patent and has normal flow.

FIGURE 56.2 The islet transplant procedure: present and future. Islet transplantation, in its current form (A), has provided insulin independence in most diabetic patients at 1 year after transplant, but this procedure is limited by the availability of suitable cadaveric donors and the requirement for lifelong immunosuppression. In the future (B), islet transplantation could be made available to a broader range of diabetic patients through the use of alternative tissue sources such as living donors, xenogeneic donors, or stem cellederived b cells. Also, as novel immunomodulatory therapies are identified, tolerance induction strategies can be developed that will prolong graft function and allow for the reduction or complete withdrawal of immunosuppressive drug therapy.

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If the skills of a local expert interventional radiologist are unavailable or if a large hemangioma is present on the right side of the liver at risk for puncture and bleeding if the percutaneous approach were to be used, surgical laparotomy and cannulation of a mesenteric venous tributary of the portal system should be considered. In this situation, complete surgical control is in place to prevent uncontrolled bleeding. Another advantage includes the potential for use of a dual-lumen catheter for cannulation of a mesenteric vein (i.e., dual-lumen 9-French Broviac line), which allows for continuous monitoring of portal pressure during islet infusion. Still, this surgical approach should be considered only when the percutaneous transhepatic approach cannot be used, because it presents several major disadvantages, including the requirement for a surgical incision, formation of adhesions, and the risk for wound infection and wound herniation, which may be exacerbated when the drug sirolimus is used after transplant, because this drug interferes with wound healing.

Risks to the Recipient Surgical Complications There are two potentially serious procedural complications in islet transplantation: bleeding from the catheter tract created by the percutaneous transhepatic approach, and portal vein thrombosis, particularly when large volumes of tissue are infused. Adverse bleeding events were noted early in the development of the Edmonton program, but these have been almost completely avoided with the routine use of effective methods to seal and ablate the transhepatic portal catheter tract on egress when the catheter is withdrawn [53]. The combination of coils and tissue fibrin glue (Tisseel) was used previously, but it has been replaced by Avitene paste (1 g Avitene powder mixed with 3 mL radiological contrast media and 3 mL saline; approximately 0.5e1.0 mL of this paste is injected into the liver tract) [54]. The use of purified islet allograft preparations has not resulted in main portal vein thrombosis in the Edmonton program, but thrombosis of a right or left branch or peripheral segmental vein has been encountered in approximately 5% of patients. Other rarely observed procedural side effects have included fine-needle gallbladder puncture, which has been avoided with the use of ultrasound-guided transhepatic portal venous access. Rarely, arteriovenous fistulae (which may require selective embolization) or hepatic steatosis have been observed [55]. Immunosuppressive Therapy and Complications Islet transplantation for T1DM represents a unique challenge in immunosuppression, because both alloimmunity and islet-specific autoimmunity must be effectively controlled to preserve graft function. An additional important consideration is that many of the immunosuppressive agents used in solid organ transplantation since the 1960s, particularly corticosteroids, are known to be toxic to islets. Previously in the Edmonton Protocol, the induction agent daclizumab (anti-CD25 [interleukin (IL)-2R] antibody) was used, but this was replaced by basiliximab when daclizumab was no longer available for clinical use. Maintenance immunosuppression consisted of sirolimus combined with low-dose tacrolimus. This regimen, described initially at the University of Alberta, has been successfully replicated at other centers as part of a multicenter Immune Tolerance Network (ITN) trial [39,56]. More recently, induction with depletional therapies including thymoglobulin (6 mg/kg total dose) or alemtuzumab (30 mg intravenously) combined with the antiinflammatory agents etanercept (antietumor necrosis factor [TNF]) and anakinra (anti-IL receptor [ILR]) have been employed in the Edmonton program. We and others found that the more standard posttransplant combination of tacrolimus (6e10 ng/mL) and mycophenolate mofetil (up to 2 g/day in a divided dose as tolerated) is much better tolerated than sirolimus. The 3- and 5-year outcome data suggest more graft durability in terms of sustained insulin independence with this regimen, which has been encouraging. In addition to the Edmonton Protocol immunosuppression described earlier, alternative regimens have been reported. The Minnesota Group, led by Dr. Bernhard Hering, used antithymocyte globulin (ATG) and etanercept (antieTNF-a antibody) induction with a combination of sirolimus and mycophenolate mofetil with or without low-dose tacrolimus for maintenance, or hOKT3g1 (Ala-Ala) (humanized anti-CD3 antibody) and sirolimus induction with sirolimus and reduced-dose tacrolimus for maintenance [47,48]. In some instances, alternative immunosuppressive agents have been used because of drug intolerance or other side effects. Islet patients often possess mild preexisting renal impairment as a result of long-standing diabetes, and this renal dysfunction may be exacerbated with calcineurin inhibitor therapy, even at the low doses involved in the Edmonton Protocol. The drug sirolimus may also have nephrotoxic side effects, which may be compounded when used in combination with a calcineurin inhibitor drug [57,58]. For these reasons, renal status must be monitored diligently in all patients after islet transplantation. In addition to its recognized nephrotoxicity, tacrolimus is associated with gastrointestinal side effects, which may lead to episodic diarrhea. Neurotoxicity may be seen with tacrolimus but is often avoided in

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low-dose regimens [59]. Sirolimus is associated with neutropenia and mouth ulceration, but these side effects can be reduced with lower target trough levels and tablet formulations. In the context of islet transplantation, sirolimus has been linked to a number of side effects including dyslipidemia, small bowel ulceration, peripheral edema, and the development of ovarian cysts or menstrual cycle irregularities in female recipients [51,60]. Combined with the observation that sirolimus is extremely poorly tolerated at high doses in this patient population, this has led to avoidance of this drug within our program. Although chronically immunosuppressed patients are at risk for developing all types of malignancy, squamous epithelial cancers most commonly occur and are most readily treatable. The Collaborative. Islet Transplant Registry reports an overall rate of less than 2% in islet transplant recipients [46]. The lifetime risk of lymphoma is estimated to be 1e2% in transplant recipients, but this risk is likely to be reduced in islet recipients, because these patients are generally not treated with glucocorticoids or OKT3.

FUTURE CHALLENGES Overcoming Tissue Shortage In its current form, islet transplantation is reserved for a small subset of patients with the most severe forms of T1DM. Even with the relatively small patient population selected for islet transplantation, the wait-list time for patients in Edmonton, which has access to organs from a large geographic region, ranges from 6 months to 2 years, depending on blood group. As islet transplantation becomes more suitable for a broader range of diabetic patients and as the incidence of diabetes increases, there will be an even more severe shortage of islet tissue for transplantation. Clinical islet programs rely on the scarce supply of pancreas organs derived exclusively from heartbeating, brain-dead cadavers. Compared with organs procured for whole-pancreas transplantation, which must fall within strict donor criteria, organs obtained for islet transplantation tend to be more “marginal” and come from older, less stable donors. Furthermore, the pancreas is particularly susceptible to toxicity from the circulating products of severe brain injury, hemodynamic instability, and inotropic support in a brain-dead organ or donation after cardiocirculatory death (DCD). The quality of the pancreas is further degraded by cold ischemic injury during transportation, which inevitably results in islet damage and loss. Brain death, with an acute cytokine and injury storm has been associated with islet injury, with consequential reduced islet recovery and viability compared to islet isolation in the absence of injury, at least in small animal models [61,62]. Because of the sporadic islet transplant experience with DCD pancreas donation in North America, a direct comparison with the standard neurological determination of death (NDD) donation has yet to be objectively conducted. A comparison study of clinical islet isolation success rates and transplantation outcomes between DCD and NDD pancreas donors from our center was published in 2015 [63]. Islet yields were similar between NDD and DCD. Likewise, metabolic function was similar between NDD and DCD, as well as the mean decrease in insulin requirement at 1 month after transplantation. Although these results support the broader use of DCD pancreata for islet isolation, a much larger DCD islet experience will be required to determine the noninferiority of both short- and long-term outcomes [63]. Unfortunately, once the pancreas is in the isolation laboratory, the extensive processing and purification steps during processing result in further islet destruction and loss, often resulting in at best 60% recovery of the estimated 107 IE/pancreas, [64]. As a result, nearly all islet recipients require islets derived from two cadaveric donors. Thus, a rapidly growing area of islet transplant research involves the development of improved cadaveric or alternative islet tissue sources for transplantation.

Living Donor Islet Transplantation One approach to alleviating islet tissue demand would be to use living donors for islet transplantation. Living donor programs in kidney, liver, and lung transplantation have moved forward successfully at most leading transplant centers worldwide, in an attempt to meet the growing demand for donor organs and improve clinical outcomes. Long-term clinical outcomes of islet transplantation alone, in specialized centers, mirror the results of whole-pancreas transplant alone; over 50% of patients have achieved insulin independence 5 years after transplant [65]. However, apart from the University of Minnesota experiences [47] and those at the University of California, San Francisco (UCSF) [66,67], the rate of single donor islet transplant success remains low (10e15%) [68]. Therefore, most patients require multiple pancreas donors to achieve insulin independence. In addition, despite remarkable progress in clinical islet transplantation since 1999, islet supply and functional viability remain

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significant challenges when islets are derived from cadaveric organ donors, even at the most experienced centers [61]. In the living donor setting, the distal half-pancreas could be procured under “ideal” circumstances without exposing the pancreas to hemodynamic instability or inotropic drugs, and the pancreas would be processed immediately without prolonged cold ischemia. Thus, the potency of islets derived from a living donor source is assumed to be far superior to cadaveric tissue. Living donor islet transplantation represents a unique opportunity to overcome donor organ shortage and procure the islet tissue under perfect conditions, with closer human leukocyte antigen matching between donor and recipient. Furthermore, the living donor islet transplant setting will provide a unique opportunity to develop protocols for pretransplant recipient conditioning for donor-specific tolerance induction. Although cadaveric islet transplantation has been an active area of clinical research involving more than 1500 patients over several decades, only three cases of living donor islet allotransplantation have been reported [69e71]. Initially, the rationale of pursuing living donor islet transplantation originated from discouraging patient outcomes in the pioneering series of deceased donor islet allografts conducted in the late 1970s [72]; the first attempts at islet autografts (after pancreatectomy for chronic pancreatitis) proved to be more promising [71,73]. The first two clinical attempts at living donor islet allotransplantation were carried out in 1978 by Sutherland and colleagues at the University of Minnesota [70,71,73]. Although neither recipient achieved sustained islet function, these pioneering efforts were truly remarkable given the early stage of clinical islet transplant development at the time. The first recipient’s graft was lost owing to a cytomegalovirus infection, and the second recipient lost function because of a sensitizing event as the result of a previously rejected kidney from the same donor (the recipient’s sister) [72]. The immunosuppression available was primitive by current standards (azathioprine and high-dose steroids), and the islets were isolated using suboptimal conditions, before the development of the Ricordi chamber and sophisticated purification schemes currently used in clinical islet transplantation. The dramatic improvement in clinical outcomes obtained in cadaveric islet transplantation since 2000 has renewed interest in developing living donor islet transplantation. The first living donor islet transplantation case attempted since the introduction of the Edmonton Protocol was carried out at the University of Kyoto in early 2005 [69]. The recipient, a 27-year-old woman, developed C peptideenegative, unstable diabetes after chronic pancreatitis as a child. Her 56-year-old mother was approved to be the donor, and islets were purified from the distal pancreas obtained during an open laparotomy. There were no surgical complications in either donor or recipient. The unpurified islet mass (408,114 IE [8200 IE/kg] in a volume of 9.5 mL after tissue digestion) was transplanted into the portal vein using the percutaneous approach under full systemic heparinization. Edmonton Protocolestyle immunosuppression was started before transplant using sirolimus and low-dose tacrolimus (started 7 days before transplant), anti-IL2R antibody (given 4 days before transplant and on the day of transplant), and anti-TNF-a blockade induction (infliximab; given 1 day before transplant). Insulin therapy in the recipient was discontinued at 22 days after transplant; this patient continued to be insulin independent with excellent glycemic control and a normal HbA1C more than 1 year after transplant [74]. The donor presented no evidence of glucose intolerance, maintained normal HbA1C values, and was C-peptide positive 30 months after transplant [72,74]. Although no definitive conclusions can be drawn from this single successful case of living donor islet allotransplantation, results from living donor islet autotransplantation suggest that the insulin independence may be achieved routinely with significantly less IE/kg recipient body weight than has been required for cadaveric allografts thus far. It is widely accepted that over 70% of patients will remain insulin free after islet autotransplantation if an islet mass exceeding 300,000 IE (2500 IE/kg) is transplanted, compared with the >10,000 IE/kg that is often required to achieve insulin independence with cadaveric islet preparations [75]. Reports from the Minnesota group showed that clinical islet autografts have a significantly lower rate of metabolic decay over time, even with a smaller islet implantation mass (108 cells/cm3) into large tissues without some sort of prevascularization or its alternative. Thus, a capillary network (and a lymphatic network) needs to be “engineered” as part of the creation of a larger structure. In a cell-free approach, vascularization and improvement of left ventricular (LV) function after MI were achieved by the sustained release of basic fibroblast growth factor (bFGF) incorporated into gelatin microspheres [106], acidic fibroblast growth factor (FGF) from ethylene vinyl acetate copolymer [107], and bFGF from heparinalginate beads [108]. Mooney and colleagues incorporated an endothelial cell mitogen (VEGF) into 3D porous poly(lactide-co-glycolide) scaffolds during fabrication [109] to promote scaffold vascularization. Sustained delivery of bioactive VEGF translated into a significant increase in blood vessel ingrowth in mice and the vessels appeared to integrate with the host vasculature. We use microencapsulated VEGF165-secreting cells (prepared

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by transfection of L929 cells) as a means of exploring this strategy, at least for microcapsules [110]. Of course, VEGF is but one angiogenic factor [111] and issues associated with the functional maturity of the vessels and the need for multiple factors may limit this strategy. In a third approach, Vacanti et al. micromachined a hierarchical branched network mimicking the vascular system in 2D. Silicon and Pyrex surfaces were etched with branching channels ranging from 500 to 10 mm in diameter [112] that were then seeded with rat hepatocytes and microvascular ECs. We covalently immobilized VEGF165 and angiopoietin-1 to porous collagen scaffolds to enhance scaffold vascularization in vitro and in vivo [113]. Such covalent immobilization offers the advantages of prolonged signaling and lowers the total amount of growth factors required; it also offers the possibility of generating capillary-like structures in the tissue engineered scaffolds in vitro. A prevascularized skeletal muscle was created [114] by coculturing skeletal muscle cells with ESC-derived EC and FBs. It appeared that up to 40% of the engineered blood vessels “connected” to the host vasculature upon implantation in this small-animal model. Finally, we adapted endothelial seeding in a modular approach to create scalable and vascularized tissue constructs. ECs were seeded onto submillimeter-sized collagen gel cylindrical modules that contained a second cell (e.g., hepatoma G2, smooth muscle cells, or [most relevant here], cardiomyocytes) [115]. With a view to creating uniform, scalable, and vascularized constructs, these modules were packed into a larger tube, formed into a sheet, or implanted directly with interconnected channels lined with ECs resulting from the random assembly of the modules. These channels connected with the host vasculature in vivo [116], creating a perfusable chimeric vasculature containing both host and donor cells and with host smooth muscle cell involvement. Embedded cardiomyocytes formed “contractile” structures near the periphery of modules, although the density of such structures was relatively low [115]. Remodeling occurred in vivo (after periinfarct injection or use as a patch), resulting in a well-distributed microvasculature (after 2 or 3 weeks in syngeneic animals), but the distribution of cardiac structures was relatively low. Tissue vascularization is an enduring challenge in the field of tissue engineering and numerous approaches have been devised to tackle it. Advances in microfabrication techniques provided us with increasing control over the spatial arrangement of cells when creating a 3D functional tissue. Building on this progress, we developed a 3D stamping technique based on conventional soft-lithography, which allowed us to pattern thin sheets of synthetic biodegradable polymers and stack them to form a scaffold with a built-in intricate microchannel network mimicking the native blood vessels. The scaffold (referred to as AngioChip) was subsequently populated with ECs within the microchannels and parenchymal cells in the surrounding matrix (Fig. 61.6A) [117]. The stable synthetic scaffold structure supports extensive tissue remodeling, resulting in a free-contracting cardiac muscle and a metabolically functional liver embedded with an internal perfusable vascular network (Fig. 61.6B). The built-in blood vessels were shown to be mechanically stable and permeable, and even allowed vascular sprouting under the guidance of angiogenic growth factors. Furthermore, we surgically connected the built-in vessel network within the engineered tissue to the host circulation system, establishing blood perfusion immediately upon implantation (Fig. 61.6C). This demonstration shows the feasibility of our bottom-up approach to overcome the lack of tissue survival after implantation caused by the absence of vascular perfusion.

Host Response and Biocompatibility Questions related to the immune and inflammatory response to tissue constructs are starting to draw attention. The host response to a tissue engineered construct is manifested by the innate and adaptive immune systems, involving both plasma (e.g., complement) and cellular components (e.g., macrophages, T cells, etc.), that are directed against engineered cells and grafts or the materials used in tissue constructs. This potent immune response is most often mediated by major histocompatibility complex mismatches between donor and host tissue in allogeneic transplantations. This response can also be manifested in situations in which autologous cells or tissues are engineered to express therapeutic but foreign factors, or if these autologous cells are placed in tissue constructs that themselves negatively affect immune consequences [118]. Immunosuppressants have enabled the successful transplantation of kidneys, hearts, and other organs. With the advent of tissue engineering, new configurations of tissues and organs (often with an added biomaterial component) are being developed and our understanding of the immune and inflammatory response to these new therapies is being shown to be inadequate. Some xenogeneic cell transplants (mice to rat) survive in situations of cardiac repair despite the species differences [119], although this may be specific to the animal model or to cardiac repair. The longevity of a transplant also depends on the ability of somatic cells to withstand and respond to the stresses of

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FIGURE 61.6 Microfabrication of vasculature [117]. (A) Schematic of the AngioChip scaffold illustrating the endothelialized lumen in the parenchymal space (left). Scanning electron micrograph of a multilayered AngioChip scaffold. Red arrows illustrate the 20-mm microholes designed to allow perfusion through the AngioChip (right). (B) Immunostaining of the hEC (human endothelial cell) vascular network for (i) CD31 (red), (ii) cardiac tissue stained for sarcomeric a-actinin (green), and F-actin (red), and (iii) liver tissue stained for E-cadherin (green), albumin (red), and 40 ,6-diamidine-20 -phenylindole dihydrochloride (DAPI) (blue). (C) Image of the AngioChip in an artery to vein graft in a rat. After the surgical anastomosis, blood perfusion was resumed.

implantation, rejection, and other injuries [120]. The classic “foreign body reaction” to biomaterials is well-known, but the details of the molecular signals (complement regulatory proteins and MMPs) that accompany this phenomenon (in the context of biomaterials) are only beginning to be defined. A variety of approaches have been undertaken or are under development to generate or improve upon graft acceptance [121]. These approaches include methods to block the innate immune response such as by using drugs or transferring genes to block nuclear factor-kB signaling pathways, for example. Other methods to block the innate response include the use of antibodies to IL-1 or TNF or the use of antiadhesion and antielastase antibodies. We must

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understand the mechanism of the host response itself better so that we can design better biomaterials, select or engineer more suitable cells, or devise better strategies to control both innate and adaptive immune responses and enable a functional integration of the new tissue with the host.

IN VIVO STUDIES In Situ Cardiac Tissue Engineering via Injection of Cells in Hydrogels Hydrogels have gained much attention as vehicles for the delivery of reparative cells into the myocardium owing to their injectability and ability to control cross-linking chemistry. For a hydrogel for use in myocardial regeneration, it needs to be: (1) biocompatible; (2) biodegradable; (3) injectable, so that it can be applied with a syringe in a minimally invasive manner; and (4) mechanically stable enough to withstand the beating environment of the heart. In addition, a biomaterial that can promote the attachment and survival of cells and localize them at the infarction site would address these limitations of poor cell retention and survival. Early studies relied on cell injection using natural hydrogels such as Matrigel [122,123] or fibrin [124e126] and reported structural stabilization, reduced infarct size, and improved vascularization upon injection of undifferentiated ESCs [122,123] or bone marrow cells [124e126]. Alginate alone was demonstrated to reduce pathological remodeling and improve function [127]. A synthetic, self-assembling peptide hydrogel (AcNRARADADARARADADA-CNH) was also used, forming a nanofibrous structure upon injection into the myocardium that promoted recruitment of endogenous ECs and supported survival of injected cardiomyocyte (CM) [128]. Insulin-like growth factor-1 bound to the self-assembling peptide was demonstrated to improve grafting and survival of CM injected into MI [129]. Laflamme and Murry demonstrated that targeting of multiple pathways, which was related to cell survival by encapsulating a number of biomolecules in Matrigel, significantly increased the survival and grafting of the hESC-derived CM injected into infarcted rat hearts [130]. They carried on studies using the same hydrogel to deliver hESC-derived cardiomyocytes in guinea pig [131] and nonhuman primates [132], demonstrating that the cells delivered with the hydrogel could indeed integrate, upon a remodeling phase that included the presence of significant ventricular arrhythmias. Zhang et al. studied the effect of injecting CM in a mixture of collagen type I and Matrigel [133], the material used by Zimmerman et al. to create engineered heart tissue [134], in MI-induced rats. An additional problem with using ECM proteins in this setting may be the immune response exhibited by rats to mouse proteins (i.e., Matrigel is a basement membrane derived from mouse sarcoma). Positive connexin-43 staining was found in cells in the biomaterial, and the biomaterial was also seen to improve the thickness and function of the heart. The main drawback is that the material takes 1 h to gel, which could allow for significant cell loss, although no cell retention studies were conducted in these experiments. We modified chitosan with the peptide QHREDGS derived from angiopoietin-1, the peptide sequence implicated in the survival response of muscle cells cultivated in the presence of this growth factor [135]. The chitosan was rendered photocross-linkable by modification with azidiobenzoic acid (Az-chitosan) [136]. Neonatal rat heart cells cultivated on cross-linked films of Az-chitosaneQHREDGS attached, elongated, and remained viable whereas they exhibited lower attachment levels and decreased in viability when cultivated on the chitosan substrates modified with the scrambled peptide sequence [137]. Interestingly, cells on Az-chitosaneQHREDGS were capable of resisting Taxol-induced apoptosis, whereas those on Az-chitosaneArg-Gly-Asp-Ser were not [137]. We also demonstrated that QHREDGS peptide was able to attenuate pathological remodeling and improve cardiac function upon MI in a rat model, acting largely by promoting the survival of spared cardiomyocytes [138]. Other studies collectively indicated that an injection of hydrogel alone, without the reparative cells, may attenuate pathological remodeling upon MI [127,139e141]. For example, injection of alginate or collagen alone improved LV function and reduced cardiac remodeling after infarction [127,142]. Two alginate biopolymers were modified to assess therapeutic potential in rat MI models. Alginate modified with 0.025% v/v polypyrrole, a conductive polymer, injected into the infarct zone showed improved arteriogenesis at 5 weeks posttreatment and significantly enhanced infiltration of myofibroblasts into the infarct area compared with saline and alginate-only controls [143]. Also, Arg-Gly-Aspeconjugated alginate and alginate alone injected into the infarct zone showed improved LV function and increased arteriole density 5 weeks after injection compared with bovine serum albumin in phosphate-buffered saline control [144]. Results from both studies again show the potential for nonecell based therapies to treat chronic heart failure.

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An interesting example is the injection of a hydrogel based on decellularized porcine ventricular myocardium, which resulted in improvements in local and global cardiac function [145,146] and thus the initiation of Phase I clinical trials with this hydrogel. It is suspected that by changing the ventricular geometry and mechanics, hydrogels reduce the elevated local wall stresses that have been implicated in pathological remodeling [147]. Finite element modeling of wall stresses indicated that upon injection of the material of elastic modulus 10e20 kPa in the infarct, the relationship between ejection fraction and the stroke volume/end-diastolic volume was improved. In addition, injections of the material in the border zone decreased end-systolic fiber stress proportionally to the volume and the stiffness of the injected material. We believe that properly tuning the mechanical properties of a hydrogel and providing bioactive molecules may offer new cell-free treatment options for MI. The death of CM by necrosis and apoptosis peaks at 6 h upon acute MI [148]. However, the persistent and progressive loss of CM in neighboring areas of the infarct continues up to 60 days after the onset of MI. During this process, up to 35% of cells at the borders of subacute and old infarcts may become apoptotic [149], compared with only 1% in the remote regions of myocardium [150]. Studies in rats and dogs demonstrated that CM loss by apoptosis persists for 1e4 months upon MI, which correlates with the progressive worsening of the pump function. Thus, developing hydrogels that specifically prevent apoptosis of the heart cells (e.g., QHREDGS peptide-modified chitosan) may result in new treatment options in the future, in which hydrogel injected alone in the border zone, without reparative cells, would act to mechanically stabilize the ventricle and prevent further apoptosis of cardiomyocytes. For example, it was shown that EC-induced cardiomyocyte protection after infarction occurs through PDGF-BB signaling. Thus, binding PDGF-BB to the self-assembling nanofibers of RAD16-II (a peptide consisting of alternating RAD domains, AcN-RARADADARARADADA-CNH2) hydrogel was evaluated as a potential therapeutic option. Sustained, targeted release of this signaling molecule to host myocardium was observed up to 14 days after injection. Injection of nanofibers with PDGF-BB at the site of infarct in rats decreased CM death and preserved systolic function after MI and showed (separately) a decrease in infarct size after ischemiaereperfusion [151]. The relative contribution of cells versus the injected biomaterial to the attenuation of pathological remodeling also needs to be assessed and the mechanism by which various cells induce functional improvements needs to be elucidated. Although with the injection of contractile cardiomyocytes the expectation is that cells will functionally couple to the host myocardium and contribute to contractile function, the same is not possible for noncardiomyocytes. The exact mechanism by which nonmyocytes impart improvement in function and attenuation of pathological remodeling is still under debate, but some researchers suggest that the transplantation of healthy cells results in the release of growth factors and other molecular signals, i.e., the paracrine effect. These help with angiogenesis, cell survival, and recruitment of progenitors. A possible drawback of using the biomaterial is that the scaffold or hydrogel may also take up space that would prevent a high tissue density until the material degrades.

Implantation of Cardiac Patches Significant progress has been made in constructing in vitro cultivation systems and biomaterial scaffolds, but fewer studies have focused on implanting cell-based cardiac patches onto viable or injured myocardium. In a pioneering study, Li and colleagues [18] implanted a construct based on neonatal rat cardiomyocytes and collagen sponges onto the surface of the cryoinjured myocardium of Lewis rats. The grafts were implanted 3 weeks after infarction. After 5 weeks in vivo, cells survived supported by blood vessel ingrowth and integrated with the surrounding tissue. However, the graft did not improve LV function. Attenuation of pathological remodeling (i.e., prevention of ventricle dilatation and maintenance of contractile function) was observed in a study by Leor et al. [20], in which cardiac constructs based on neonatal rat cardiomyocytes and porous alginate scaffolds were implanted onto myocardium of SpragueeDawley rats that underwent permanent main coronary artery occlusion. The grafts were implanted 7 days after MI. After 9 weeks of implantation, the grafts demonstrated integration with host myocardium at the anchorage sites as well as inflammatory infiltrates and the presence of fibrous collagen. Zimmerman et al. [15] placed cardiac tissue rings cultivated in the presence of mechanical stimulation onto uninjured hearts of Fisher 344 rats for 14 days. They noticed that although both cells and collagen were isolated from Fisher rats, immunosuppression was required to maintain of heart tissue upon implantation. In the absence of immunosuppression, even in the syngeneic approach, cardiac constructs completely degraded after only 2 weeks in vivo. It is unknown what caused the response; it is possible that it was the remainder of serum or chick extract.

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Regardless, the finding has significant implications in the potential implantation of cardiac patches in clinical settings. To decrease the potential immunogenicity of their engineered tissue, Zimmerman and colleagues discarded all xenogenic components from their culture [152]. This included cultivating the engineered heart tissues in serumfree and Matrigel-free conditions. Mixed heart cell populations rather than cardiomyocyte-rich populations were used and the culture medium was supplemented with triiodothyronine and insulin [152]. Other studies also established the need for nonimmunogenic media. Schwarzkopf et al. used autospecies sera, in this case rat, to culture rat cardiomyocytes [153]. The metabolic activity of cells was significantly higher than for the cells cultivated in conventional culture medium with fetal bovine serum. Zimmermann et al. demonstrated integration and electrical coupling of a complex multiloop graft to native myocardium in rats with left anterior descending coronary artery ligation (Fig. 61.7A and B). Functional improvement was demonstrated not to be merely the result of scar stabilization or paracrine effects [134]. Functional integration of cardiac cell sheets to the heat-injured myocardium was also demonstrated [154]. In a study performed by Weinberger et al., human engineered heart tissue strips composed of fibrinogen and Matrigel with suspended human iPSC-derived cardiomyocytes were created [155]. The engineered strips, which were intended to limit the number of cells that were washed out after a cell injection, were sutured onto the hearts of cryoinjured guinea pigs. Hearts of the guinea pigs treated with the engineered strips had an improved LV function compared with animals that received a cell free patch, and the graft was found to be electrically coupled to the host myocardium. However, 28 days after implantation, cells were found in the spleen and lungs of some of the guinea pigs. Although cardiac patches still require further development and regulatory approval, there are numerous in vivo studies improving their cell retention and proving their effectiveness. A cardiac patch developed by Zhang and colleagues was transplanted onto an MI porcine model to test whether including multiple cell types and growth factors would aid in cell retention [156]. The patch was composed of fibrin loaded with insulin growth factor and was encapsulated with a combination of human iPSC-derived cardiomyocytes, ECs, and smooth muscle cells. After transplantation, the group saw that the cell-laden patch increased ejection fraction and the ventricular wall thickness while it decreased the infarct size, which suggested that incorporating a growth factor into the cardiac patch and multiple cell types may be required for the success of such a therapeutic approach (Fig. 61.7C and D).

FIGURE 61.7 In vivo integration of engineered cardiac patches. (A) A multilooped, mechanically stimulated cardiac construct based on neonatal rat cardiomyocytes and collagen hydrogel was used to repair myocardial infarction (MI) in the rat heart [134]. (B) Histology of the multilooped patch showing integration with the surrounding myocardium [134]. (C) A cardiac patch implanted on an MI porcine model was shown to decrease the infarct size compared with animals not treated with a patch [156]. (D) Immunostaining of the implanted patch shows increased vascular density at the border zone. Image on the left is stained for CD31 (green), smooth muscle actin (SMA) (red), and cardiac troponin T (cTnT) (white). In addition, less apoptotic cells were found at the border zone. Image on the right is stained for apoptotic cells (terminal deoxynucleotidyl transferase dUTP nick end labeling assay) (red), cardiac troponin I (green), and 40 ,6-diamidine-20 -phenylindole dihydrochloride (DAPI) (blue) [156].

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The first human clinical trial using a patch-based therapeutic was performed by Larghero and colleagues. In this case report, hESC-derived cardiac progenitors were embedded in a fibrin glue scaffold [157]. The pericardium over the infarcted area was retracted, creating a pocket with which the patch was placed. At 3 months after surgery, the patient’s ejection fraction had improved from 26% to 36% and a reduction in end-diastolic and systolic volumes was noted. Despite the increase in symptomatic performance, this trial was conducted only on a single patient; therefore, it is difficult to draw conclusions. However, this trial shows the potential for such a treatment and future clinical trials. In addition to engineering the patches of myocardium, Zimmermann and colleagues designed the first biological assist device [158]. The authors mechanically stimulated a hollow-spherical construct consisting of collagen I and neonatal rat cardiomyocytes until a beating pouch-like structure was created. The pouch was then placed over uninjured rat hearts in such a manner that the right and left ventricles were covered. Fourteen days after implantation, the pouch covered the epicardial surface of the heart and exhibited blood vessel ingrowth. Roche et al. created an alternative ventricular assist device that uses soft robotics to pump the heart manually. The sleeve consists of actuators aligned helically to mimic the twisting motion of the beating heart. It was tested in vivo on a pig model with acute heart failure and the researchers observed the reestablishment of cardiac output. The group hopes to see the technology used in the clinic as a device to extend the life of a patient waiting for transplant [159]. Badylak et al. implanted ECM derived from porcine urinary bladder into a surgically created 2-cm [2] defect in the LV free wall of dogs. Eight weeks after implantation, the ECM patches showed higher regional systolic contraction compared with the control group in which a material (Dacron) used for myocardial defects was employed. Histological analysis suggested that cardiomyocytes accounted for about 30% of the remodeled tissue in the ECM scaffolds [160]. In another study, this improvement in the heart function was attributed to an increase in the myocyte content in the ECM patches between weeks 2 and 8. The relationship between the myocyte content and the extent of mechanical function was observed to be linear. There was also some evidence (a decrease in cardiomyocyte diameter and an increase in the overall area occupied by cardiomyocytes over time) suggesting the possibility of cardiomyocyte proliferation in the patches [161]. Limitations related to the source of autologous cardiomyocytes motivated studies that used nonmyocyte-based patches for MI repair. Smooth muscle cells seeded with poly (ε-caprolactone-co-L-lactide) sponge reinforced with poly-L-lactide fabric were used in a modified endoventricular circular patch plasty procedure (Dor procedure). Cell-seeded grafts resulted in improved LV function (as assessed by echocardiography) compared with cell-free controls [162]. A patch made of dermal FBs seeded onto knitted Vicryl mesh (Dermagraft) was used in an attempt to increase angiogenesis upon MI. When placed over the infarcted regions on the hearts of SCID mice, the grafts improved microvessel density within the damaged myocardium [163]. There appears to be a consensus regarding the requirement for multiple cell types, specifically FBs and ECs, in addition to cardiomyocytes for successful cardiac tissue graft survival and vascularization in vivo. In one approach, omentum was used to prevascularize cardiac patches based on neonatal rat cardiomyocytes and alginate scaffolds modified with angiogenic factors. After excision and implantation into the infarcted rat myocardium, the vascularized cardiac patch showed structural and electrical integration into host myocardium and attenuated pathological remodeling of the ventricle significantly better than did the in vitro cultivated patch alone [164]. In another strategy, a simultaneous triculture scaffold-free approach was used to generate beating cardiac patches based on hESC. Upon implantation into the hind-limb muscle of nude rats, these patches, which were composed only of enriched cardiomyocytes, did not survive to form significant grafts. However, patches containing ECs (either HUVECs or hESC-derived ECs) and FBs in addition to cardiomyocytes persisted in the (noninfarcted) rat heart and resulted in 10-fold larger cell grafts compared with cardiomyocyte-only patches. The preformed human microvessels also anastomosed with the rat host coronary circulation and delivered blood to the grafts [165]. These studies demonstrated the feasibility of cardiac patch implantation, but further studies are necessary to estimate the effect of culture conditions and scaffold type on the in vivo outcome. Although significant progress has been made in the area of biomaterials and bioreactors, it is unknown which cultivation conditions and what biomaterial will best preserve contractile function and prevent pathological remodeling upon implantation. Thus, studies that investigate this systematically and correlate in vitro parameters (e.g., force of contraction) to in vivo outcomes (e.g., fractional shortening) are required. The host response to the patch and the nature of the immunologic situation further complicate these studies.

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Delivery Route for Cells and Patches It remains to be determined how and where in the heart the cells and patches should be applied for optimal therapeutic benefit. The ventricular wall is composed of the epicardium, the outer epithelial lining of the muscle; the myocardium, composed of working CMs and support cells such as FBs; and the endocardium, consisting of a specialized endothelial lining that is in contact with blood filling the ventricle. Most cell therapy approaches focus on the direct injection of a cell suspension into the myocardium using either catheters or open chest surgery [166e168]. These approaches are inherently limited by the small volume of free space available for cell engraftment. The cell density in the myocardium is extremely high (108 cells/cm3 [169]); upon injury such as MI the dead cells are replaced by a dense collagenous scar, leaving no space for injected cells to engraft. The injected cells are rapidly washed out or ejected, resulting in a maximal engraftment efficiency of 10% [20,130,170,171], which necessitates the injection of as many as 1 billion CMs for sizable grafts in monkeys [132]. The injection of isolated cells also offers little opportunity to mature or modify the electrophysiological properties of the graft. The immaturity of the injected cells was postulated to have caused malignant arrhythmias in monkeys that received hESC-derived CMs [132]. Thus, cardiac patches may offer a unique solution to the challenges of cell delivery, engraftment, and graft function. The size, shape, and functional properties of cardiac patches can be tailored through in vitro cultivation with electrical field stimulation [54,73] or mechanical stimulation, whereas their application on the ventricular surface could provide sufficient space for implant engraftment to a clinically relevant size. In fact, in a clinical trial, hESC-derived cardiac progenitors were applied during open heart surgery onto the epicardial surface of the heart using a fibrin hydrogel and a pericardial patch over the hydrogel [157]. The epicardium is generally considered a safe location for patch deployment, owing to the high risk for embolization and thrombosis associated with patch application inside the ventricle on the endocardial surface of the ventricular wall.

SUMMARY Functional viable cardiac patches have been engineered based on neonatal rat cardiomyocytes and hESC- and iPSC-derived cardiomyocytes. Various biomaterials have been tested for this purpose and in vitro culture systems have been developed that enhance cardiac construct differentiation (mechanical and electrical stimulation), improve cardiomyocyte survival at high density (medium perfusion), and enhance maturation of cardiomyocytes derived from human pluripotent stem cells, paving the way to engineering autologous cardiac patches of clinically relevant size. The field of microphysiological organ systems and organ-on-a-chip engineering has emerged and matured to the level required to use these miniaturized 3D cardiac tissues in drug discovery, safety testing, and disease modeling, relying on human cardiomyocytes derived from iPSCs. Because the in vivo studies conducted thus far used different cell sources, biomaterials, animal models, delivery times after infarction, and experimental time frames, direct comparison cannot be made between methods. Although all reported studies have shown some form of improvement, complete myocardial regeneration has not been achieved. Perhaps a valid question to be answered in the future is: What is the required level of myocardial regeneration in terms of survival and attenuation of symptoms? Complete regeneration is an ambitious goal that may not be required. Future studies must also increase their time frames to assess the long-term effects of these treatments better. Although the completely artificial heart will remain a dream, the near future may bring a clinically relevant autologous cardiac patch as evidenced by rapid progress in engineering cardiac patches based on stem cellederived cardiomyocytes. Work on recellularizing decellularized hearts may represent the first step toward “the heart in a box” envisioned several decades ago.

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Engineered heart tissue grafts improve systolic and diastolic function in infarcted rat hearts. Nat Med 2006;12: 452e8. [135] Dallabrida SM, Ismail N, Oberle JR, Himes BE, Rupnick MA. Angiopoietin-1 promotes cardiac and skeletal myocyte survival through integrins. Circ Res 2005;96:e8e24. [136] Yeo Y, et al. Photocrosslinkable hydrogel for myocyte cell culture and injection. J Biomed Mater Res B Appl Biomater 2007;81:312e22. [137] Rask F, et al. Photocrosslinkable chitosan modified with angiopoietin-1 peptide, QHREDGS, promotes survival of neonatal rat heart cells. J Biomed Mater Res A 2010;95A:105e17. [138] Reis LA, et al. Hydrogels with integrin-binding angiopoietin-1-derived peptide, QHREDGS, for treatment of acute myocardial infarction. Circ Heart Fail 2015;8:333e41. https://doi.org/10.1161/CIRCHEARTFAILURE.114.001881. [139] Fujimoto KL, et al. Synthesis, characterization and therapeutic efficacy of a biodegradable, thermoresponsive hydrogel designed for application in chronic infarcted myocardium. Biomaterials 2009;30:4357e68. [140] Dobner S, Bezuidenhout D, Govender P, Zilla P, Davies N. A synthetic non-degradable polyethylene glycol hydrogel retards adverse postinfarct left ventricular remodeling. J Card Fail 2009;15:629e36. [141] Leor J, et al. Intracoronary injection of in situ forming alginate hydrogel reverses left ventricular remodeling after myocardial infarction in Swine. J Am Coll Cardiol 2009;54:1014e23. [142] Dai W, Wold LE, Dow JS, Kloner RA. Thickening of the infarcted wall by collagen injection improves left ventricular function in rats: a novel approach to preserve cardiac function after myocardial infarction. J Am Coll Cardiol 2005;46:714e9. pii: S0735-1097(05)01201-5. [143] Mihardja SS, Sievers RE, Lee RJ. The effect of polypyrrole on arteriogenesis in an acute rat infarct model. Biomaterials 2008;29:4205e10. [144] Yu J, et al. The effect of injected RGD modified alginate on angiogenesis and left ventricular function in a chronic rat infarct model. Biomaterials 2009;30:751e6. [145] Seif-Naraghi SB, et al. Safety and efficacy of an injectable extracellular matrix hydrogel for treating myocardial infarction. Sci Transl Med 2013;5:173ra125. https://doi.org/10.1126/scitranslmed.3005503. [146] Wassenaar JW, et al. Evidence for mechanisms underlying the functional benefits of a myocardial matrix hydrogel for post-MI treatment. J Am Coll Cardiol 2016;67:1074e86. https://doi.org/10.1016/j.jacc.2015.12.035. [147] Wall ST, Walker JC, Healy KE, Ratcliffe MB, Guccione JM. Theoretical impact of the injection of material into the myocardium: a finite element model simulation. Circulation 2006;114:2627e35. [148] Anversa P, et al. Apoptosis and myocardial infarction. Basic Res Cardiol 1998;93(Suppl. 3):8e12. [149] Yaoita H, Ogawa K, Maehara K, Maruyama Y. Apoptosis in relevant clinical situations: contribution of apoptosis in myocardial infarction. Cardiovasc Res 2000;45:630e41. [150] Olivetti G, et al. Acute myocardial infarction in humans is associated with activation of programmed myocyte cell death in the surviving portion of the heart. J Mol Cell Cardiol 1996;28:2005e16. [151] Hsieh PCH, Davis ME, Gannon J, MacGillivray C, Lee RT. Controlled delivery of PDGF-BB for myocardial protection using injectable selfassembling peptide nanofibers. J Clin Invest 2006;116:237e48. [152] Naito H, et al. Optimizing engineered heart tissue for therapeutic applications as surrogate heart muscle. Circulation 2006;114:I72e8. [153] Schwarzkopf R, et al. Autospecies and post-myocardial infarction sera enhance the viability, proliferation, and maturation of 3D cardiac cell culture. Tissue Eng 2006;12:3467e75. [154] Furuta A, et al. Pulsatile cardiac tissue grafts using a novel three-dimensional cell sheet manipulation technique functionally integrates with the host heart, in vivo. Circ Res 2006;98:705e12. [155] Weinberger F, et al. Cardiac repair in Guinea pigs with human engineered heart tissue from induced pluripotent stem cells. Sci Transl Med 2016;8:363ra148. [156] Ye L, et al. Cardiac repair in a porcine model of acute myocardial infarction with human induced pluripotent stem cell-derived cardiovascular cell populations. Cell Stem Cell 2014;15:750e61. https://doi.org/10.1016/j.stem.2014.11.009. [157] Menasche´ P, et al. Human embryonic stem cell-derived cardiac progenitors for severe heart failure treatment: first clinical case report. Eur Heart J 2015;36:2011e7. https://doi.org/10.1093/eurheartj/ehv189. [158] Yildirim Y, et al. Development of a biological ventricular assist device: preliminary data from a small animal model. Circulation 2007;116: I16e23. [159] Roche ET, et al. Soft robotic sleeve supports heart function. Sci Transl Med 2017;9. https://doi.org/10.1126/scitranslmed.aaf3925.

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[160] Badylak SF, et al. The use of extracellular matrix as an inductive scaffold for the partial replacement of functional myocardium. Cell Transplant 2006;15(Suppl. 1):S29e40. [161] Kelly DJ, et al. Increased myocyte content and mechanical function within a tissue-engineered myocardial patch following implantation. Tissue Eng Part A 2009;15:2189e201. https://doi.org/10.1089/ten.tea.2008.0430. [162] Matsubayashi K, et al. Improved left ventricular aneurysm repair with bioengineered vascular smooth muscle grafts. Circulation 2003; 108(Suppl. 1):II219e225. [163] Kellar RS, et al. Scaffold-based three-dimensional human fibroblast culture provides a structural matrix that supports angiogenesis in infarcted heart tissue. Circulation 2001;104:2063e8. [164] Dvir T, et al. Prevascularization of cardiac patch on the omentum improves its therapeutic outcome. Proc Natl Acad Sci USA 2009;106: 14990e5. [165] Stevens KR, et al. Physiological function and transplantation of scaffold-free and vascularized human cardiac muscle tissue. Proc Natl Acad Sci USA 2009;106:16568e73. [166] Makkar RR, et al. Intracoronary cardiosphere-derived cells for heart regeneration after myocardial infarction (CADUCEUS): a prospective, randomised phase 1 trial. Lancet 2012;379:895e904. [167] Perin EC, et al. A phase II dose-escalation study of allogeneic mesenchymal precursor cells in patients with ischemic or nonischemic heart failure. Circulation Research 2015;117:576e84. https://doi.org/10.1161/CIRCRESAHA.115.306332. [168] Traverse JH, et al. Effect of the use and timing of bone marrow mononuclear cell delivery on left ventricular function after acute myocardial infarction: the TIME randomized trial. J Am Med Assoc 2012;308:2380e9. https://doi.org/10.1001/jama.2012.28726. [169] Mandarim-de-Lacerda CA, Pereira LMM. Numerical density of cardiomyocytes in chronic nitric oxide synthesis inhibition. Pathobiology 2000;68:36e42. [170] Reinecke H, Murry CE. Taking the death toll after cardiomyocyte grafting: a reminder of the importance of quantitative biology. J Mol Cell Cardiol 2002;34:251e3. https://doi.org/10.1006/jmcc.2001.1494. [171] Mu¨ller-Ehmsen J, et al. Survival and development of neonatal rat cardiomyocytes transplanted into adult myocardium. J Mol Cell Cardiol 2002;34:107e16.

Further Reading Sefton MV, Zandstra P, Bauwens CL, Stanford WL. In: del Nido PJ, Swanson SJ, editors. Sabiston and Spencer surgery of the chest. Elsevier; 2005.

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C H A P T E R

62 Bioengineering of Liver Tissue Pilar Sainz-Arnal1,2, Iris Pla-Palacı´n1, Natalia Sa´nchez-Romero1, Manuel Almeida1, Sara Morini1,3, Estela Solanas1, Alberto Lue1,4, Trinidad Serrano-Aullo´1,4, Pedro M. Baptista1,5,6,7 1

Health Research Institute of Arago´n (IIS Arago´n), Zaragoza, Spain; 2Instituto Aragone´s de Ciencias de la Salud (IACS), Zaragoza, Spain; 3Universidade de Lisboa, Lisbon, Portugal; 4Lozano Blesa University Hospital, Zaragoza, Spain; 5Center for Biomedical Research Network Liver and Digestive Diseases (CIBERehd), Zaragoza, Spain; 6Instituto de Investigacio´n Sanitaria de la Fundacio´n Jime´nez Dı´az, Madrid, Spain; 7Universidad Carlos III de Madrid, Madrid, Spain

INTRODUCTION The liver is the largest internal gland in the body and one of the most critical organs in metabolic homeostasis because it is responsible for many essential metabolic exocrine and endocrine functions. Among them, it is responsible for the detoxification and elimination of a variety of substances, regulation of glucose levels, maintenance of blood homeostasis, and production of many products including lipids, proteins, vitamins, and carbohydrates. In addition to this, the liver possesses a unique regenerative capacity; it is able to regenerate most of its function after losing up to three-quarters of its mass as a result of partial hepatectomy or toxic injury. However, in severely ill patients with an end-stage liver disease, this regeneration capacity is severely limited and liver transplantation (LT) becomes the only lifesaving intervention. This also applies to patients with decompensated cirrhosis, hepatocellular carcinoma (HCC), inborn errors of metabolism, and fulminant hepatic failure. Moreover, LT remains the only available cure for a wide range of liver diseases. Nearly a half century has passed since Dr. Starzl [1] performed the first liver transplant. However, it was not until 1967 when short-term success was achieved with 1-year survival after transplantation [2]. Since then, the success and the survival rates of LT have increased owing to improved immunosuppressive regimens, surgical techniques, and donorerecipient pairing [3]. However, today, the number of patients waiting for a liver transplant far exceeds the number of transplants performed. Although this gap was reduced with the increased use of marginal organs such as steatotic livers, employment of hepatitis C virus (HCV)-positive donor organs for HCV-positive recipients, and donation after cardiac death, the number of patients waiting for a liver transplant continues to exceed the number of available livers. According to reported by the Organ Procurement and Transplantation Network in 2014, 6729 liver transplants were performed in adults whereas around 14,600 people were on the waiting list to receive a donor’s liver. Although 6729 transplants were performed, 1821 patients died while waiting for a donor and 1290 patients were removed from the list because they were too sick for transplantation during that year [4]. Despite attempts to reduce the waiting time, two-thirds of all patients will never receive one. Liver cell transplantation has also been considered a possible treatment for patients with life-threatening liver diseases. In this case, a preparation of isolated cells is injected directly into the liver via a portal vein or into the spleen. Engraftment of these cells and their therapeutic potential to correct many metabolic deficiencies has also been demonstrated by several studies [5]. In 1992, a group of scientists from Michigan University performed the first successful hepatocyte transplantation into a French-Canadian woman with familial hypercholesterolemia [6]. Afterward, other successful cases followed, but not all patients had noticeable progress [7], which was also observed with other diseases [8,9]. Although the long-term safety of this technique has been confirmed [10,11], the success

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of this procedure has not yet been enough to evade the need for whole-organ transplantation. One possible explanation for these results could be the small number of hepatocytes that engraft in the recipient owing to the poor quality and quantity of the cells used [12]. Hence, it is accurate to say that as a result of hepatocytes, the shortage of human liver donors is the limiting factor in the development of this therapeutic approach. Other sources such as hematopoietic stem cells [13], bone marrowederived mesenchymal stem cells [14,15], and fetal liver progenitors [16,17] have also been used (Fig. 62.1). In addition to cellular therapies, other experimental approaches are in development with high therapeutic potential. These novel approaches aim to stimulate the diseased liver to regenerate itself or substitute it with a bioengineered liver capable of performing all native liver functions [18]. In this regard, and with the available technology, liver bioengineering comprises the creation and use of a supporting scaffold and large numbers of hepatic and vascular cells to generate functional tissue. Considerable effort has been made in trying to develop biomaterials capable of mimicking the liver extracellular matrix (ECM) to replicate the benefits offered by this structure regarding cell adhesion, growth and viability, and maintenance of the differentiation state and metabolic functions. Despite this, in the first studies performed with synthetic materials, the maintenance of the vascular network was not successfully achieved, resulting in low hepatocyte survival [19]. Some of these obstacles were solved by administering growth factors to promote angiogenesis and increase cell viability [20,21], but also by developing advanced bioreactors to improve hepatocyte seeding [22]. However, the absence of specific hepatic signals for cell growth and differentiation proved to be another critical limitation, probably owing to the lack of a real three-dimensional (3D) environment in which hepatocytes could find or create their native niche. Several research groups have addressed this issue by employing tissue decellularization to prepare bioscaffolds devoid of all cellular material from a harvested tissue or organ. In 2009, Baptista et al. successfully repopulated decellularized rat and ferret livers with

FIGURE 62.1 Available strategies for the isolation and generation of hepatic cells and tissues. ES, embryonic stem cells; hES, human embryonic stem cells; ICM, inner cell mass; iPS, induced pluripotent stem cells.

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human fetal liver and endothelial cells [23,24]. Immediately afterward, Uygun’s group achieved the decellularization of rat livers, which were then repopulated with rat primary hepatocytes and anastomosed ex vivo with a rat for up to 8 h [25]. It is still difficult to know the real value of this technology in clinical applications, but the potential to change transplantation medicine and patients’ lives is evident. Hence, in this chapter, we will describe some of these novel regenerative medicine approaches and the potential applications for these tissue-engineered liver constructs as biologic surrogates for a large variety of biomedical and pharmaceutical applications.

HEPATIC TISSUE ENGINEERING It is widely accepted today that 3D cultures mimic in vivo natural conditions much better than do traditional 2D cultures, in which cells are grown on a flat surface in Petri dishes. The reason for this is that 3D cultures more closely reflect in vivo cell behavior, providing a physical and mechanical environment that promotes cell proliferation, differentiation, and expression of specific growth factors and other molecules. Hepatic tissue engineering has been able to create bioengineered grafts with the medical purpose of organ replacement, or restoration, which recreate physiological 3D microenvironments that more accurately mimic the in vivo environment. These promising techniques rely on the combination of different disciplines such as engineering, biomaterials, cells, and growth factors. Starting with the biomaterials used, these can be made of either natural or artificial components capable of supporting 3D cellular growth. Some examples of their applications in tissue engineering are described subsequently.

Chitosan Chitosan is a natural polysaccharide derived from chitin, a structural polymer of arthropod exoskeletons [26], which has been widely used in tissue engineering because of the structural similarity with glycosaminoglycans, another component of liver ECM [27,28]. Fan et al. [28] combined chitosan with galactosylated hyaluronic acid, making porous 3D sponges that induced hepatocyte aggregation that displayed liver-specific metabolic activities. Gong et al. used chitosanegelatin liver scaffolds with a hierarchical predesigned channel network inside, providing a porous structure. They compared this fabricated scaffold with conventional porous ones, obtaining higher proliferation rates (62%) of HepG2 [29]. Jiankang et al. [30] also designed a model of a porous scaffold made of chitosan and gelatin with internal structures such as portal vein, hepatic chambers, well-organized fluidic channels, and central vein, demonstrating improved hepatocyte aggregation and liver-specific functions. In another experiment, Zang et al. [31] determined that 1% of genipin cross-linked with chitosanegelatin scaffolds present better results in HepG2 performance compared with glutaraldehyde, and r1-(3-dimethylaminopropyl) 23-ethyl-carbodiimide hydrochloride cross-linked scaffolds. Galactosylated chitosan (GC) is also commonly used, combined with hyaluronic acid, an anionic glycosaminoglycan widely present in mammal ECM, to create porous sponges in which hepatocytes and endothelial cells were cultured [32]. This hybrid scaffold could represent a cell storage and delivery vehicle for a bioengineered liver tissue. Chung et al. [33] combined GC with alginate, creating a highly porous 3D sponge for hepatocyte anchorage. They demonstrated better spheroid formation and higher cell viability compared with scaffolds made only of alginate. Chitrangi et al. compared three different scaffold constructions for the differentiation of human umbilical cord-derived mesenchymal stem cells into hepatocyte-like cells: dextranegelatin, chitosanehyaluronic acid, and gelatin-vinyl acetate. They reported better results in hepatocyte differentiation when using those scaffolds made of gelatin, probably owing to its presence in natural ECM composition [34].

Collagens Type I collagen, a crucial component of mammal ECM, is one of the most used materials in tissue engineering, especially for hepatocyte culture and primary hepatocyte transplantation [35]. Risbud et al. published several articles in which they compared the use of hydrogel scaffolds coated with collagen or with chitosan, demonstrating better human umbilical vein endothelial cell attachment and growth in the first ones. They also suggested that collagen hydrogel-coated textile scaffolds represented an excellent system in which hepatocytes and nonparenchymal cells could be cocultured in parallel [36,37].

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Matrigel consists of a complex mixture of ECM proteins (collagen IV), proteoglycans (heparan sulfate), and growth factors (epidermal growth factor and transforming growth factor-b) secreted by EngelbrechteHolmeSwarm mouse sarcoma cells that resemble the ECM of many tissues. Although the main inconvenience of Matrigel is that it is not a well-defined culture condition, many authors use this compound in vitro. For example, Fan et al. combined Matrigel with agarose to construct a hybrid hydrogel to obtain 3D constructs for tissue engineering [38]. Yarmush et al. also developed 2D culture systems that rely on using mixed collagen I and Matrigel sandwiches. These systems showed that primary hepatocyte function was maintained for several weeks without a noticeable loss of function [39].

Alginate Alginate, a hydrophilic porous polysaccharide matrix that allows cell growth and differentiation, is also commonly used. Shteyer et al. [40] suggested that alginate scaffolds improve animal survival after an acute hepatic failure owing to an 87% partial hepatectomy in mice, achieving in vivo hepatocellular function of primary hepatocytes. They compared these results with collagen scaffolds and obtained inferior results. Lin et al. used alginate scaffolds for bone marrowederived mesenchymal stem cell differentiation into hepatocyte-like cells [41]. Tai et al. [42] suggested in vitro differentiation of human mesenchymal stem cells into hepatocyte-like cells using an ArgeGlye Aspemodified chitosanealginate scaffold. This construct was implanted in rats with a 70% partial hepatectomy and used for the in vivo delivery of liver-like cells. Many other related studies describe the power of alginate matrices as scaffolds in liver tissue engineering. They enhance hepatocyte viability, morphology, and function, maintaining hepatocellular performance [43], [44].

Polyglycolic Acid, Polylactic-co-glycolic Acid, and Polycaprolactone Synthetic materials have also been used for this purpose, such as polylactic-co-glycolic acid (PLGA) or polyglycolic acid (PGA). Lees et al. [45] generated 3D scaffolds made of PLGA coated with laminin in which they seeded human embryonic stem cells. These scaffolds were then transplanted into the liver lobules of mice. The researchers observed the formation of teratomas that produced typical proteins of hepatic lineages and also from pancreatic and neuronal lineages. Furthermore, they found the assembly of an ECM rich in laminin, collagen IV, and collagen I, and positive CD34 endothelial progenitors, which represented vasculature formation. Barralet et al. [46] developed a PGA fiber mesh scaffold stabilized with polycaprolactone to study the behavior of human biliary epithelial cells. They compared this technique with alternative models as collagen I sandwiches or Matrigel, showing spheroid aggregation, long-term proliferation (6 months), and phenotypic stability, which would represent a good strategy for bile duct tissue engineering. Other studies [22,47] achieved hepatic tissue generation with primary hepatocyte seeding in this PGA scaffolds, with the capacity of albumin and urea secretion.

Decellularized Extracellular Matrix Decellularization techniques represent a popular technique for generating naturally derived scaffolds. In this process, whole organs are perfused with detergent solutions to remove the cellular contents, preserving an intact ECM and vascular tree. The main advantage of this technique is that there is no need to create an ECM in vitro because the ultrastructure and composition of the native ECM are preserved. In the same way, the vascular network and bile duct system are also well-preserved after liver decellularization [23,24,48]. These scaffolds are then recellularized with different cell types and numbers using different seeding conditions, with the objective of creating functional bioengineered organs able to be transplanted into recipient animals and, in the long-term, into patients with end-stage liver failure. Human, pig, and rat livers are the scaffolds mainly used for liver tissue engineering, but there are several studies in which spleen was used (Fig. 62.2) [49]. Regarding the decellularization process, there are differences reported by several authors. The most commonly used reagents are sodium dodecyl sulfate [50] and Triton X-100 [23,51,52]. Some other examples are enzymes, proteases, ethylenediaminetetraacetic acid, or deoxycholic acid. There are also multiple concentrations and combinations of these reagents [53,54], several methods of detergent perfusion exist, as reported by several researchers. Some use the portal vein [50], others the portal vein and hepatic artery [23,51,55], and others the inferior vena cava [56]. Cell seeding methods are also variable [53,54]; the portal vein and hepatic vein are the most commonly used.

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Dissociaon to single cells

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Liver organoids

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Liver decellularizaon

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FIGURE 62.2 Biomedical applications of liver bioengineered tissues. (A) The resulting single cells isolated from the digestion of human hepatic tissue samples can be seeded in a three-dimensional (3D) environment, originating spheroids and organoids with multiple applications. (B) Cadaveric donor livers can be decellularized and afterward recellularized using the patient’s own cells, with the ultimate goal of obtaining a functional liver for transplantation.

Regarding cell types, primary hepatic cells, both parenchymal (hepatocytes) and nonparenchymal cells (sinusoidal endothelial cells, stellate cells, and Kupffer cells), have been used for liver tissue engineering, as well as pluripotent stem cells, fetal progenitor cells, and immortalized cell lines. Cell number is also a significant variable to take into account [48,57]. To mimic better what is observed in vivo, it is essential to recreate the hepatic spatial distribution (e.g., 80% of the liver parenchyma is composed of hepatocytes). However, there are issues regarding this: not all the cells that comprise the liver structure are easily isolated and maintained in vitro (e.g., hepatocytes do not efficiently expand in vitro, and liver sinusoidal endothelial cells are difficult to obtain) [57]. Furthermore, high cell seeding densities usually provoke cell aggregation, blocking the vascular network perfusion. On the other hand, low cell seeding densities will limit cellecell interactions, limiting cell growth, affecting both cell viability and survival in the scaffold. Taking this into account, different recellularization methods have been developed [48,57,58]. Soto-Gutierrez et al. [48] compared three of them regarding better hepatocyte survival and function: direct parenchymal injection, continuous perfusion, and multistep infusion. The first consists of direct cell injection into different lobes of the liver scaffold. In the second technique, the total amount of cells that are going to be seeded are suspended in the culture media and perfused at a determined flow rate. Finally, the cells are added to the culture media in multiple steps, with time intervals between them, and perfused at a determined flow rate. They determined that the best results were using the multistep perfusion system, which not only allowed hepatocyte engraftment around 90%, but also were able to produce albumin and metabolize ammonia and cytochrome P450 (CYP)1A/2 activity for 7 days in vitro. The flow seeding conditions are also critical. Low flows should be used to allow cell attachment and survival in the scaffold surface. Moreover, every type of cell has its particular cell attachment conditions, with the need to find a particular speed that allows for the most significant cell engraftment as possible, and the delivery into the native spatial location inside a decellularized liver scaffold. Other shortcomings that have to be solved are the long-term cell survival or graft thrombosis after implantation, owing to lack of a functional vascular network in the current generation of bioengineered livers. What can be determined from this is that a lot of relevant issues still need to be addressed in this novel field of liver bioengineering before clinical applications. Although much has been accomplished so far, mimicking in vitro the complexity observed in vivo is still challenging.

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Future studies may lead to a better understanding of the production of human-sized transplantable bioengineered livers able to restore diseased organs or replace failed organs.

LIVER SPHEROIDS, ORGANOIDS, AND AGGREGATES: CANCER, BIOARTIFICIAL LIVER, TRANSPLANTATION RESEARCH, AND TOXICOLOGY AND DRUG DEVELOPMENT Another strategy for the generation of 3D liver tissue constructs is using the hepatocyte capacity to form hepatospheres, which are spontaneous nonadherent spheroid cell aggregations that mimic in vivo cellular environment. The generation of these structures can be achieved by multiple methods such as nonadherent round-bottom wells using methylcellulose-containing medium, rotary and rocked cultures, hanging drops, or 3D matrices [59]. These hepatic structures have revealed higher liver-specific hepatospheres [60,61], which solves the problem of culturing hepatocytes in 2D systems, where dedifferentiation quickly develops. Cocultures with endothelial and other nonparenchymal cells were also developed using this technique (Fig. 62.2A).

Cancer Research Liver spheroids have emerged as a promising tool in cancer research [62,63]. This 3D culture method has been used as an intermediate model between in vitro and in vivo microenvironments [64]. Studies have demonstrated that liver spheroids preserve a more significant number of liver cells functions better than 2D culture, including albumin and urea production and bile secretion [62,63]. Moreover, the 3D culture conserves cell polarization, resulting in higher concordance with in vivo conditions rather than 2D cultures. This is particularly important in cancer research because 3D cultures better resemble the interplay of the tumor and its microenvironment. This culture configuration allows for the formation of a stable tissue in which cellecell and cellematrix interactions are better preserved than in 2D cultures [65]. Hence, liver spheroids more accurately reflect the tumor complexity and heterogeneity and constitute a therapeutically relevant model, displaying pathophysiological gradients of in vivo tumors, such as pH and oxygen gradients, the penetration rate of growth factors, and the distribution of proliferating and necrotic cells [64]. Liver spheroids and aggregates have also been used to evaluate the mechanism of tumor growth, metastasis development, and antitumor therapeutic agents [63,65,66]. Not long ago, unlike other cancers, there was a lack of evidence and data published about the use of liver spheroids in primary liver cancer research. An increase was observed in the number of studies using liver aggregates for in vitro modeling of HCC [67,68]. Song et al. cultivated multicellular tumor spheroids to elucidate the mechanisms of environment-mediated chemoresistance and metastatic transformation in HCC. In that study, the authors created different multicellular spheroids that contained different human HCC lines and stromal cells, including hepatic stellate cells (HSCs). The authors observed that the presence of activated HSCs increased the compactness of spheroids and enhanced resistance to chemotherapeutic agents such as sorafenib and cisplatin compared with other stromal cells. The authors also observed that this increase in drug resistance was related to HSC production of collagen 1A1. On the other hand, the results of the study showed that activation of HSC promoted HCC migration by upregulating matrix metalloproteinase 9. Globally, the authors demonstrated that interaction between HCC cells and HSC induced chemoresistance and gave the tumor a more aggressive and invasive profile; they suggested that HSC could be the target for novel therapeutic strategies [68]. In the same way, Chen et al. used HCC spheroids to screen candidate metastasis-associated genes. The authors built two different HCC spheroids using two different cell lines, one with high-metastatic (MHCC97H) behavior and other with a less aggressive profile (Hep3B). Then the authors evaluated the expression of candidate metastasis-associated genes related to cell adhesion, matrix secretion, and invasion. The results showed an evident change in the gene expression of 123 genes between the two spheroids. In the MHCC97H spheroids, the authors found that the number of upregulated genes related to adhesion molecules mediating cellematrix interactions and matrix secretion was significantly higher. In contrast, the Hep3B spheroid presented an increased expression of adhesion molecules maintaining cell adhesion. With these data, the authors could identify a specific gene expression pattern that may determine the malignant phenotype of the HCC spheroid [67]. The findings of these studies represent the potential of the liver spheroids in the field of cancer research. 3D culture models offer a more accurate environment of in vivo microarchitecture to study tumor growth and progression.

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In the future, the introduction of new biomaterials and biofabrication techniques will also reproduce the interaction between cells and matrices more precisely. For these reasons, the results obtained from studies performed with 3D culture methods are more consistent owing to their similitude with the in vivo reality. A strong implication of liver spheroids in cancer research is their potential in personalized medicine. Future studies in patients with primary liver cancer or liver metastasis may use host-tumor based spheroids to evaluate pharmacologic activity and toxicity.

Bioartificial Liver and Transplantation Research The liver is a vital organ with complex functions including gluconeogenesis, albumin and urea synthesis, lipid metabolism, drug detoxification, waste removal, and immune and hormonal modulation. Liver diseases are prevalent worldwide; most cases lead to progressive impairment of liver function. The only available treatment for lifethreatening conditions such as end-stage liver disease and acute liver failure is LT, but because of donor shortage and an increase in the number of potential LT candidates, alternative measures are needed [69]. Owing to their ability to resemble in vivo structures and preserve organ-specific functions, liver spheroids have also been used in bioartificial livers and transplantation research [62,63]. Liver spheroids have a potential use in this field, directly transplanted or employed in hybrid and bioartificial liver support systems (Fig. 62.2) [70e72]. In 2005, Gan et al. explored the role of a spheroid-based hybrid artificial liver support system (HALSS) for treatment of severe liver failure. In this study, the HALSS consisted of a plasma separator, a hemoabsorbant, and a bioreactor with pig-derived hepatocyte spheroids in its extrafiber space. This system has been used in 10 patients with severe liver failure. The authors observed that after 24 h of spheroid culture, they produced higher levels of albumin and urea in respect to 2D cultures, suggesting better preservation of cell function. The authors also observed that patients treated with HALSS had a better survival rate (30% versus 0%), low levels of total bilirubin and alanine transferase, and improvement in coagulation compared with those in the control group. With these results, the authors concluded that spheroid-based HALSS could be useful in supporting liver function in patients with severe liver failure [71]. A preclinical study by Glorioso et al. evaluated the neuroprotective effect of the spheroid reservoir bioartificial liver (SRBAL) in a porcine model of acute liver failure. In that study, 18 pigs with an acute liver failure caused by a drug overdose were randomized into three groups: standard therapy, standard therapy plus no cell liver support device, and standard therapy plus the SRBAL device. After 90 h of treatment, the authors observed that pigs treated with the SRBAL improved survival compared with standard therapy and no cell device therapy (83%, 0%, and 17%, respectively) and had significant improvement in the rate of ammonia detoxification, peak levels of serum ammonia, and intracranial pressure. The authors also found that hepatocyte spheroids remained highly functional and that survival and device function were directly related to hepatocyte cell dose, treatment duration, and membrane pore size [72]. These studies reflect the potential of the spheroid-based bioartificial liver as a promising treatment for liver failure. Spheroid aggregates offer a readily available source of donor cells and reach the minimal cell dose to obtain a significant clinical effect in animals and humans. Regardless, randomized clinical studies are needed to determine whether these bioartificial support systems are safe and feasible in individuals with acute liver failure, with the aim of expanding its use to clinical practice and avoiding LT. Other studies have evaluated the effect of direct transplantation of liver spheroids. No controlled trial has yet evaluated the clinical benefit of liver cell therapy in liver diseases in humans, and slightly more than 100 patients have been treated with hepatocyte transplantation [73]. However, the benefit is well-known in animal models. In 2002, Hamazaki et al. demonstrated that intraperitoneal transplantation of microencapsulated multicellular spheroid of rat hepatocytes improved survival in acute liver failure induced by 90% hepatectomy. Nevertheless, data are needed on liver spheroids transplantation in humans [70]. Results from experience with hepatocyte transplantation in humans and with liver spheroids in animals are promising, and cell-based therapies emerged as a “bridge to transplant” for both metabolic liver disease and liver failure. The most significant limiting factor is that sources of hepatocytes for transplant are scarce and tissues that are available come from organs that are unsuitable for LT [74]. To avoid this problem, alternative sources of cells have been evaluated. Pettinato et al. evaluated the differentiation of human induced pluripotent stem cells (hiPSCs) into hepatocyte-like cells through direct Wnteb-catenin pathway inhibition [75]. The authors developed a new multicellular spheroid-based hepatic differentiation protocol starting from hiPSCs. The resultant spheroids were transplanted into a rat model of acute liver failure. The results showed that transplantation significantly prolonged the mean survival and ameliorated the liver function compared with the control group. These results indicated that this differentiation program is feasible and highly efficient and suggested that the differentiation of stem cells

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may provide an alternative source for hepatocyte-like cells with comparable differentiation and function [75]. Once again, these results confirm the power of 3D culture to preserve liver-specific functions throughout time. Finally, liver spheroid transplantation could also have a role in personalized gene therapy. A study suggested that transplantation of spheroids of cells after genetic modification with nonviral DNA vectors may increase the therapeutic potential of this technique. In their work, Uchida et al. evaluated this hypothesis [76]. The authors transplanted subcutaneously hepatocyte spheroids transfected with erythropoietin DNA in rats. They observed that the spheroid-treated group showed a significantly higher hematopoietic effect than animals transplanted with a cell suspension from monolayer cultures. Moreover, the spheroid system contributed to the preservation of native functions of hepatocytes in the host tissue, confirming once again that 3D culture delays dedifferentiation and loss of functions of these cells [76]. These results introduced a new perspective for liver spheroid transplantation. In the future, genetically modified spheroid transplantation systems may be used to overcome genetic diseases. The actual state of the art suggests that we may soon be able to create and transplant genetically modified spheroids of hepatocyte-like cells derived from the patient’s own stem cells, which will represent a revolution in personalized medicine.

Toxicology and Drug Development The liver is the organ where most drugs are metabolized and transformed into metabolites or active compounds. These substances by-products may become toxic to the liver itself and the rest of the body, a phenomenon known as drug-induced liver injury (DILI). To launch a single drug into the market is a long (12e15 years) and costly ($3e5 billion) process [77]. After an initial screening, lead candidate compounds are characterized in vitro and in vivo for their absorption, distribution, metabolism, excretion, and toxicity properties before proceeding with clinical trials. However, most compounds (>90%) fail during these final stages: 43% of these failures occur owing to a lack of efficacy and 33% to the appearance of adverse effects [78], mainly because of DILI [79]. Hence, drug withdrawals at clinical stages in humans mainly result from the use of inappropriate or inaccurate in vitro and in vivo liver models in the course of drug studies. On the other hand, the liver is the target organ of some prevalent diseases, such as infectious hepatitis B virus, HCV [80], malaria [81], overnutrition-induced (type 2 diabetes, nonalcoholic fatty liver disease, fibrosis, and cirrhosis) [82e84] or tumoral diseases (HCC represents the sixth most common cancer worldwide) [85]. Considering all of this, liver models’ results are necessary for the development of novel drugs, not only for the study of xenobiotics’ metabolism and toxicity but also for the development of specific drugs for liver diseases. Hence, more realistic in vitro human liver models are needed that resemble as closely as possible in vivo liver structure, physiology, and the pharmacological response.

Limitations of Current In Vitro Liver Models to Test Drugs As mentioned before, maintaining liver parenchymal function ex vivo is essential to generating stable systems for efficacy and toxicology drug studies, so fully functional hepatocytes are needed. For that, the 3D relationship of cells within the differential microenvironments of the liver (e.g., periportal versus pericentral), the regional hemodynamic flow patterns, and other physiological factors such as oxygen tension and cytokine profiles have to be simulated in vitro. However, cell-based models that are routinely used in drug testing are simple monoculture systems (typically standard microtiter plate formats) employed under static, nonphysiologic 2D conditions, which make them suboptimal models for drug efficacy and safety testing, unable to mimic or predict more complex mechanisms of action [86]. Hepatocyte viability in suspension decreases significantly after 4 h. Because of that, for years cryopreserved human hepatocytes in monolayer cultures have been the reference standard to test drug metabolism and toxicity [87]. However, cryopreservation also reduced hepatocyte viability, and function and their culture in monolayer downregulates cell receptors involved in cellecell and celleECM interactions, drastically reducing cell functionality over time [87]. The development of 2D culture models such as sandwich culture increased basal and induced drugmetabolizing enzyme activities and simulated in vivo biliary excretion rates [87,88]. However, dedifferentiation of hepatocytes in long-term cultures and the lack of nonparenchymal cells that interact with hepatocytes continued to be inherent disadvantages of these models [89]. The coculture of hepatocytes with other liver cells, such as stellate cells, Kupffer cells, liver sinusoidal endothelial cells, or liver epithelial cells diminishes these limitations to some extent, improving the longevity and functionality of cells and producing the higher expression of CYP and phase

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II isoforms compared with monotypic culture [90e93]. Nevertheless, cocultures are usually based on the random mixing of different cell types and thus do not account for their particular anatomical relationship. Looking for more relevant models, to emulate the 3D organization and morphology of hepatocytes within the liver, 3D cultures have been developed. 3D cultures range in complexity from monotypic or heterotypic spheroids [94,95] to 3D scaffold systems [96] and more advanced models using microfluidic in vitro systems [96,97]. Multiple commercial 3D coculture platforms have been developed for drug screening and drug studies such as the “HepatoPac” platform [98], the 3D InSight Human Liver Microtissues of InSphero, the HepaChip in vitro microfluidic system [99], and the HmREL microliver platforms [100]. Although some issues have been addressed for specific applications with these models, others continue to be biologically and technically challenging [88,101].

Organoids in Drug Development Organoids represent more complex models that try to simulate 3D cellecell and celleECM relationships in more relevant physiological conditions that mimic the liver microenvironment arrangement amenable to the highthroughput screening of compounds feasible enough to guarantee long-term studies. The optimal liver role depends on the coordinated function of parenchymal and nonparenchymal cells within the hepatic acinus as well as hepatic blood microcirculation. Aspects of the microcirculation can be simulated in vitro via perfusion models to create a dynamic in vivo like environment. Different macroscopic perfused in vitro liver systems initially developed as bioartificial liver devices have been created [97], which provide evidence that perfusion can improve longevity and function in sophisticated hepatic systems, and thus show better in vivo mimicry. Although these models represent the most physiologically practical systems, their size makes them unfeasible for use in drug testing studies because they lack the throughput and analytical flexibility for drug screening. The use of these organoids in drug development involves their miniaturization to a microscopic level. This new class of in vitro tools, often called “on-a-chip" tissue models, can mimic the architecture of small tissue sections and individual characteristics of the dynamic in vivo flow environment, while offering more precise spatial and temporal control of soluble factors. Apart from being amendable to high-throughput screening approaches, these models can be engineered for real-time monitoring of the state of cells and their extracellular environment, which is crucial for determining cellular mechanisms of action in drugs [97]. Several organoid systems have been developed for drug screening and testing. One of the best examples of these organoids is in microfluidic systems. In 2006, Kane et al. developed a microfluidic coculture system of hepatocytes and T3-J2 fibroblasts in an 8  8-well array, demonstrating stable albumin and urea excretion for 32 days. Some years later, HmREL Corporation developed a similar microfluidic in vitro liver platform for drug screening for commercial purposes (HmRELflow) [100]. This platform, formed by multiple fluidically interconnected microscale cell culture compartments, enables simulation of the interaction of test substrates with two or more organs, which provides an enhanced prediction of human response. In fact, in vivo-like absorption, distribution, metabolism, bioaccumulation, and toxicity of naphthalene were demonstrated when lung, adipose, and liver cells were fluidically connected [100]. Furthermore, the size of the system enabled microscopic imaging, oxygen sensing, physiologically appropriate ratios of chamber sizes, hydrodynamic shear stress, and less consumption of media and cells. Even so, some issues such as sample removal, complexity in maintaining recirculation, and cell monolayers on chips and not physiological tissue constructs, significantly limit the model. Some years later, Au et al. developed another microfluidic model, a microfluidic organoid for drug screening (MODS) platform [102]. The novelty of this system compared with the previously developed MODS was the ability to evaluate different conditions simultaneously and the automation of time-consuming processes such as the generation of mixtures and the formation of serial dilution series, which can result in more efficient screening of lead drug candidates. Vernetti et al. developed and characterized a sophisticated system to investigate drug safety and efficacy in liver models of disease. This system includes a human 3D, microfluidic, four-cell, sequentially layered, self-assembly liver model as well as fluorescent protein biosensors for mechanistic readouts and a microphysiology system database to manage, analyze, and model data [103]. Hollow-fiber reactors have also been adapted to drug testing. In 2010, Schelzer et al. developed a microscale prototype of a hollow-fiber reactor. In this model, the bioreactor consisted of four cell chambers, each of which included four compartments (one for cells, two for culture medium, and the last for oxygen supply) connected to provide the cells with a physiologically based environment [104]. The prototype allowed for small numbers of cells and limited reagent use, microscopic evaluation of the cells, and monitoring of oxygen concentrations. Later, a similar system

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with coculture of parenchymal and nonparenchymal liver cells was developed for studies of pharmacokinetics and drug toxicity, showing maintained albumin synthesis and CYP activity for 2e3 weeks [105]. Nevertheless, some limitations arise in this kind of systems, such as the lack of physiologic gradients typically seen in liver tissue, the complexity of many tubing lines, and the limited throughput, because only a few different conditions can be assessed simultaneously. As also mentioned previously, decellularization constitutes a novel approach in liver models [23,25]. This macroscopic model that can be used to investigate liver development and regeneration can also be miniaturized for highthroughput drug studies [106]. Apart from physiological models, organotypic models of liver diseases are being developed for drug testing. Drug metabolism, toxicity, and efficacy in diseased livers differ substantially compared with healthy conditions, so accurate models of disease are required. In this sense, Skardal et al. developed liver-based cell organoids in a rotating-wall vessel bioreactor that inoculated with colon carcinoma cells to generate liver-tumor organoids for in vitro modeling of liver metastasis [65]. Leite et al. developed hepatic organoids with fibrotic features such as HSC activation and collagen secretion and deposition to study drug-induced liver fibrosis [107]. Similarly, Lee et al. generated a reversible- and irreversible-injured alcoholic liver disease model in spheroid-based microfluidic chips in which rat primary hepatocytes and HSCs are cocultured [108]. Although enormous advances have been made to develop more realistic and predictive in vitro liver models for drug testing, the field is still dawning. There are critical issues that should be solved for the field to move forward. Standardizing model and platform characterizations for drug-based studies (viability, secretory capacity, enzymatic and toxicology activities, and drug transporter activity) should be established for each model. Building specificity and sensitivity of the systems, recreating more accurately parenchyma zonation, developing better detection systems and better materials [97], and finding new, unlimited, fully functional cell sources [109] are some challenges in developing in vitro liver models for drug studies.

CONCLUSIONS AND FINAL PERSPECTIVES This chapter has focused on strategies of liver tissue engineering with potential for developing innovative treatments and efficient models for studying the liver. Knowledge of regenerative biology and medicine has increased exponentially, giving new hope to the development of effective treatments for liver disease. Much work still lies ahead to obtain final therapies or models that accurately represent the liver to its fullest. Standardization of cell isolation protocols and media formulation is needed to achieve a high grade of reproducibility of results among laboratories around the world. Despite the obstacles, some 2D liver culture models, especially cocultures with hepatocytes and nonparenchymal cells (which better represent the in vivo microenvironment of the liver) can be considered useful models for studying the acute phase response, mutagenesis, xenobiotic toxicity, lipid, and drug metabolism in the liver. Liver organoids are also considered a possible tool to assess cell changes that lead to tumorigenesis and cancer progression, and for drug screening and testing. Improvement and standardization of the protocols used are needed to enhance the production rate and ameliorate the features of the obtained organoids. Regarding the bioreactor systems used in liver tissue engineering, clinical assessment of the efficacy and safety of bioartificial liver systems in treating different end-stage liver diseases is near completion. Then, the field can concentrate on reducing the costs, finding more suitable cell sources, and optimizing bioreactor technologies. Finally, the use of whole-organ liver scaffolds is promising in the quest of bioengineering a whole liver for transplantation. Livers from different species have been decellularized using various protocols to obtain acellular bioscaffolds. Nonetheless, complete recellularization using all necessary liver cell types, including Kupffer, sinusoidal endothelial, and stellate cells, and the generation of a fully functional liver have not yet been accomplished. Hence, one of the most significant challenges in whole-organ liver bioengineering is an appropriate cell source to repopulate the acellular scaffold and achieve well-defined complete revascularization of the decellularized liver. Induced pluripotent stem cell (iPSC) technology has the potential to provide a source of cells for whole-liver bioengineering for humans. However, bioengineering a fully functional organ with a size comparable to that of humans has yet to be achieved by using iPSC technology. Altogether, the strategies for liver tissue engineering discussed in this chapter may have an impact on innovative personalized tissue engineering-based treatments as alternative and effective therapies for different liver diseases or pathologies.

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Acknowledgments This work was supported by Instituto de Salud Carlos III through a predoctoral fellowship, i-PFIS IFI15/00158 (I. PeP), Gobierno de Aragon and Fondo Social Europeo, through a predoctoral fellowship DGA C066/2014 (P. S-A), and by two predoctoral fellowships from Fundac¸a˜o para Cieˆncia e Tecnologia, Portugal (PD/BD/114057/2015) (S.M.). NSR was supported by a POCTEFA/RefBio II research grant. PMB was supported by the PI15/00563 Research Project from Instituto de Salud Carlos III, Madrid, Spain.

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C H A P T E R

63 Regenerative Medicine in the Cornea Fiona Simpson1, Emilio I. Alarcon2, Jo¨ns Hilborn3, Isabelle Brunette1,4, May Griffith1,4 1

Maisonneuve-Rosemont Hospital Research Centre, Montreal, QC, Canada; 2University of Ottawa Heart Institute, Ottawa, ON, Canada; 3Uppsala University, Uppsala, Sweden; 4University of Montreal, Montreal, QC, Canada

INTRODUCTION Structure and Function of the Cornea The cornea is the transparent front covering of the eye that transmits and focuses light into the eye for vision. It is composed of three main cellular layers: an outermost stratified epithelium, a middle stroma, and an innermost endothelial layer. The total thickness is approximately 550 mm in the center and 750 mm at the periphery [1]. The epithelium of the cornea forms the outer layer and primary protective barrier. It is composed of stratifying, nonkeratinized epithelial cells. Epithelial cells also secrete antiinflammatory and antimicrobial factors as an insoluble layer that maintains the tear film [2]. Because cells are lost on the anterior surface of the cornea, new layers of epithelium originate from a basal layer generated by corneal stem cells in the corneoescleral limbus [3]. The stroma is composed of a hydrated, largely collagenous extracellular matrix (ECM) containing a network of fibroblast-like cells called keratocytes. The single endothelial layer actively osmoregulates the entire structure and maintains hydration while pumping out excess fluid. The cornea is avascular and relies on its extensive network of sensory neurons and their interaction with the corneal cells to maintain tissue integrity and heal wounds [4]. The external location of the cornea makes it prone to injury and infection. Diseases of the cornea are the fourth largest cause of vision loss globally according to the World Health Organization [5]. Corneal blindness is treated by transplantation with donated human corneas, but there is a severe worldwide shortage of goodquality corneas. Not taking into consideration patients in poorly served remote areas or those contraindicated for conventional surgery, an estimated 12.7 million patients are awaiting corneal transplantation globally [6]. This need is particularly elevated in low- to middle-income countries, e.g., China and India, where 2 million and 7 million patients, respectively, are awaiting surgery.

Treatment Options, State of the Art, and Need for Corneal Regenerative Medicine The only widespread treatment is corneal transplantation using a full-thickness human donor cornea through a procedure known as penetrating keratoplasty (PK). PK therefore relies on the ability to access donor corneas via eye banks. In 2012, there were 184,576 corneal transplants in 116 countries globally [6]. This means that of 70 patients needing transplantation only 1 is treated, and this is mainly because of a severe global shortage of human donor corneas. However, even with adequate access to donated tissues, although initial graft success rates are high in developed nations, e.g., 90% over 2 years in Sweden [7], only about 64% of grafts survive over 10 years [8]. In developing nations where there is a preponderance of severe pathologies that place patients at high risk for rejection, the graft survival rate over 2 years was 52% [9].

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Corneal prostheses known as keratoprostheses (KPros) have been developed mainly as alternatives to donor corneas for high-risk grafts, because these patients are generally contraindicated or not prioritized for conventional donor transplantation in countries where there is a severe donor shortage. The traditional KPro that is used clinically is composed of an optical core, often made from poly(methyl methacrylate) and a skirt that interfaces with the patient’s eye. The two most widely used KPros (the Boston KPro and osteoodonto-keratoprosthesis) and their variants contain a biological interface. However, the irreversible nature of the surgery, potential complications, and the need for sustained antibiotics and immune suppression cause them to remain an option for end-stage eyes, so they are not ideal replacements for donor human corneas [10]. For more detailed reviews of KPros, see Avadhanam et al. [10]and Salvador-Culla et al. [11]. Regeneration of the corneal epithelium from cell-culture expanded limbal epithelial stem cells pioneered by Pelligrini et al. was a major breakthrough in modern corneal blindness treatment and is in clinical application as the first European Commission approved cell-based advanced therapy medicinal product (ATMP) [12,13]. A direct method for limbal cell replacement known as simple limbal epithelial transplantation (SLET) that bypasses stem cell expansion has also entered clinical application [14,15]. However, limbal stem transplantation is most successful when using an autograft, and in cases where the pathology involves deeper layers, follow-on transplantation with a donated cornea is still needed [16]. The first successful regeneration of the cornea using a cell-free method was reported by Griffith and colleagues and resulted from the use of recombinant human collagen (RHCIII) hydrogels [17]. Ten patients underwent the implantation of proregeneration scaffolds by anterior lamellar keratoplasty [17]. Patients were given partialthickness implants composed of RHCIII cross-linked using 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimidee N-hydroxysuccinimide, resulting in the regeneration of the epithelium, stroma, and corneal nerves (Fig. 63.1). Because the neocorneas were derived from the patients’ own cells growing into cell-free scaffolds, there were no immune problems and patients did not require immune suppression beyond the 6e7 weeks of local steroid eye drops prescribed prophylactically. Regenerative medicine remains the main option for restoring the structure and function of the cornea to the semblance of a normal cornea. In this chapter, we review current methods used to regenerate the human cornea. Because of the wide variety of methods used, we will review only examples that are in clinical evaluation and are being evaluated preclinically, as well as some novel methods.

REGENERATIVE MEDICINE APPLIED TO KERATOPROSTHESIS DEVELOPMENT KPros under development and clinical evaluation incorporate elements allowing regeneration. Specifically, devices have been designed to allow the regrowth of corneal epithelium to cover the device to maintain the tear film and prevent infection or extrusion of the implant. This is done by surface modification of the interface opposing the patient’s corneal tissues with either surface chemistries or lithographic techniques that pattern the surface on a nanoscale to microscale. The macromolecules of ECM proteins such as collagen, fibronectin, laminin, their derivative cell adhesive peptides (e.g., isoleucineelysineealanineevaline, TyreIleeGlyeSereArg [YIGSR], and ArgeGlyeAsp [RGD]), or other ECM analogs have been evaluated for their ability to modulate epithelial growth over KPros. Previous work has shown that corneal epithelial cell adhesion and growth were significantly enhanced by tethering of laminin or fibronectin adhesion promoting peptide via flexible polyethylene glycol (PEG) chains, more so than by tethering of fibronectin or simple coating of the surface with matrix proteins [18,19]. Modification with fibronectin-based ArgeGlyeAspeSer (RGDS) [20e22], laminin-based YIGSR [23,24], and a novel collagenbased peptide, GlyeProeLeu [25], was also shown to improve epithelial cell adhesion in vitro. Surface modification peptide combinations, e.g., well-known cell adhesion peptides RGDS and YIGSR together with their synergistic counterparts ProeHiseSereArgeAsn and ProeAspeSereGlyeArg, have resulted in surfaces that provide improved corneal epithelial cell adhesion and growth [23]. Growth factors such as epidermal growth factor or insulin-like growth factor-1 have been tethered onto polydimethylsiloxane and polymethacrylic acid-co-2-hydroxyethyl methacrylate, respectively [18,26]. Both studies showed that the growth factors enhanced corneal epithelial cell attachment and growth and that the use of a PEG tether has been shown to improve cell coverage of the polymers significantly in vitro. Both sets of results strongly suggested that the spacer molecules provided the correct microenvironment for the epithelial cells by exposing the bioactive motifs to allow the cells to reach confluence, compared with little or no epithelial growth on the surfaces that were coated only with the bioactive factors.

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FIGURE 63.1 Corneal features in a healthy, unoperated subject, alongside those of operated patients, at 24 months after implantation of a biosynthetic cornea or a human donor cornea. (Top row) Anterior segment optical coherence tomography (ASOCT) images of a healthy cornea, biosynthetic implant, and human donor transplant by penetrating keratoplasty. Areas of wound-healing activity exhibit high reflectivity (white areas). (AeO) In vivo confocal microscopy (IVCM) images. Intact epithelium of the unoperated cornea (A), regenerated corneal epithelial cells on the implant surface (B), and regenerated epithelium of the penetrating graft (C). Regenerated nerves (E) at the subbasal epithelium in an implanted cornea were parallel and morphologically similar to the normal cornea (D), whereas regenerated subbasal nerves were also observed in a cornea transplanted with human donor tissue (F). Anterior stromal cell (keratocyte) nuclei (GeI) and posterior keratocytes (JeL) were present, with varying density, in all corneas. The endothelium (MeO) in all corneas exhibited a characteristic mosaic pattern. Reproduced from Fagerholm P, Lagali NS, Merrett K, Jackson WB, Munger R, Liu Y, et al. A biosynthetic alternative to human donor tissue for inducing corneal regeneration: 24 month follow-up of a Phase I clinical study. Science Transl Med 2010; 2: 46ra61, courtesy of AAAS.

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Modification of the physical profile of a KPro surface by lithography is another method to enhance the celletissue device interface. Surface patterning of a KPro was first reported by Myung et al. using a photolithographically patterned device composed of a PEGepoly(acrylic acid) (PAA) central core and a poly(hydroxyethyl acrylate) microperforated skirt. Type I collagen was then coupled to the hydrogel using photochemical surface modification to promote epithelial migration into the wound site in rabbit models in vitro and in vivo [27]. When the collagen-patterned PEGePAA hydrogels were implanted into rabbits, the result was good tolerance in 9 of 10 subjects but an optical haze remained at the surgical site [28]. An additional study suggested that the collagen was evenly distributed but cell migration was slower than in surgical controls and took 5 instead of 2.5 days for wound closure [29]. A comparative study of PEG-diacrylateePAA versus PEG-diacrylamideePAA hydrogels indicated that the PEG-diacrylamide was more biocompatible than the original formulation [30]. A 2015 study compared porous versus nonporous PEG-diacrylamideePAA hydrogels with surface modification with type I collagen and fibronectin [31]. The porous hydrogels were synthesized using templates composed of polystyrene microbeads, which were removed after cross-linking by washing in methyl ethyl ketone buffer. Both the porous and nonporous formulations supported the multilayered growth of corneal fibroblasts, which suggested the suitability of porous PEGePAA hydrogels for KPro skirt material.

REGENERATION OF CORNEAL LAYERS Corneal Epithelium Work in regeneration of the corneal epithelium has focused on methods that circumvent the rejection of allogeneic donor cells, because this is a huge problem that necessitates systemic immunosuppression even in patients receiving human leukocyte antigenematched tissue, which has its own risks and side effects [32]. This includes the development of new therapeutic stem cell sources from transdifferentiation of progenitor cells from other sites such as oral mucosa (or other mucous membranes) to conversion of fibroblasts or mesenchymal stromal/stem cells (MSCs) into corneal lineages by induced pluripotency [33]. Transdifferentiation of oral mucosa has been tested in patients with bilateral corneal epithelial stem cell depletion. In 2004, Nishida et al. reported the first successful human cultivated oral mucosal epithelial transplantation surgeries [34]. The procedure was conducted in four participants with bilateral total limbal stem cell deficiency caused by StevenseJohnson syndrome (SJS) or ocular cicatricial pemphigoid. The oral mucosal cells were cultured on temperature-responsive cell-culture surfaces with 3T3 feeder cells and transplanted onto one eye of each patient, with successful, stable outcomes. However, all transplanted eyes had some peripheral corneal neovascularization. The procedure was later improved by culture of oral mucosa on human amniotic membrane (HAM) [35]. This technique has been successfully used in patients with SJS, alkali and temperature burns, and aniridia [36e40]. Other stem cells such as umbilical cord lining-derived stem cells have been successfully tested in rabbit models for use in regenerating the corneal surface [41]. MSCs have also been tested as potential sources of corneal limbal epithelial stem cell replacements in rat and rabbit models, but results have been mixed [42e44]. A range of biomaterials have been employed as substrates for corneal epithelial cells for use as implants. Fibrin and HAMs have been the staples. However, materials such as collagen membranes and silk fibroin are being evaluated. Electrospun poly(D,L-lactide-co-glycolide) (PLGA) has also been used as substrate to expand cultured limbal cells and as a human limbal cell explant to restore the stem cell niche [45,46]. PLGA was used at a ratio of 50:50 lactic acid to glycolic acid. The electrospun meshes underwent g-sterilization and dry storage at 20 C. The meshes degraded at 4e6 weeks in vitro and allowed for the culture of confluent limbal cells. Both cultured and explanted cells generated differentiated and stem cell populations when transplanted into an ex vivo rabbit corneal model.

Corneal Endothelium The expansion of primary corneal endothelial cells (CECs) in culture before reinjection has been successfully tested in a feline model. CECs were obtained from Descemet membranes excised from feline corneas and expanded for up to two passages after cell detachment. Sixteen animals underwent surgery of the right eye. Eight animals underwent 7-mm central endothelial scraping and injection with 2  105 or 1  106 cultured CECs, and 100 or 350 mM Rho-associated, coiled-coil containing protein kinase (ROCK) inhibitor (ROCKi). Two underwent

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18-mm total scraping followed by intracameral injection of 1 M CECs and ROCKi in 0.2 mL. Six negative controls underwent central or total scraping followed by injection of ROCKi. After surgery, the corneas grafted with CECs performed better than the surgical controls but the endothelium had incomplete functionality and was inferior to the unoperated corneas. CECs have been differentiated from human embryonic stem cells (hESCs) but in vivo analysis has thus far been limited to animal models. In 2014, Zhang et al. reported human corneal endothelial-like cells derived from the periocular mesenchymal precursor (POMP) phase of hESCs [47]. POMP cells were transferred into coculture with human corneal stromal cells in lens epithelial celleconditioned media for differentiation into CEC-like cells. The CEC-like cells were sorted using fluorescence activated cell sorting and cultured to generate CEC-like sheets. Transplantation into rabbits undergoing Descemet membrane stripping resulted in gradual restoration of transparency. Two methods of deriving CECs from the neural crest (NC) phase of hESCs have been tested. McCabe et al. derived CEC-like cells from dual Smad inhibitoregenerated, feeder-free hESC NC cells with platelet-derived growth factor-b, and Dickkopf-related protein 2 [48]. Song et al. derived CEC-like cells from feeder-supported hESCs differentiated into NC cells cultured in bovine CEC conditional medium and fresh CEC media [49].

Corneal Stroma Funderburgh and colleagues worked on stroma regeneration and showed that direct injection of stem cells isolated from limbal biopsies can prevent corneal scarring [50]. Limbal stem cell biopsy-derived stromal cells were derived from mesenchymal cells from human cadaveric cornea-sclera rims. In culture, these cells expanded to produce a thick lamellar structure with aligned collagen and proteoglycans resembling the ECM. Grafts in a mouse corneal wound model prevented the formation of scar tissue and were well-tolerated. Corneal stroma sheets have also been constructed in vitro for use as potential replacements for pathologic or thinned stromas. These are discussed subsequently in a discussion on self-assembled corneal constructs.

FULLY CELL-BASED, SELF-ASSEMBLED CORNEAL CONSTRUCTS Self-assembly of corneal stromal components, largely collagens, has been extensively examined by Germain and colleagues, who used the method of ascorbic acid stimulation of fibroblastic cells to secrete ascorbic acid to prepare a range of tissues. In earlier work, Proulx et al. constructed a corneal stroma by growing stromal fibroblasts in cell culture medium supplemented with ascorbic acid for 28e35 days [51]. The ascorbic acid induces fibroblasts to lay down collagen and other ECM components to form sheets. These 35- to 55-mm-thick sheets are then stacked to form a corneal stroma. After fabrication of a stacked stroma, Proulx et al. added an endothelium and epithelium to construct a three-layered cornea. First, CECs were seeded on top of the reconstructed stroma and allowed to grow into a monolayer in endothelial growth medium. After 2e7 days, a plastic ring was placed on top of the cultured construct that was then turned upside down. Corneal limbal epithelial cells are then seeded on top, after which the entire construct is cultured in epithelial growth medium supplemented with more ascorbic acid. After achieving epithelial confluence, the entire construct is air-lifted and cultured at the aireliquid interface to promote stratification of the corneal epithelium. The resulting cornea was significantly thinner than a native human cornea. However, it was transparent and the authors suggested that full thickness might be achieved by increasing the number of stacked layers of fibroblasteECM. A variation on the use of ascorbic acid was reported by Karamichos et al. [52], who employed ascorbic acid with and without transforming growth factor-b to induce elaboration of ECM by human umbilical cord MSCs (cord stem cells) to form a cornea stroma-like structure. The cord stem cells themselves differentiated into stroma-like cells. As reported for the corneal stromal fibroblasts, it took 4 weeks to produce a construct 24 mm in diameter and 30 mm thick. However, in the cord stem cell matrices, 50% fewer cells than fibroblast-produced matrices were present, which showed that the ECM laid down by the cord stem cells was denser. Self-assembled corneal stromal constructs were tested in animals. The iteration of self-assembled corneal stromal was reported by Syed-Picard et al. [53]. Human cornea stromal stem cells were seeded onto micropatterned PDMS substrates supplemented with stem cell growth medium. After 48 h, the medium was switched to a keratocyte differentiation medium containing ascorbic acid, fibroblast growth factor-2, and transforming growth factor-b3. After 10 days, a sheet was produced that could be detached from the PDMS substrate. After 5 weeks of implantation into stromal pockets in the corneas of immune-competent mice, the sheets were

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well-integrated with persistence of human donor cells. They had also acquired transparency during the implantation period (Fig. 63.2). Thicker sheets can be produced using a similar method, by stacking of self-assembled stromal constructs. Corneal stromal constructs were generated from both feline and human donor corneas and six sheets of stromal cells were stacked to generate each graft, which was then implanted into the corneas of cats in stromal pockets [54]. The animals each received two 4-mm-long and 300-mM-long incisions poles apart in the superior nasal and inferior temporal quadrants, after which a graft was placed in each intrastromal pocket. Four animals received human xenografts and four animals received feline allografts. Contralateral eyes underwent surgery without grafting. Preoperatively, the grafts were slightly hazy but they were restored to complete transparency after day 37. The grafts remained transparent and avascular with reinnervation and showed normal corneal sensitivity (Fig. 63.3). Overall, the surgery was well-tolerated with no symptoms of immune rejection.

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FIGURE 63.2 Formation of scaffold-free tissue sheet with parallel cell and matrix organization. Light micrographs of (A) top view and (B)

cross-sectional view of the polydimethylsiloxane substrate show grooves approximately 10 mm wide, 10 mm apart, and 5 mm deep. (C) Phase contrast image shows corneal stromal stem cells (CSSC) cultured on the grooved substrate. (D) For better visualization, CSSC were labeled with DiI (red) and cultured on grooved substrate. (E) Two-photon micrograph of 10-day cultures of CSSC on grooved substrates in keratocyte differentiation medium shows deposition of parallel organized collagenous matrix (green). Nuclei (blue) were stained by SYTOX-green (blue). (F) After 10 days of culture, a robust tissue sheet is formed that can be separated from the substrate using forceps. Scale bars: (A) and (B) ¼ 50 mm; (C)e(E) ¼ 100 mm. Reproduced from Syed-Picard FNN, Du YY, Hertsenberg AJJ, Palchesko RR, Funderburgh MLL, Feinberg AWW, et al. Scaffold-free tissue engineering of functional corneal stromal tissue. J Tissue Eng Regen Med 2016;12(1):59e69, with permission from Wiley and J. Funderburgh.

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FIGURE 63.3 Clinical outcome of the tissue engineered (TE)eintrastromal grafts. (A) Representative slit-lamp photos at 4 days (column 1), 1 month (column 2), and 4 months (column 3) after transplantation of a TE-stroma in the superior nasal quadrant. Regular diffuse illumination (first row) does not allow visualization of the clear graft, which can only be seen with direct tangential illumination (second row; arrow; same graft and 10 magnification for all six photographs). (BeG) Clinical evolution of grafted and control eyes. (B) Transparency score. (C) Anterior chamber cell score. (D) Anterior chamber flare score. (E) Intraocular pressure (IOP). (F) Graft thickness. (G) Esthesiometry. The parameters described in this figure were measured in all eyes at all indicated time points, with no missing data. The most representative photos are shown. SN graft, superior nasal graft; IT graft, inferior temporal graft. Reproduced from Boulze Pankert M, Goyer B, Zaguia F, Bareille M, Perron MC, Liu X, et al. Biocompatibility and functionality of a tissue-engineered living corneal stroma transplanted in the feline eye. Invest Ophthalmol Vis Sci 2014;55(10):6908e20.

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Self-assembled stromal constructs therefore show promise for regenerative medicine applications; however, they have some significant limitations. Allogeneic cells may be problematic in clinical applications, so patientderived autografts are more likely to be preferred. This poses significant surgical limitations, because patients will need to undergo surgery to isolate the donor tissue and then wait a significant period for the fabrication of a construct. This type of custom construct will then be limited to clinics with on-site cleanroom facilities. The possibility remains of decellularizing the stromal ECM and implanting it, but this has been attempted only in cartilage [55].

CELL-FREE BIOMATERIALS A range of ECM-based or inspired proregeneration scaffolds have been proposed as alternatives to the largely collagenous corneal stroma. These range from human and xenogeneic decellularized ECM (from pig corneas to fish scales) to matrices made from extracted human and animal collagen, RHC, or short peptide analogs of ECM proteins.

Decellularized Extracellular Matrix as Implants The aim of decellularization is to remove the cellular content while preserving the native ECM that is believed to promote regeneration when it is used as cell-free or preecell seeded implants. A wide range of decellularizing methods have been tested, including agents such as enzymes, chelating agents, chemicals (acid and alkali treatment and alcohols), detergents, hypertonic and hypotonic solutions, and physical methods (freezing, pressure, sonication, and mechanical agitation) [56]. Automated devices for decellularization have also been developed [57]. In 2011, Daoud et al. reported the clinical evaluation results of 150 patients who received anterior lamellar keratoplasty with human decellularized corneas sterilized with g-radiation [58]. These implants showed high levels of epithelialization within a few days and no postoperative infection or rejection in all but four cases, all of which occurred in patients with preexisting corneal melting. In 2015, Zhang et al. described the implantation of decellularized porcine corneal stromas into the corneas of 47 patients who had corneal fungal ulcers [59]. They reported that 41 of the implants became transparent over time and 34 patients (72%) showed vision improvement. These results are promising, but caution must be exercised for xenogeneic transplantation because severe allergic reactions have been reported [60,61]. The use of poor-quality cadaveric human corneas as decellularized implants instead of living allografts is still caused by cornea shortage problems and requires confirmation of safety, because immunogenicity and risk of disease transmission remain considerations. The main structural protein in decellularized ECM is collagen. The scales of fish comprise 41e81% of connective tissue protein and collagen and have been examined as collagen implants. Emergency patients with perforated corneal ulcerations or lacerations referred for immediate treatment at European hospitals often receive a temporary human cadaveric donor cornea. This so-called “0-cornea” is used to close the eye only temporarily because it does not meet all requirements of human donor corneas used for PK. However, an “0-cornea” is not always immediately available; hence, the ologen Biocornea was developed (Fig. 63.4) This is a temporary cornea made from decellularized and decalcified fish scales, designed to seal perforated corneas temporarily for up to 72 h while waiting for a donor human cornea to become available. When tested in rat corneas [62], these fish

FIGURE 63.4 Ologen Biocornea derived from fish scale has a morphology similar to human cornea extracellular matrix and is transparent. Courtesy of Aeon Astron Europe B.V.

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scaleederived collagen matrices were not reepithelialized during the 3-week study period and integration was incomplete in two of six animals, leading to the loss of implants. Some melting was also observed. As implants within stromal pockets, the grafts resulted in haziness. The ologen Biocornea was also reported to be able to seal full-thickness cornea perforations in minipig models in which there was no leakage of aqueous humor, the anterior chamber of the eye retained its normal depth, and there was only mild swelling of the wounded cornea [63]. Histopathological examination of the patched cornea showed that the temporary repair did not trigger immune cell invasion and prevented epithelial ingrowths into the wound. An important consideration for the use of xenoderived, decellularized cornea stromas is the heterogeneity of the decellularization process and the risk for allergic reactions to xenogeneic biomaterials (compared with allogeneic biomaterials) and pathogen transmission if tight processing controls are not adhered to [64]. China Regenerative Medicine International is moving ahead with the industrial-scale use of decellularized pig corneas as alternatives to human donor corneas [65].

Collagen-Based Implants The 4-year follow-up of the first successful regeneration of multiple cell types in the human cornea showed that implants made from carbodiimide-cross-linked RHCIII were able to engraft stably without immune suppression [66]. Interestingly, even at 4 years after the operation, the regeneration process in several patients appeared to be ongoing. Histopathology of a graft from a patient who elected to be regrafted (owing to a problem with contact lenses needed to correct astigmatism resulting from suturing) showed that the regenerated neocornea had a structure similar to that of a normal, healthy cornea. For corneas with inflammation or more severe pathologies such as chemical burn, however, implants made from RHCIII alone were not able to prevent neovascularization, as shown in rabbit models of alkali burn [67]. The incorporation of a network of 2-methacryloyloxyethyl phosphorylcholine (MPC), a synthetic lipid with known inflammation suppressing properties [68] properties, blocked neovascularization [69]. When the RHCIIeMPC implants were grafted into three patients as tectonic patches for relief of symptoms from chronic corneal ulceration from burns and a previously rejected graft, i.e., individuals with severe pathologies causing pain, severe discomfort, and photophobia who were diagnosed to be at high risk for rejecting conventional donor corneas, RHCIIIeMPC implants restored surface integrity and relieved the patients of the symptoms [69]. Two of the three patients also showed improvement in vision. A more recent clinical study of RHCIII-MPC implants as anterior lamellar grafts in six patients over an average of 24 months showed stable restoration of surface integrity, relief of symptoms in patients with ulceration, and improvement in touch sensitivity indicating regeneration of corneal nerves [70]. In Islam et al. patterning RHCIIIeMPC hydrogel constructs with microcontact printing, another form of soft lithography, showed that different widths of fibronectin stripes were able to modulate cell adhesion and proliferation characteristics [71].

Peptide Analogs of Extracellular Matrix Full-length ECM macromolecules are often hard to extract in large quantities and to purify. Animal source material has the risk of pathogen transmission and requires screening before use in humans [72]. In addition, these macromolecules are relatively difficult to functionalize compared with synthetic polymers. Hence, the development and use of short analogs of ECM components would allow for more effective processing and applications. Peptide analogs have been tested for the fabrication of scaffolds. Peptide amphiphiles (PAs) containing the ArgeGlyeAsp (RGD) cell adhesion motif from fibronectin were developed by Miotto et al. to promote corneal regeneration (PAs) [73]. When assembled into film coatings, such PAs enhanced the adhesion, proliferation, and alignment of human corneal stromal fibroblasts while inducing the formation of three-dimensional (3D) lamellar-like stromal tissue [74]. Uznalli et al. tested PAs based on laminin peptides [75]. When used as injectable scaffolds in rabbit models of corneal wound healing, they reported an increase in keratocyte migration into the surgically induced wounds that led to enhanced stromal regeneration. In addition, PAs using matrix metalloproteinase sequences were developed that self-detach after reaching confluency [76,77]. Finally, a PA based on lumican that forms nanotape structures was reported to support human corneal fibroblasts and increase collagen production [78]. Islam et al. reported the use of a short peptide analog of collagen in fabricating corneal implants [79]. A sequence from O’Leary et al. [80] conjugated to multiarm PEG through thiol-maleimide chemistry through a “C-G” peptide linker resulted in a hybrid hydrogel consisting of the collagen-like peptide (CLP) and PEG. CLPePEG hydrogels

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were successfully tested as a proregeneration scaffold in corneas of minipigs as anterior lamellar grafts. Regeneration of corneal epithelium, stroma, and nerves were recorded over the 12-month observation period (Fig. 63.5). The timing and similarity of the regeneration were comparable to those of control implants made from RHCIIIeMPC previously tested in human patients [69,70].

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FIGURE 63.5 Postsurgical corneal regeneration at 12 months after collagen-like peptide (CLP)epolyethylene glycol (PEG) implantation. (A) Optically clear CLPePEG implant (arrow) stably integrated within the pig cornea, compared with its healthy, unoperated contralateral cornea (B). Blood vessels are seen in the implanted cornea but stop at the margin of the implant. (C) Hematoxylin-eosin staining of a representative regenerated CLPePEG neocornea compared with a healthy control cornea (D) showing similar morphology. Scale bars ¼ 50 mm. (E) In vivo confocal microscopy shows the regenerated nerve (arrows) in CLPePEG cornea that follow a parallel pattern similar to that of the unoperated cornea (F). (G) Sodium dodecyl sulfateepolyacrylamide gel electrophoresis separated proteins from the central cornea area of both implanted and control corneas show the presence of a1 and a2 chains for type I collagen (a1[I] and a2[I], respectively) and a1 chain for type V (a1[V]), in the regenerated CLPePEG implanted cornea, similar to recombinant human collagen (RHC)-2-methacryloyloxyethyl phosphorylcholine (MPC) and the unoperated healthy control corneas. (H) Analysis of normalized, relative protein content shows that the levels of a1(I) and a2(I) in CLPePEG implanted corneas were similar to those of the control healthy corneas (no statistical significance) whereas the level of a1(V) was higher than that of the control (*P < 0.05). In RHC-MPCeimplanted corneas, levels of both a1(I) and a2(I) were significantly lower than in the unoperated controls (*) whereas a1(V) were similar. No differences were observed between RHCIII-MPC and CLPePEG implants in general. (I) Fourier transform infrared analysis of regenerated pig neocorneas 12 months after grafting with CLPePEG implants shows the presence of both CLP and PEG. e, epithelium; i, implant; s, stroma. Reproduced from Islam MM, Ravichandran R, Olsen D, Ljunggren MK, Fagerholm P, Lee CJ, et al. Self-assembled collagen-like-peptide implants as alternatives to human donor corneal transplantation. RSC Adv 2016;6(61):55745e9.

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CELLeBIOMATERIAL COMPOSITES Cellebiomaterial composites have also been proposed as artificial corneas. Collagen has been used to develop composites of epithelium and stroma from isolated stem cells, e.g., as Real Architecture for 3D Tissue equivalents. A first-in-human study is under way. Silk, particularly the structural protein fibroin derived from the domesticated silkworm Bombyx mori, has been explored as potential corneal stromal scaffolds by several groups in conjunction with therapeutic stem cells. Several examples are discussed here. Groove-patterned, RGD-functionalized silk substrates have been fabricated and seeded with human corneal stromal stem cells (hCSSCs) and human corneal fibroblasts (hCFs) as potential human corneal stromas [81]. The hCSSCs effectively differentiated into stromal cells that elaborated a cornea-specific ECM to mimic human corneal stromal tissue. These constructs were 90e100 mm thick and contained ECM components such as collagen, keratin sulfate, lumican, and keratocan. On the other hand, the hCFs differentiated into myofibroblasts that deposited less-organized collagen in a pattern resembling that of corneal scar tissue. In 2016, Wang et al. reported the successful implantation of helicoidal multilamellar RDG functionalized silk fibroin films in rabbits [82]. At 180 days after the operation, the implants were transparent without signs of neovascularization or immunogenic response. Wang et al. reported the construction of a 3D corneal model using silk with endothelium, stroma, and innervation supported on a free-floating PDMS scaffold in culture media [83], which allowed an aireliquid interface. This and previous models using collagen showed that an aireliquid interface enable optimal corneal cell and nerve ingrowth in multilayered corneal constructs [84e86]. Silk fibroin has also been blended with chitosan to form scaffolds [87]. These were seeded with cornea stromal fibroblasts to form constructs that were implanted into the corneas of 15 rabbits for 12 weeks to allow regeneration. The regenerated rabbit corneas were described as being comparable to healthy, unoperated corneas, with expression of K3/12 expression in the corneal epithelial cells and vimentin in the stromal cells. Hazra et al. reported the creation of a silk fibroin film generated from nonemulberry feeding silkworms, Antheraea mylitta (Am) [88]. The Am film supported the proliferation and differentiation of limbal stem cells. After implantation in rabbits, the films were clear and well-tolerated without neovascularization at 2-month follow-up.

COMPOSITE IMPLANTS INCORPORATING SPECIFIC BIOACTIVE FUNCTIONS Infections are an important cause of vision loss and corneal blindness [2]. Hence, implants that contain nanoparticles designed to combat microorganisms or viruses have been under development. Collagen-based implants incorporating antibiotics (vancomycin) or nanoparticles releasing the antiviral drug acyclovir have been tested and found to be effective at blocking the activity of Staphylococcus aureus in rabbit infection models and herpes simplex virus serotype 1 (HSV-1), respectively, in vitro [89,90]. A variant on implants with antiviral properties is a collagen-based implant that incorporates corneal cells engineered by gene transfer to secrete the innate antiviral cathelicidin peptide, LL-37 [91]. The secreted LL-37 was able to inhibit viral binding and reduce the incidence of plaque formation and plaque size in vitro but was insufficient to protect cells completely from HSV-1 infection. Silver nanoparticles (AgNPs) have been reported to have broad antiinfective and antiinflammatory properties. Collagen implants incorporating AgNPs have also been developed, along with a strategy to prestabilize and incorporate AgNPs with different colors into collagen matrices to fabricate corneal implants of different colors and antimicrobial properties [92].

CHALLENGES From a regulatory perspective, regenerative implants and grafts in the cornea are primarily classified as medical devices or ATMPs, respectively. The most straightforward type of implant to manufacture is a fully synthetic scaffold, especially one that is thermostable and can be sterilized by g-irradiation or electron beam. These scaffolds do not have to contend with rules concerning xenografts. Scaffold derived from animals, either through decellularization or by isolating structural proteins, requires more testing to comply with requirements to prevent pathogen transmission. The most complex regulatory submissions are cell and scaffold composites, which are regulated under the ATMP guidelines. Pellegrini et al. successfully initiated the first application for an ATMP

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product in Europe, but this remains an area in which there are many safety considerations and few examples to guide both the applicant and the regulatory agencies [93]. It will likely remain incumbent on the applicants to demonstrate that their product is effective, reproducible, and sterile. Extensive risk analysis of all phases of the production and implantation of an ATMP will be required to generate sufficient confidence that the product will be safe for an extended market.

CONCLUSIONS AND FUTURE PERSPECTIVE There is a wide variety of solutions for regenerating the cornea. Most recently, 3D bioprinting was used to print corneal stromas complete with cells [94]. Progress has been made in cell-based therapies, ECM-based scaffolds, nanomaterials, and combinations of advancements in cell and biomaterials. The challenge to replace human donor corneas and KPros as the current standard of care remains complex. Advances in cell-based therapies have significantly improved the probable efficacy of corneal autografts, and biomaterials have improved the structural properties of ECM-based scaffolds; however, no single therapy has emerged as a stable, affordable solution for the many causes of corneal damage. It is possible that in the future, biomaterial scaffolds may be combined with surgical techniques such as SLET or cultured cell solutions to customize commercially available scaffolds for individual patients. Nanomedicine also opens the door to the possibility of incorporating therapies for eye infections directly into surgical solutions, which may speed postoperative recovery and prevent latent infection recurrence in patients with corneal blindness caused by infections such as HSV.

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Self-assembled collagen-like-peptide implants as alternatives to human donor corneal transplantation. RSC Adv 2016;6(61):55745e9. [80] O’Leary LER, Fallas JA, Bakota EL, Kang MK, Hartgerink JD. Multi-hierarchical self-assembly of a collagen mimetic peptide from triple helix to nanofibre and hydrogel. Nat Chem 2011;3(10):821e8. [81] Wu J, Rnjak-Kovacina J, Du Y, Funderburgh ML, Kaplan DL, Funderburgh JL. Corneal stromal bioequivalents secreted on patterned silk substrates. Biomaterials 2014;35(12):3744e55. [82] Wang L, Ma R, Du G, Guo H, Huang Y. Biocompatibility of helicoidal multilamellar arginine-glycine-aspartic acid-functionalized silk biomaterials in a rabbit corneal model. J Biomed Mater Res B Appl Biomater 2015;103(1):204e11. [83] Wang S, Ghezzi CE, Gomes R, Pollard RE, Funderburgh JL, Kaplan DL. In vitro 3D corneal tissue model with epithelium, stroma, and innervation. Biomaterials 2017;112:1e9. [84] Suuronen EJ, McLaughlin CR, Stys PK, Nakamura M, Munger R, Griffith M. Functional innervation in tissue engineered models for in vitro study and testing purposes. Toxicol Sci 2004;82(2):525e33. [85] Suuronen EJ, Nakamura M, Watsky MA, Stys PK, Mu¨ller LJ, Munger R, et al. Innervated human corneal equivalents as in vitro models for nerve-target cell interactions. FASEB J 2004;18(1):170e2. [86] Griffith M, Osborne R, Munger R, Xiong X, Doillon CJ, Laycock NL, et al. Functional human corneal equivalents constructed from cell lines. Science 1999;286(5447):2169e72. [87] Guan L, Ge H, Tang X, Su S, Tian P, Xiao N, et al. Use of a silk fibroin-chitosan scaffold to construct a tissue-engineered corneal stroma. Cells Tissues Organs 2013;198(3):190e7.

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[88] Hazra S, Nandi S, Naskar D, Guha R, Chowdhury S, Pradhan N, et al. Non-mulberry silk fibroin biomaterial for corneal regeneration. Sci Rep 2016;6:21840. [89] Bareiss B, Ghorbani M, Li F, Blake JA, Scaiano JC, Zhang J, et al. Controlled release of acyclovir through bioengineered corneal implants with silica nanoparticle carriers. Open Tissue Eng Regen Med J 2013;3:10e7. [90] Riau AK, Mondal D, Aung TT, Murugan E, Chen L, Lwin NC, et al. Collagen-based artificial corneal scaffold with anti-infective capability for prevention of perioperative bacterial infections. ACS Biomater Sci Eng 2015;1(12):1324e34. [91] Lee CJ, Buznyk O, Kuffova L, Rajendran V, Forrester JV, Phopase J, et al. Cathelicidin LL-37 and HSV-1 corneal infection: peptide versus gene therapy. Transl Vis Sci Technol 2014;3(3):4. [92] Alarcon EI, Vulesevic B, Argawal A, Ross A, Bejjani P, Podrebarac J, et al. Coloured cornea replacements with anti-infective properties: expanding the safe use of silver nanoparticles in regenerative medicine. Nanoscale 2016;8(12):6484e9. [93] Pellegrini G, Lambiase A, Macaluso C, Pocobelli A, Deng S, Cavallini GM, et al. From discovery to approval of an advanced therapy medicinal product-containing stem cells, in the EU. Regen Med 2016;11(4):407e20. [94] Isaacson A, Swioklo S, Connon CJ. 3D bioprinting of a corneal stroma equivalent. Exp Eye Res. 2018;173:188e93.

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64 Alimentary Tract Richard M. Day University College London, London, United Kingdom

INTRODUCTION The alimentary tract is a hollow organ that starts at the mouth and terminates at the anus. It conducts a number of highly complex and diverse functions that are regulated by distinct cellular and functional differences along the tract length, which allow it to provide the body with nutrients, water, and electrolytes. To achieve this, it is essential that the lumenal contents be propelled along at a rate that will allow efficient digestion and absorption to take place while also enabling waste products to be stored and excreted in a controlled manner. In addition to this, an important symbiotic relationship exists between bacterial species that colonize the alimentary tract of the host [1]. Therefore, the surface of the alimentary tract must also provide a barrier against unwanted entry of organisms and toxins. If the barrier function is breached, specialized cells and tissues within the gut wall provide an important component of the immune system to protect the host. Dysfunction of the alimentary tract may result from a variety of congenital and acquired conditions that can affect any of its physiological functions. This chapter will discuss knowledge regarding tissue engineering different regions of the alimentary tract, highlighting successful strategies as well as failures and some obstacles that have yet to be overcome in this rapidly evolving field.

ESOPHAGUS The esophagus is a muscular tube measuring approximately 25 cm in length in adult humans. It functions primarily as a conduit to connect the pharynx with the stomach, providing coordinated peristaltic contractions in response to swallowing to propel food into the stomach. The esophageal mucosa is lined by stratified, squamous, nonkeratinized epithelium. The submucosa contains muscle, nerve, blood vessels, lymphatics, and mucosal glands. The muscularis has two layers consisting of an outer longitudinal layer and an inner circular layer. Both layers consist of striated muscle in the upper portion and smooth muscle in the lower third, continuous with the muscle layers of the stomach. The myenteric plexus exists between the muscle layers. The esophagus has no serosa and its vascular supply is less extensive compared with the intraabdominal portions of the gut. Sphincters at the upper and lower ends of the esophagus ensure food is transferred appropriately between it and the pharynx or stomach. The upper esophageal sphincter, found in the upper 3e4 cm of the esophagus, and the lower esophageal sphincter, located 2e5 cm above the gastroesophageal junction, remain tonically and strongly constricted to prevent air from entering the esophagus during respiration between swallowing and reflux of stomach contents into the esophagus between peristaltic waves, respectively. Regenerative medicine techniques are being explored for a number of conditions affecting the esophagus. Gastroesophageal reflux disease is one of the most common disorders affecting the gastrointestinal tract, resulting from lower esophageal sphincter incompetence. Medical therapy is generally safe and effective in most cases, but for patients for whom this option fails, antireflux surgery or endoscopic procedures that involve injecting bulking materials may be used in an attempt to narrow the lumen of the lower esophagus. Attempts have been made to restore physiological function of the esophagus using regenerative medicine. The feasibility of using a suspension of muscle precursor cells to restore gastroesophageal function in a model of gastroesophageal reflux disease has been explored Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00064-3

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[2]. Muscle precursor cells isolated from expanded satellite cells derived from skeletal muscle fibres were injected into the gastroesophageal junction after cryoinjury. Histology showed an increase in myofibers at the site of injection that had fused into newly formed or preexisting myofibers. The need remains to demonstrate that cells injected in this manner can contribute to functional improvement of damaged esophageal sphincter, but the feasibility of using this approach offers a promising therapy for this common condition. Esophageal reconstruction is a requirement for congenital esophageal atresia, burns, malignancy, or severe benign disease. Surgical techniques include stretching, circular myotomy, and interposition of stomach or colon, but these approaches are frequently associated with complications including stricture, leakage, elongation, and gastroesophageal reflux. An artificial esophageal construct has been sought for many years. To be effective, the construct must be implantable without rejection and biocompatible to support appropriate tissue growth, and retain biomechanical characteristics of native esophageal tissue, i.e., be soft and elastomeric while maintaining a tubular structure when implanted in vivo. Attempts to tissue engineer replacement esophageal tissue have included both patch and circumferential implantation of constructs composed of synthetic as well as natural scaffold materials. A full-length of tissue engineered esophagus has not been produced, but a number of incremental advances toward this goal have been achieved. Early attempts exploring the use of a nondegradable prosthetic tubes in canine models were of limited success [3]. Surgical reconstruction techniques of the esophagus after resection for structures and malignancies have involved the transfer of small segments or patches of skin and other tissues on a vascular pedicle, with moderate results [4e6]. Tissue engineered sheets of autologous oral mucosal epithelial cells were successfully transplanted by endoscopy in a canine model [7]. The transplanted sheets adhered to the underlying esophageal muscle layers created by endoscopic submucosal dissection and enhanced wound healing without postoperative stenosis. Because the interaction between the epithelium and mesenchymal cells is thought to reduce fibrosis and scarring that cause stenosis, the investigators suggested that this approach may offer a novel therapy to reduce scarring and prevent painful constriction that can be associated with endoscopic submucosal dissection for the removal of large esophageal cancers. A variety of scaffold materials to support cell and tissue esophageal constructs have been investigated [8e12]. Acellular scaffolds composed of extracellular matrix components have been explored primarily because they are assumed to be advantageous over synthetic scaffold materials owing to their ability to promote cell attachment, growth, and cellecell signaling among different tissue components. Decellularized esophageal tissue can be produced via repeated detergent-enzymatic treatment that results in a scaffold with biocompatibility suitable for the growth of esophageal epithelial cells [13,14]. Based on these preclinical findings, it has been envisaged that human donor esophageal tissue might one day be used in a manner similar to that described for the tissue engineering of human airway tissue [15]. Scaffolds derived from small intestinal submucosa (SIS) have been widely investigated for tissue engineering replacement esophageal constructs. SIS consists of extracellular matrix material harvested from porcine small intestine and has been used extensively in tissue engineering experiments since it was originally described by Matsumoto and colleagues in 1966 for use in large vein replacement in dogs [16e20]. It has been successfully applied to regenerative medicine applications in humans, including repair of hernias, diaphragms, and tympanic membranes, and for large wound coverage [21e24]. The success with using SIS as a scaffold to promote tissue regeneration appears to relate to the retention of collagen (types I, II, and V), growth factors (transforming growth factor, fibroblast growth factor 2, and vascular endothelial growth factor), glycosaminoglycans (hyaluronic acid, chondroitin sulfate, and heparin sulfate), proteoglycans, and glycoproteins (fibronectin) during the fabrication process [25,26]. It is thought that the resulting scaffold has a composition closely resembling native extracellular matrix, which makes it ideally suited for the attachment and growth of new tissue. The extent of circumferential replacement of esophageal tissue appears to have an impact on the outcome of attempts to tissue engineer esophagus, with patches producing better results compared with tubular segments. Lopes and colleagues successfully used SIS patches to repair defects to the anterior wall of cervical or abdominal esophagus in rats without signs of stenosis over 150 days [27]. Likewise, Badylak and colleagues used SIS patches (or urinary bladder submucosa) to repair esophageal defects created in dogs without clinical signs of esophageal dysfunction [28]. However, the latter study reported signs of stenosis in dogs receiving complete circumferential segmental grafts of SIS [10]. Doede and colleagues reported similar findings, with severe stenosis occurring when relatively short tubular lengths (4 cm) of SIS were used in alloplastic esophageal replacement in piglets [28]. Thus, although scaffolds consisting of only extracellular matrix have shown the capacity to promote cell growth in vitro and tissue regeneration of patch defects in vivo, replacement of circumferential defects without stricture

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formation remains difficult to achieve. The outcome of repairing full circumferential defects using SIS may be improved by optimizing the poor mechanical properties of the scaffold material. To address this, hybrid scaffolds composed of SIS combined with synthetic polyesters (poly[3-hydroxybutyrate-co-3-hydroxyhexanoate] and poly [lactide-co-glycolide]) were assessed for their feasibility as a tissue engineering scaffold for esophageal constructs [29]. Improved biocompatibility was observed with the hybrid scaffold compared with scaffolds produced from the synthetic material alone. Epithelial-mesenchymal cell signaling is likely to have an important role in facilitating reconstruction of the esophageal construct after implantation. A similar effect has been shown in bladder reconstruction, in which the presence of urothelium led to infiltration of fibroblasts into acellular matrices and apparent transdifferentiation into a smooth muscle phenotype [30]. Signaling from the mesenchymal cell population appears to be equally important in promoting growth of overlying epithelium [31]. Complete reepithelialization with little inflammatory response and evidence of skeletal muscle regeneration was observed when bone marrow mesenchymal stem cells were seeded onto the SIS scaffold implanted in a canine model [32]. Moreover, the presence of epithelialmesenchymal signaling may prevent stricture formation in an esophageal construct, a problem frequently encountered with many of the scaffolds tested [10,33]. Similar signaling properties have been demonstrated in bladder reconstruction in which acellular collagen scaffolds seeded with urothelium and smooth muscle cells prevented tissue contraction [34]. Likewise, the interaction of muscle with the ablumenal surface of esophageal scaffolds at the time of implantation of partially circumferential grafts appears to have accounted for the reduced stricture formation observed in a canine model of esophageal reconstruction described by Badylak and colleagues [11]. It can be concluded from these observations that careful consideration of the order in which cells are added to the tissue engineered construct will improve the likelihood of achieving a successful outcome. In addition to SIS, gastric acellular matrix was used as scaffold by Urita and colleagues to regenerate esophagus in a rat model [35]. Grafts of gastric acellular matrix were used to patch defects in the abdominal esophagus and animals were killed at points between 1 week and 18 months. Although regeneration of the muscle layer or lamina muscularis did not occur, there was no evidence of stenosis or dilatation at the graft site. The matrix obtained in this study was from whole stomachs, but the investigators suggested that gastric acellular matrix may provide an autologous source of naturally derived extracellular matrix scaffold in a clinical setting, because the portion of stomach destroyed to obtain the matrix is minimal. It remains to be seen whether this approach is feasible in a larger animal model, but the use of autologous acellular matrix scaffolds avoids concerns related to the use of xenogenic scaffold materials such as porcine-derived SIS. In addition to the risk of transmitting viral pathogens and prions, cultural and religious beliefs may need to be considered when using acellular matrix scaffolds derived from certain species. In addition to SIS, other biological materials have been investigated for esophageal tissue engineering. Extracellular matrix scaffold has been generated from ovine forestomach tissue [36]. Moving away from mammalian sources of scaffold material, Franck and colleagues reported that a bilayer silk fibroin matrix composed of porous silk fibroin foam annealed to a homogeneous silk fibroin film exhibited improved cell attachment and spontaneous differentiation of esophageal epithelial cells toward a suprabasal cell lineage compared with SIS scaffolds [37]. Although the esophagus can be considered one of the less complex regions in the alimentary tract, several significant hurdles still need to be overcome before tissue engineering and clinical replacement of full-length esophageal segments become a clinical reality in humans. Unlike patch grafts, replacement of longer lengths of tissue will be unable to rely on adjacent esophagus to cover the surface area of larger scaffolds via guided tissue regeneration. Improved methods from isolating and expanding the different esophageal cell populations will therefore be a prerequisite for successful tissue engineering of larger constructs. Kofler and colleagues identified subsets of ovine esophageal epithelial cells that may help achieve this [38]. PCK-26epositive esophagus epithelial cells demonstrated high proliferative capacity and uniform coverage on collagen scaffolds, which the investigators suggested could have an important role for successfully tissue engineering esophagus. Further refinement of the scaffold material may also improve epithelialization. For example, the inclusion of copper into acellular porcine SIS scaffolds was reported to enhance epithelialization of the scaffold in a canine model of esophageal repair [39]. Failure to regenerate a functioning muscle layer may not be problematic for short or noncircumferential grafts, but for longer lengths of esophagus the presence of an innervated functional muscle layer will be essential. A retrospective study investigating the temporal appearance and spatial distribution of nervous tissue in a canine model of esophageal reconstruction using porcine urinary bladder submucosa showed the presence of nerve tissue within sites of the remodeling scaffold [40]. Although the study was unable to demonstrate whether the nervous tissue was functional or to distinguish among the various subsets of neurons, it opens the possibility of using similar models to identify mechanisms that promote innervation that will facilitate the tissue engineering of functional tissue. Peristalsis of food also depends on the correct orientation of muscle fibers in the wall of the alimentary tract.

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To address this, promising results have been obtained with orienting smooth muscle tissue on unidirectional scaffolds for tissue engineered esophagus in rats [41]. Oriented stands of smooth muscle mimicking the configurations found in the native organ were engineered when cells were seeded onto unidirectional scaffolds. These were assembled with esophageal epithelium to create a hybrid approach.

STOMACH Gastric disease affects approximately 10% of the world’s population and includes gastritis, peptic ulcers, and gastric cancer. Insufficient stomach mass, which may arise from gastrectomy or congenital microgastria, is associated with increased patient morbidity. The stomach functions as a digestive organ and reservoir and is anatomically divided into four regions (cardiac, fundus, corpus, and pylorus). The gastric glands are tubular structures whose cellular composition and function are specialized according to each region of the stomach. Such specialization, combined with the harsh environment created by the lumenal contents, makes tissue engineering of the stomach as whole challenging. The size and shape of the stomach vary, depending on its contents. The stomach wall contains outer longitudinal and inner circular layers of smooth muscle, with an innermost layer of oblique muscle fibers. These layers facilitate important functions including storage of ingested food in the stomach until it can be accommodated in the lower portion of the alimentary tract, mixing of the food to form chyme, and regulation of food transit into the small intestine at an optimal rate for digestion and absorption. Stomach emptying is controlled by the gastric food volume and the release of the hormone gastrin, as well as feedback signals from the duodenum. Tissue engineered neostomach constructs to patch partial gastrectomy have been explored in a canine model using a two-part sheet composed of an outer layer of collagen sponge and a temporary inner silicone sheet to protect the collagen from degradation by the acidic stomach juices and provide mechanical support [42]. After removal of the silicone sheet at 4 weeks, evidence of stomach regeneration was observed and complete coverage of the scaffold had occurred by 16 weeks, confirmed by the presence of mucosa and a thin muscular layer. Acid production capacity was present in the regenerated stomach wall but the contractile response to acetylcholine was poor [43]. Technical difficulties associated with suturing and endoscopic removal of the silicone sheet in this model were addressed by creating a tissue engineered sheet without silicone that had sufficient strength to allow suturing and resist anastomotic dehiscence [44]. The silicone sheet was replaced by a biodegradable copolymer of poly(D,L-lactide) and ε-caprolactone (PDLCL) on the mucosal side of the collagen scaffold, both of which were completely absorbed at 16 weeks’ implantation. Although regeneration of the stomach mucosa was observed, the replacement of the silicone sheet with PDLCL did not provide sufficient mechanical strength to prevent significant shrinkage of the scaffold. The feasibility of creating new stomach tissue using stomach-derived organoid units harvested from neonatal and adult rats has been investigated [45]. The organoid units were seeded onto polymer scaffold tubes to form constructs that were implanted into the omentum of adult syngeneic rats. At 4 weeks, the construct was anastomosed to the small intestine. Histology of the tissue engineered stomach tissue was similar to native stomach, with gastric pits, squamous epithelium, and positive staining for a-actin smooth muscle in the muscularis and gastrin indicating the presence of a well-developed gastric epithelium. The same approach has been used to tissue engineer stomach neoconstructs in an autologous large animal model [46]. A limitation of approaches using autologous or allogenic organoid units is the tissue source. To address this, three-dimensional gastric organoid tissue from human pluripotent stem cells (PSCs) have been generated by temporal manipulation of the fibroblast growth, WNT, bone morphogenetic protein, retinoic acid and epidermal growth factor (EGF) signaling pathways [47]. The primitive gastric organoids exhibited molecular and morphogenetic developmental stages similar to those observed in the developing antrum of murine stomach. The proliferative zones contained LGR5þ cells, mucus-secreting antral and pylorus cells, and gastric endocrine cells; however, the acid secreting (corpus) region was not developed in this model. A critical factor for the development of all the stomach regions from stem cells appears to be the Barx-1 gene [48]. Noguchi and colleagues demonstrated Barx-1einducing culture conditions generated spheroids of fully functional stomach-like tissue in vitro from mouse embryonic stem cells. The spheroids were able to develop into functional corpus and antrum tissue that secreted pepsinogen and acid. The relevance of mesenchymal stem cells to optimize and condition the cellular milieu within the tissue engineered construct has also been shown with stomach tissue engineering. Tissue regeneration after the creation of a full-thickness stomach defect in rats was enhanced when SIS scaffolds were used in conjunction with mesenchymal

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stem cells [49]. Whereas contractility in response to a muscarinic receptor agonist, a nitric oxide precursor, or electrical field stimulation was observed in all groups, smooth muscle layers were both longer and better structured compared with SIS grafts not seeded with mesenchymal stem cells.

SMALL INTESTINE The small intestine in adults measures approximately 6 m in length from the duodenojejunal flexure to the ileocecal valve. Its primary function is absorption of nutrients from the lumen, a process facilitated by specialized mucosal surface features (folds of Kerckring, villi, and microvilli) that increase the absorptive surface area about 600-fold to approximately 250 m2. The mucosa is lined with epithelium overlying lamina propria containing vascular and reticular stroma, large aggregates of lymphoid tissue called Peyer patches, and a strip of smooth muscle called the muscularis mucosa. Intestinal stem cells reside at the base of epithelial invaginations into the mucosa called crypts and develop into all four lineages of epithelial cells that line the intestine [50]. Epithelial cells migrate along the cryptevillus axis, differentiating and maturing toward the lumen of the bowel where they become senescent over the course of a few days and are shed into the lumen of the bowel. A lack of intestinal epithelial stem cell markers has hampered identification and isolation of pure populations of cells for regenerative medicine purposes. Studies have shown that Musashi-1 may be a marker of intestinal stem cells [51,52], and a Sox9-(enhanced green fluorescent protein [EGFP]) mouse model has been used to enrich multipotent intestinal epithelial stem cells [53]. Using a culture system that mimics the native intestinal epithelial stem cell niche, these cells are capable of generating “organoids” that contain all four epithelial cell types of the small intestinal epithelium. Furthermore, Sox9-EGFP multipotent intestinal epithelial stem cells express CD24, which may facilitate their enrichment by fluorescence activated cell sorting using widely available antibodies. The submucosa consists of fibrous connective tissue that supplies blood and lymphatic vessels to the mucosa. The muscularis propria consists of an inner layer of circular muscle and an outer longitudinal muscle layer. The muscularis propria is covered by the adventitia, a layer of loose connective tissue, and the serosa, a mesothelial lining of peritoneum. The function of the small intestine cannot be replaced by transposing another part of the gut. Intestinal ischemia and bowel resection for tumors and inflammatory bowel disease can result in short bowel syndrome when more than 75% of the small intestine is lost. Short bowel syndrome is often associated with intestinal failure and the requirement of lifelong nutritional support (total parenteral nutrition), which is frequently accompanied by severe complications such as liver failure, line sepsis, and poor long-term survival rates. The length of residual intestine is critical for these patients; thus, techniques for increasing absorptive surface area have been sought for many years. Surgical options for increasing the absorptive surface or slowing the transit time to enhance absorption have been reported, but these approaches require longer residual intestinal segments and most have only limited long-term clinical success [54e57]. Small bowel transplantation is a viable option for some patients, but this procedure has limitations including the availability of donor tissue, the need for long-term immunosuppression, graft versus host disease, and potential posttransplant lymphoproliferative disorder [58]. The amount of small bowel required for successful nutritional rehabilitation depends on factors including the patient’s age, the amount of small bowel present, the presence or absence of the ileocecal valve, and the amount of large bowel present. Therefore, small bowel elongation of just a few centimeters could allow many patients to become independent of total parenteral nutrition. Distraction enterogenesis has been devised as a novel method to increase intestinal length by applying linearly directed force, resulting in increased surface area and epithelial cell proliferation [59,60]. Devices used for distraction enterogenesis include extralumenal, radially selfexpanding shape memory polymer cylinders [61]; biodegradable springs composed of polycaprolactone created to lengthen intestinal segments mechanically while avoiding the need for subsequent retrieval [62]; and doubleballoon catheter devices [63]. Biological mechanisms that account for distraction enterogenesis are unknown [64], but the effect can be enhanced by adding exogenous glucagon-like peptide 2 [65] or by codelivering microspheres that provide sustained release of basic fibroblast growth factor (bFGF), which results in improved vascularity [66]. The use of growth factoreembedded scaffold materials is an effective method for improving the short half-lives of growth factors. An example used to achieve growth factoreembedded scaffolds for distraction enterogenesis include subcritical CO2 to embed heparin-binding EGF-like growth factor into polyglycolic acid/ poly-L-lactic acid scaffolds [67]. Local delivery of the trophic growth factor improved structure of the tissue engineered intestine.

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Several different approaches, using either guided tissue regeneration or tissue-engineering neotissue constructs, were taken to regenerate small intestine that used combinations of various synthetic and natural scaffold materials, different cell types, and surgical procedures. Early attempts to patch bowel defects using the serosal surface of another piece of intestine resulted in its being covered with regenerated mucosa [68,69]. This paved the way for other researchers to investigate a variety of scaffold materials, such as polytetrafluoroethylene tubing, for the ingrowth of neointestine via guided tissue regeneration [70]. The use of nonresorbable materials for studying intestinal morphogenesis and regeneration continues to be interesting [71], but the use of resorbable scaffold biomaterials for intestinal tissue engineering has become the predominant approach. Chen and Badylak used SIS to patch partial defects in the small bowel wall of a canine model [72]. Histological evaluation showed the presence of mucosa, varying amounts of smooth muscle, sheets of collagen, and an outer serosal layer. However, in the same study, attempts to use a tubular configuration of SIS were unsuccessful. The tubes either leaked or became obstructed; this occurred primarily because the SIS was unable to maintain lumenal patency when exposed to the moist lumenal contents. Similar limitations with the mechanical integrity of the scaffold material were reported by Pahari and colleagues, who used guided tissue regeneration to create a segment of new intestine in rats using acellular dermal matrix (AlloDerm) rolled into tubes [73]. Building on the work, elongation of the intestine was achieved using an acellular biologic scaffold to create autologous bioartificial intestinal segments (BIS) [74]. The BIS were demonstrated to have functional absorptive characteristics [75]. Other approaches using biologically derived scaffolds included the use of allogenic aortic graft segments interposed in an excluded small bowel segment wrapped in omentum, which resulted in intestinal-like wall transformation of the aortic graft [76]. In an attempt to maintain an open lumen in the tissue engineered intestine, Hori and colleagues reported that scaffolds composed of sheets of acellular collagen sponge wrapped on a temporary silicone stent and covered with omentum guided tissue regeneration of almost all layers of the gastrointestinal tract in a canine model, but only a thin muscularis mucosa was present and the muscularis propria was absent [77]. The same group explored the addition of mesenchymal stem cells seeded onto a collagen scaffold, which it was hypothesized might differentiate “site-specifically” into muscle cells and regenerate the muscle layer [78]. Intestinal regeneration occurred but muscle regeneration in an organized manner was not observed. Wang and colleagues used a rat model to evaluate the feasibility of regenerating tubular intestine using sheets of rat-derived SIS wrapped around a silicone stent [79]. The tubular graft was interposed in the middle of a Thiry-Vella loop (a defunctionalized segment of ileum that is brought out as a double ileostomy) in Lewis rats. The silicone stent was left in place for 3 weeks to maintain lumenal patency during tissue regeneration. At 4 weeks, an epithelial layer had begun to form and completely covered the lumenal surface by 12 weeks. The neomucosa had a typical morphology containing goblet cells, Paneth cells, enterocytes, and enteroendocrine cells. Although the regenerated bowel contained bundles of smooth muscle-like cells, especially near the sites of anastomosis, the quantity and organization of the muscle layer differed from those found in native small intestine; they were predominantly circular muscle with no longitudinal muscle. The use of a ThiryVella loop in the model created by Wang may have facilitated mucosal development in the neointestine by protecting it from alimentary transit and creating an isolated environment that avoided the food stream and digestive enzymes. Lee and colleagues observed only minimal intestinal regeneration in a rat model used to evaluate SIS scaffolds. From this, they concluded that SIS scaffolds alone were insufficient to regenerate small intestine and suggested that the use of appropriate progenitor cells is probably necessary to facilitate the regeneration of small intestine [80]. Many of the studies reporting intestinal tissue engineering strategies are based on methodologies used in the pioneering work conducted by Vacanti and colleagues in the 1980s and 1990s that combined intestinal tissue with scaffolds [81]. Important to these studies were previous investigations by Tait and colleagues that showed intestinal tissue could be separated by enzymatic digestion to produce organoid units [82]. These clusters of cells contained all of the elements of the intestinal mucosa including stem cells and mesenchyme, which could be used to regenerate intestinal neomucosa expressing digestive enzyme activities and glucose transport capacity similar to those of agematched native intestinal mucosa. When organoid units were subcutaneously grafted, they displayed different epithelial populations consistent with epithelial transit amplifying and stem cell populations [83]. Subsequent studies demonstrated that transplanting organoid units onto biodegradable polymer scaffolds followed by implantation into the omentum of syngeneic adult animals resulted in the formation of neointestinal cysts attached to a vascular pedicle with mucosa facing a lumen that contained mucoid material [84]. The mucosa of the neointestine created with this technique showed morphological similarities to native intestine, including the formation of a primitive cryptevillus axis lined with columnar epithelial cells and goblet cells and a polarized epithelium with the brush border enzyme sucrase expressed at the apical surface and laminin at the basolateral surface, and

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transepithelial resistance similar to that of native intestine [85]. For all of these promising findings, it is essential that the tissue engineered construct facilitates nutrient absorption. Anastomosis of the tissue engineered cyst-like structures to native jejunum in adult rats provided continuity with the native intestinal tract, which resulted in a more developed neomucosa containing significant increases in villus number, villus height, and surface length of the cyst compared with nonanastomosed cysts [86]. The investigators postulated that anastomosis may have facilitated neomucosal growth in the cysts by draining lumenal contents or via stimulatory factors present in the lumenal contents of the native intestine in continuity with the neomucosa. The anastomosed neointestine was also shown to express of Naþ-dependent glucose transporter SGLT1 [87] and a mucosal immune system with intraepithelial and lamina propria immune cells similar to that of native jejunum [88]. The native small intestine has a great adaptive and compensatory capacity in response to massive small bowel resection, which is considered to be controlled by humoral factors. The mucosa of the neointestine was also shown to possess this adaptive capacity after massive small bowel resection, resulting in a significant regenerative stimulus for the morphogenesis and differentiation of the tissue engineered intestine [89]. An improvement in intestinal function, capable of facilitating patient recovery after massive small bowel resection, was putatively demonstrated when cysts containing neointestine were anastomosed to native small bowel during an 85% enterectomy in rats [90]. The study showed that animals with tissue engineered intestine returned to their preoperative weight more rapidly compared with animals undergoing small bowel resection alone. These findings significant because they are the first to suggest that tissue engineered intestine may provide a therapeutic intervention for managing patients with short bowel syndrome. Although it is tempting to speculate that the observed effects resulted from neointestine restoring absorptive function after small bowel resection, the mechanism underlying the beneficial effects remain uncertain [91]. It has been postulated that the amount of intestine replaced by the anastomosed neointestine (approximately 4 cm) was far shorter than the amount resected, probably approximately 10% of the original length and is unlikely to have added sufficient mucosal surface area to account for the increase in postoperative weight observed. Furthermore, the improved nutrition may have resulted from the tissue engineered intestine slowing intestinal transit, leading to increased absorption and weight gain, a principle that could be achieved with simpler remedial surgical procedures [91]. To gain a better understanding of the mechanisms underlying the formation of neointestine, the model has been transitioned from a rat to a mouse model, which has enabled the use of transgenic tools for lineage tracing. This demonstrated that epithelium, muscularis, nerves, and blood vessels are derived from multicellular organoid units derived from donor small intestines of transgenic mice [92]. Studies also investigated the effects of donor age and region where intestine crypts are harvested [93]. In mice, higher efficiency of enterosphere formation was observed with crypts harvested from tissue collected from the proximal small intestine, and also in young mice. A significant drawback with this approach is the need for large amounts of donor tissue to harvest a sufficient number of organoid units to seed scaffolds that will generate a relatively short length of neointestine, which is likely to offer only limited therapeutic value [91]. A solution might exist with the use of yet unexplored alternative sources of intestinal epithelial stem cells, such as bone marrowederived cells and PSCs circulating in the peripheral blood [94,95]. New methods for generating organoid units that address the limitations of harvesting donor intestinal tissue have also been explored. McCracken and colleagues reported a protocol for differentiating human PSCs into threedimensional (3D) human intestinal tissue, developing into intestinal tissue containing all major types of intestinal epithelial cells and mesenchymal components [96]. To address challenges associated with the limited availability of autologous donor intestine in patients with short bowel syndrome, protocols exist for generating enteroids from minimal quantities of starting material [97]. Also, refinement of the harvested donor stem cells or manipulation of growth factors in the local environment may provide a method for enhancing the quality of the neointestine. The morphology of tissue engineered intestine was improved by seeding scaffolds with intestinal stem celleenriched crypts [98]. Greater circumferential mucosal engraftment and an average villous height closer to native intestine were achieved with the purified crypts collected using a filtration-based system compared with scaffolds seeded with a villous fraction containing differentiated epithelial cells. Likewise, manipulation of the expression of growth factors that control the growth and differentiation of the intestine during development might provide a valuable approach for improving the formation of tissue engineered small intestine. Tissue engineered organoid units overexpressing fibroblast growth factor 10 resulted in larger tissue engineered constructs, with longer villi and a greater proportion of proliferating epithelial cells [99]. Organoid units derived from human postnatal, small bowel resection specimens, were seeded onto biodegradable scaffolds and implanted into nonobese diabetic/severe combined immunodeficiency g chainedeficient mice [100]. After 4 weeks, the tissue engineered small intestine contained the four major types of differentiated intestinal epithelial cells, muscularis, and intestinal subepithelial myofibroblasts.

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Bone marrowederived epithelial cell adhesion moleculeepositive, and CD133-positive cells were used to recellularize human cadaveric small bowel specimens that had been chemically decellularized [101]. After recellularization, the tissue engineered small intestine contained mucin-positive goblet cells, cytokeratin 18epositive epithelial cells in villi, and smooth muscle cells in muscularis mucosa. Cryopreservation of the organoid units could be used to delay production of tissue engineered small intestine, a procedure that could be particularly helpful in patients who are critically ill and require delayed autologous implantation of a tissue engineered construct. Cryopreservation using vitrification led to higher viability of organoid units compared with standard snap-freezing; the thawed organoid units were capable of producing tissue engineered small intestine [102]. Another important aspect of intestinal tissue engineering is the ability for the neointestine to repair, regenerate, and remodel. The latter is particularly important when considering the use of engineered intestinal tissue for children, in whom the length of the intestine increases significantly during development. The trophic effects of glucagon-like peptide-2 (GLP-2) have been evaluated on neointestinal growth [103]. GLP-2 is an endogenous regulatory peptide with potent trophic effects on intestinal mucosal growth and an ability to modulate the expression of Naþ-glucose cotransporter 1 (SGLT1). Adult rats with neointestinal implants that received subcutaneous injections of a GLP-2 analog twice daily for 10 days had enhanced mucosal growth and increased expression of SGLT1 compared with control rats. These findings indicate that the neointestine is capable of responding to external regulator signals that could be used to further expand the surface of the neointestine. There is a lack of preclinical models for observing intestinal tissue regeneration and improved intestinal function on a scale that can be feasibly translated into humans. Intestinal tissue engineering has been investigated using a large-animal model designed to emulate conditions required for human therapy [46]. Tissue scaffolds were seeded with organoid units isolated from the jejunum of 6-week-old piglets and implanted into the omentum of that animal. However, the study provided only limited information on issues related to the scaling-up of a technique for use in humans because the neointestine was not anastomosed to the native intestine and the scaffolds used were similar in size to those used in previous small-animal models. A functional mucosal barrier is an essential element of intestinal tissue engineering for which scalability also needs to be considered. Although transepithelial resistance of the mucosa created in neointestinal cysts is similar to that of native intestine [85], the creation of larger intestinal constructs will require rapid coverage of the scaffold surface to ensure the barrier function is established. This process might be accelerated by including materials in the scaffold that promote epithelial cell spreading. Yoshida and colleagues investigated the effect of transplanting organoid units onto denuded colonic mucosa of syngeneic recipient rats [104]. The addition of bFGF facilitated neomucosal growth and improved restoration of intestinal epithelial cell coverage over the denuded mucosa compared with the control group. Other approaches might include including inorganic materials into hybrid scaffolds, such as bioactive glass, shown to increase epithelial cell migration via bFGF in an indirect manner [105]. As well as stimulating regeneration of the mucosa, delivery of growth factors may provide a strategy for regenerating the muscularis propria. Local delivery of bFGF from scaffolds, via either incorporation into the collagen coating of scaffolds or encapsulation into microspheres, was also shown to increase the engraftment and density of seeded smooth muscle cells and blood vessel formation after 28 days’ implantation in the omentum of rats [106]. Rapid vascular in growth into the tissue engineered intestine will be essential to maintain the viability and engraftment of cells seeded on the scaffold. Gardner-Thorpe and colleagues observed that tissue engineered intestine exhibited lower levels of bFGF and vascular endothelial growth factor (VEGF) and a fixed capillary density compared with native juvenile bowel [107]. This led that group to evaluate a polymeric microsphere system to deliver encapsulated VEGF and stimulate angiogenesis in the maturing neointestine [108]. Capillary density in the muscular and connective tissue layers was significantly increased in the presence of microspheres containing VEGF, as were the size and weight of the constructs. Interestingly, the rate of epithelial cell proliferation also increased in constructs implanted with VEGF-releasing microspheres, possibly related to the improved vascularization of the construct providing greater nutritional support to the rapidly proliferating epithelium. The need for neovascularization is not restricted to tissue engineering tissues of the alimentary tract. A number of different approaches are being used to tackle this problem [109]. It remains to be seen whether any of these approaches will provide a sufficient stimulus to promote arteriogenesis required for sufficient vascularization of larger tissue constructs. Furthermore, a functional lymphatic system in the neointestine is essential to establish normal nutrient absorption, fluid homeostasis, and immunological functions. Lymphangiogenesis is reported to occur in the neointestine created by the organoid unit-cyst model in rats [110]. Although angiogenesis has been demonstrated in intestinal tissue engineering using small-animal models, it is not certain whether the provision of thin-walled endothelium lined structures will be sufficient to support the functionality of a larger tissue construct. Therefore,

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techniques to promote the formation of medium-sized blood vessels via arteriogenesis are likely to be required to facilitate complete integration of large-scale intestinal constructs with a functional capacity. The small bowel has an extensive vascular system fed by arcades of arteries in the mesentery derived from the superior mesenteric artery. Translation of the existing tissue engineering models to a scale suitable for implantation into humans will require the formation of a similar vascular system consisting of medium-sized blood vessels to maintain viability of a larger tissue construct as well as enable absorption of fluid and dissolved nutrient material from the intestine into the portal blood, which will require a vascular system similar to that found in native intestine. One approach to enabling immediate perfusion of the tissue engineered construct might involve using the existing vascular system in decellularized tissue. Preservation of the vascular structure in decellularized porcine small bowel has been used to engineer tissue via innate vascularization [111]. The decellularized scaffold was repopulated by endothelial cells and exhibited patent vessels after arterial and venous microanastomosis. Improved methods for seeding and maturing larger tissue engineered intestinal constructs will be needed to ensure that the limited cells available are delivered to the tissue construct efficiently and uniformly. These obstacles may be overcome with the development of bioreactor systems that will assist with the long-term culture and bioengineering of tissues by providing an in vitro environment that is similar to normal physiological conditions. Kim and colleagues designed a perfusion bioreactor specifically for intestinal tissue engineering [112]. The use of bioreactors that provide magnetic force may also be used to deliver cells labeled with superparamagnetic iron oxide nanoparticles into hollow tubular scaffolds with more uniform distribution [113]. Techniques used to tissue engineer vascular grafts might also provide solutions that can be translated to intestinal tissue engineering. For example, centrifugal casting onto decellularized laser-porated natural scaffolds has been reported to enable the rapid fabrication of tubular tissue in a bioreactor-free manner [114]. The type of scaffold material chosen for tissue engineering is an important consideration. An optimal scaffold material must be capable of withstanding the intestinal microenvironment, which poses significant challenges in terms of biocompatibility, mechanical properties, and longevity. It must allow transplanted cells to engraft and proliferate rapidly while enabling tissue perfusion of nutrients and remodeling to ensure complete integration with the host. The composition, geometry, and topography of scaffolds used for intestinal tissue engineering may influence the properties of cells grown on their surface. Compared with natural extracellular matrixederived scaffolds, biodegradable synthetic polymer scaffolds provide more control over scaffold properties, such as scaffold architecture, degradation rates, and mechanical properties. Boomer and colleagues evaluated a selection of synthetic tubular scaffolds composed of poly(glycolic acid), poly(ε-caprolactone), poly(L-lactic acid), and polyurethane with either nanofiber or macrofiber structures [115]. Implantation of the scaffolds into the peritoneal cavity of rats revealed different rates of tissue infiltration and scaffold degradation. The inclusion of extracellular matrix components such as hyaluronic acid (a nonsulfated glycosaminoglycan found ubiquitously in connective, epithelial, and neural tissue) was found to enhance the physicochemical properties of gelatinecollagen scaffolds, including attachment, growth, and viability of Caco-2 cells [116]. Histological organization of cells resembling intestinal circular and longitudinal smooth muscle has also been achieved using scaffolds that consist of two layers of orthogonally oriented fibers [117]. The function of the geometry of the cryptevillus microenvironment in regulating intestinal cell proliferation and differentiation was explored by Wang and colleagues [118]. Caco-2 cells migrating over microwell structures showed increased metabolic activity and lower levels of differentiation compared with cells cultured on flat surfaces, which suggested that the structure of crypts may have a role in retaining a proliferative phenotype. Likewise, scaffold architecture is a parameter that can be used to enhance cell infiltration and mass transfer of nutrients to ensure the viability of tissue is maintained. The porosity of the scaffold has a critical role in cell survival and ultimately the viability of the tissue construct. This was illustrated in a study looking at the inclusion of 250-mm pores in multilayered electrospun scaffolds [119]. When implanted in vivo, scaffolds with greater macroporosity were associated with increased blood vessel development and improved survival of intestinal smooth muscle cells, which suggested that macropore connectivity can be optimized to enhance angiogenesis and improve cell viability. Compression molding combined with particulate leaching has been used to create multilayered scaffolds with differential porosities and pore sizes [120]. These structural features were found to influence the behavior and interaction (bridging versus penetration) of different cell types found within the small intestine (epithelial and smooth muscle cells). The impact of the spatial geometry and mechanics of the microenvironment are likely to have an important role in the physiological functionality of tissue engineered constructs. DiMarco and colleagues used a combination of experimental and finite element analysis to investigate critical variables that control intestinal organoid contraction [121]. Adjustment of ambient oxygen concentration, tailoring the density of the collagen type I matrix, addition of R-spondin1, and culture geometry were found to influence contractile behavior of the organoids. Contractile

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behavior occurred only within a narrow range of collagen densities, which the investigators suggested acted as a permissive switch to enable contraction. The inclusion of biomimetic topographical features might provide further control of cell behavior. To replicate the irregular multiscale features of native intestine, Koppes and colleagues created replicas of decellularized porcine small intestine by chemically depositing Parylene C to create molds for polydimethylsiloxane (PDMS) substrates [122]. The PDMS substrates exhibited multiscale folds, crypt and villus structures, and submicron features of the basement membrane. Finer control over the impact of microenvironmental cues is likely to improve scaffold performance, which affects tissue formation, development, differentiation, and functional behavior. Culture of primary intestinal organoids in recombinantly engineered extracellular matrix allowed improved physical manipulation of the scaffolds biomechanical cues while retaining the efficiency of organoid formation to match that obtained with natural collagen I matrices [123]. This provided a means to optimize the properties of the scaffolds to achieve improved matrix performance and identification of microenvironmental cues crucial for bioengineering tissue constructs. Porous protein scaffold systems composed of silk fibroin have also been used to replicate the tissue architecture and microenvironments of the intestine [124]. This type of insight into how properties of the scaffold materials influence the functional behavior of cells is crucial if physiological features such as the stem cell niche, cryptevillus axis, and peristalsis are to be achieved in the tissue engineered construct. Combined with the revolution in additive manufacturing, an icreased understanding of the biochemical and mechanical cues that control tissue regeneration will lead to a step change in state-of-the-art technology available to create scaffolds tailored to replicate the native scaffold of tissues in terms of topographical, mechanical, and biochemical features. Lee and colleagues fabricated scaffolds with a high surface areaeto-volume ratio using 3D printing technology [125]. The growth of smooth muscle cells in vitro was found to be influenced by the geometry of the scaffold. Scaffolds with small villi (0.5 mm) had increased cell density compared with scaffolds containing large villi (1 mm) after 14 days of culture. A 3D printable Matrigeleagarose system was described for the printing of intestinal epithelial cells [126]. The agarose component provided mechanical stability for the 3D printed structure and the Matrigel provides essential protein components for cell growth and spreading. The development of bioinks with improved mechanical properties and biocompatibility will enable additive manufacturing such as 3D printing to make tremendous opportunities for regenerative medicine [127]. Instilling peristaltic activity to the tissue engineered intestine to establish gut motility will require correctly oriented smooth muscle cell regeneration and reinnervation. Advances have been made in technologies to achieve peristalsis through the combination of autologous smooth muscle cells and biomaterials to produce patch or tubular constructs [128]. Maintaining native smooth muscle organization appears to be critical to achieving functionally contracting smooth muscle [129]. Intact smooth muscle strips retained neural and glial markers and exhibited periodic contraction whereas smooth muscle cells cultured after enzymatic digestion did not. Innervation is not only an unmet need for bioengineering gastrointestinal tissue constructs. Enteric neuropathies such as achalasia, gastroparesis, intestinal pseudoobstruction, and chronic constipation are functional gastrointestinal disorders that result from primary and secondary forms of degenerative disease that affect the nerves and muscles in the gastrointestinal tract [130]. Mouse enteric neural crest cells transplanted into aganglionic gut spread along the endogenous myenteric plexus to form functionally integrated branching networks closely associated with endogenous neural glial networks, providing evidence for the use of enteric neural stem cell therapies [131]. The appendix, a vestigial organ, might provide a potential source of autologous neural stem cells for enteric neural cell therapy [132]. Geisbauer and colleagues investigated whether a mixture of enteric cells isolated from longitudinal and circular muscle of the gut could be used as a potential source of neural crest stem cells for cell therapy. When the isolated cells were mixed in collagen containing bFGF and injected into an aganglionic segment of jejunum, ganglion-like structures were generated with elongated synapses [133]. Innervation of tissue engineered constructs is fundamental to achieving or restoring gastrointestinal transit. Colon smooth muscle cells cultured in composite chitosan scaffolds were innervated by differentiated functional neurons derived from cocultured neural progenitor cells [134]. Likewise, human smooth muscle cells and neural progenitor cells were engineered into innervated sheets of smooth muscle and wrapped around tubular chitosan scaffolds [135]. After subcutaneous implantation, the construct became vascularized and the lumenal patency was maintained. In addition to regulating peristalsis, the enteric nervous system in the intestine controls villi activity and the modulation of secretions from gut epithelial cells. Gut endocrine cells has an important role in regulating gastrointestinal activity by releasing serotonin, secretin, cholecystokinin, gastrin, and enteroglucagon and will be an essential component of the tissue engineered intestine. Nakase and colleagues investigated the regeneration of endocrine cells and the nerve system in a canine patch model of tissue engineered small intestine using a collagen sponge scaffold loaded with autologous gastric smooth muscle cells [136]. At 24 weeks after implantation of the scaffolds into the

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middle of an isolated ileal loop, the location and number of endocrine cells staining positive for chromogranin A were almost identical to those of native mucosa. Nerve fibers were present in the regenerated smooth muscle layer and villi, but the myenteric plexus of Auerbach and the submucosal plexus of Meissner were not visible. The density of smooth muscle cells implanted into the scaffolds did not affect the thickness of the regenerated smooth muscle layer, which remained approximately half that of the native smooth muscle layer, indicating that other cues will be necessary to increase its thickness. The investigators suggested that the thickness of the muscle layer might be limited by the blood supply available to the regenerating tissue, which might be increased by the delivery of angiogenic factors from the scaffold. Grikscheit and colleagues also reported that ganglion cells were distributed in the locality of the Auerbach and Meissner’s plexuses in tissue engineered small intestine [90]. Regeneration of the small intestine remains at an early stage of clinical development and has yet to provide a clear demonstration of improvement in nutrient absorption that will be valuable in a clinical therapeutic setting in humans. Although no models have unequivocally demonstrated functional neointestine with peristaltic activity, they indicate that it is feasible to engineer tubular segmental replacement of small bowel that incorporates innervated smooth muscle layers. Based on these findings, it may be possible to achieve incremental steps toward tissue engineering the intestine. For example, combining existing and established surgical procedures with the principles of tissue engineering and regenerative medicine may improve existing clinical outcome measures. An example of this type of approach was demonstrated by Nakao and colleagues, who showed that the longitudinal staples used during Bianchi’s procedure could be replaced with SIS graft [137]. Refinement of other existing techniques could yield further advances toward viable options for tissue engineered intestinal constructs.

COLON The colon is an important organ for water and sodium resorption and a storage pouch for waste products. Patients who undergo total colectomy are at risk for significant morbidities [138]. The surgical creation of an ileal pouch to create a reservoir provides only a limited solution and patients may still experience inflammation of the pouch (pouchitis), malabsorption, diarrhea, cramping abdominal pain, and fever [139]. Tissue engineering approaches similar to those used for the small intestine have been applied to the colon [140]. Consequently, many of the same challenges exist. Tissue engineered colon was achieved by seeding organoid units harvested from the sigmoid colon of neonatal Lewis rats, adult rats, and tissue-engineered colon itself onto a polymer scaffold that was implanted into the omentum of syngeneic adult Lewis rats. Tissue-engineered colon was generated from each of the donor tissue sources and the resulting architecture of the neocolon was similar to that of native tissue. When anastomosed to the native bowel, there was gross evidence of fluid absorption by the tissue engineered colon. The choice of scaffold material used for colon tissue engineering will have an impact on the viability of the neocolon. Two of the main biological scaffold materials explored in gastrointestinal tissue engineering are SIS and chitosan. Relatively few studies have made direct comparisons of these materials in terms of their biocompatibility. Denost and colleagues looked at the in vitro and in vivo properties of two bioscaffolds composed of these materials [141]. No substantial difference was observed in vitro in terms of cell attachment and proliferation, but the chitosan hydrogel facilitated improved healing of a preclinical in vivo model of colonic wall defect, including regeneration of the smooth muscle layer. A significant drawback reported with biological scaffolds such as chitosan is their weak mechanical properties. To enhance mechanical strength, Zakhem and Bitar reported using chitosan fibers circumferentially aligned around tubular chitosan scaffolds [142]. Tensile strength and strain, burst pressure, and Young’s modulus were all increased in scaffolds that contained the fibers.

ANAL CANAL Controlled storage and timely disposal of feces relies largely on the appropriate function of sphincter muscles that constrict the anal canal and maintain fecal continence. Fecal incontinence is a common disease, particularly in aging societies in which it has a huge impact on quality of life and incurs colossal health costs. Conservative estimates indicate that approximately 2% of community-dwelling adults experience regular fecal incontinence [143]. This figure increases to 50% in the institutionalized and geriatric population [144]. Conservative treatments for fecal

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incontinence are ineffective in patients with more than mild symptoms and surgical interventions produce poor long-term benefit, with frequent complications [145]. Despite holding considerable promise, cell therapy for incontinence affecting the alimentary tract remains relatively unexplored in humans [146,147]. Although the technical feasibility of injecting autologous myoblasts for treating fecal incontinence in humans has been demonstrated, these studies have been unable to demonstrate integration of cells into the damaged sphincter or a direct improvement in the functional integrity of the sphincter muscle. A fibrin-based bioengineered in vitro model of the internal anal sphincter (IAS) was described that demonstrated physiological functionality. This type of approach is likely to offer value in studying complex physiological mechanisms underlying sphincter malfunction [148,149]. The bioengineered sphincter was surgically implanted into the subcutaneous tissues of syngeneic mice and responded to the local delivery of bFGF, resulting in improved muscle viability, vascularization, and survival of the graft [150]. This field of research has continued to advance; the first study demonstrated the feasibility of transplanting bioengineered intrinsically innervated human IAS into mice. Isolated human IAS smooth muscle cells were cocultured with immortomouse fetal enteric neurons. The construct implanted into RAG1/ mice became neovascularized and the physiological function of the myogenic and neuronal components was retained. The IAS construct exhibited characteristics of IAS physiology [151]. The feasibility of implanting a bioengineered sphincter construct consisting of human IAS smooth muscle cells was explored in the perianal region of athymic rats [152]. The constructs were well-tolerated and the recipients were able to produce stool normally. For this study, vascularization was increased by delivering platelet-derived growth factor. In addition to the anal sphincter, proof-of-concept studies have demonstrated that it is feasible to bioengineer autologous bioengineered innervated pylorus constructs that consist of circumferentially aligned smooth muscle cells that exhibit tonic contractile phenotype and basal tone [153]. However, the feasibility of scaling-up this type of approach from a rodent model to humans remains uncertain. Although it is technically possible to bioengineer rings of muscle in vitro on a scale comparable to that of human sphincter muscle [154], innervation, vascularization, and cell viability in larger constructs have yet to be tested in larger-sized preclinical models. Regenerative medicine may also offer solutions to conditions in which existing medical and surgical procedures have failed. A condition in which this affects the alimentary tract is perianal fistulas that result from a connection between the anal canal and the perianal skin surface, creating an abnormal passageway for the discharge of pus, blood, and in some cases feces, resulting in significant morbidity. The goals of fistula treatment are eradication of perineal sepsis and fistula closure while posing a minimal risk for causing sphincter muscle damage. A difficulty in treating perianal fistulas is avoiding abscess formation caused by healing of the skin before closure of the tract. To address this, collagen anal fistula plugs have been devised to treat fistulas. Although early studies reported good healing rates with little or no risk to continence, long-term follow-up has revealed variable and disappointing success rates (24e78%) [155]. Reports of the plugs failing owing to dislodgment from the tracts indicate that this approach may not provide an ideal scaffold material to promote guided tissue regeneration and closure of the tract [155]. A possible solution to this problem is the use of scaffold materials that provide both optimal conditions for rapid cell infiltration when implanted into tissue cavities and mechanical strength to maintain an open scaffold structure [156].

IN VITRO MODELS Improved understanding of gastrointestinal developmental biology opens up new opportunities for creating 3D tissue constructs that can be used to modeling disease, understand embryonic development, and provide construct sources for therapeutic applications [157]. Principles of regenerative medicine are increasingly being used to fabricate biomimetic models of the gastrointestinal tract. Challenges that exist with developing in vitro tissue models of gastrointestinal tissue include mimicking the 3D microenvironment, interactions among different cell types, and the microbiome. New technologies are being applied to address these, including using microfluidics to create channels lined by living cells in microengineered biomimetic systems that might offer new opportunities to replace conventional animal models in preclinical toxicology testing. This approach has been applied to a variety of organs including the intestine to provide organs-on-chips that exhibit physiological properties including peristalsis-like movement [158]. Dynamic culture in a defined perfusion bioreactor has also been reported to result in tissue models that are physiologically closer to native small intestine [159]. Microfluidic cell culture devices have been designed that contain villi- and crypt-like structures that resulted in epithelial cells tightly connecting to each other and displaying absorption and paracellular transport function [160]. The incorporation of physiological parameters also

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appears to be important for establishing 3D in vitro tissue models of small intestine. Bioreactors used to culture decellularized segments of porcine jejunum have been used to coculture human Caco-2 cells with human microvascular endothelial cells. Compared with routine static Caco-2 assays, culture under dynamic conditions resulted in cell morphology that more closely resembled normal primary enterocytes [161]. Recapturing essential features of the cellular microenvironment is essential if in vitro tissue engineered models are to be used for functional studies such as cell growth, differentiation, absorption, or hostemicrobial interactions. The inclusion of native features such as accurately sized intestinal villi has been shown to facilitate cell differentiation along the villous axis [162]. Methods used to realize such structures include 3D natural and synthetic hydrogels created using a combination of laser ablation and sacrificial molding to achieve microscale structures that mimic the density and size of human intestinal villi [163]. The microbiome of the gastrointestinal tract is increasingly being recognized as a critical component to maintaining physiological homeostasis. Biomimetic in vitro intestinal models for investigating the adhesion and invasion profile of commensal and pathogenic organisms therefore have significant value in understanding microbe-induced intestinal disorders. To explore this interaction, synthetic 3D tissue scaffolds that support coculture of epithelial cell types have been used to provide microbial niches along the cryptevillus axis for modeling the interaction of a variety of commensal and pathogenic organisms [164].

CONCLUSION The alimentary tract is a complex organ that is essential for maintaining physiological homeostasis. Tissue engineering and regenerative medicine for hollow visceral organs have been proposed as an approach for replacing damaged or diseased tissue, as demonstrated in humans with bladder [165] and airway tissue [15]. Although these “first-in-human” studies have rightly attracted much attention, there is a long way to go until a similar approach becomes routine in the relatively complex structures of the alimentary tract. The past few decades have delivered a series of important studies that have used the innate ability of the alimentary tract to regenerate. Further studies are needed to demonstrate whether these approaches are transferable and of clinical value to humans. Fundamental challenges such as scalability will need to be addressed to enable the results obtained in small-animal models to progress into preclinical models applicable to humans. This will require refinement of the scaffolds used and the ability to seed limited quantities of cells available in an efficient manner onto the scaffolds. These are challenges that can be overcome and will allow regenerative medicine to progress in the alimentary tract in humans.

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Gastroenterology July 2009;137(1):53e61. [150] Hashish M, Raghavan S, Somara S, Gilmont RR, Miyasaka E, Bitar KN, et al. Surgical implantation of a bioengineered internal anal sphincter. J Pediatr Surg January 2010;45(1):52e8. [151] Raghavan S, Gilmont RR, Miyasaka EA, Somara S, Srinivasan S, Teitelbaum DH, et al. Successful implantation of bioengineered, intrinsically innervated, human internal anal sphincter. Gastroenterology July 2011;141(1):310e9.

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[152] Raghavan S, Miyasaka EA, Gilmont RR, Somara S, Teitelbaum DH, Bitar KN. Perianal implantation of bioengineered human internal anal sphincter constructs intrinsically innervated with human neural progenitor cells. Surgery April 2014;155(4):668e74. [153] Rego SL, Zakhem E, Orlando G, Bitar KN. Bioengineered human pyloric sphincters using autologous smooth muscle and neural progenitor cells. Tissue Eng January 2016;22(1e2):151e60. [154] Parmar N, Day RM. Appropriately sized bioengineered human external anal sphincter constructs. Surgery January 2015;157(1):177e8. [155] Adamina M, Hoch JS, Burnstein MJ. To plug or not to plug: a cost-effectiveness analysis for complex anal fistula. Surgery January 2010; 147(1):72e8. [156] Blaker JJ, Pratten J, Ready D, Knowles JC, Forbes A, Day RM. Assessment of antimicrobial microspheres as a prospective novel treatment targeted towards the repair of perianal fistulae. Aliment Pharmacol Ther September 1, 2008;28(5):614e22. [157] Howell JC, Wells JM. Generating intestinal tissue from stem cells: potential for research and therapy. Regen Med November 2011;6(6): 743e55. [158] Huh D, Kim HJ, Fraser JP, Shea DE, Khan M, Bahinski A, et al. Microfabrication of human organs-on-chips. Nat Protoc November 2013;8(11): 2135e57. [159] Schweinlin M, Wilhelm S, Schwedhelm I, Hansmann J, Rietscher R, Jurowich C, et al. Development of an advanced primary human in vitro model of the small intestine. Tissue Eng September 2016;22(9):873e83. [160] Chi M, Yi B, Oh S, Park D-J, Sung JH, Park S. A microfluidic cell culture device (mFCCD) to culture epithelial cells with physiological and morphological properties that mimic those of the human intestine. Biomed Microdevices 2015;17(3):9966. [161] Pusch J, Votteler M, Go¨hler S, Engl J, Hampel M, Walles H, et al. The physiological performance of a three-dimensional model that mimics the microenvironment of the small intestine. Biomaterials October 2011;32(30):7469e78. [162] Costello CM, Hongpeng J, Shaffiey S, Yu J, Jain NK, Hackam D, et al. Synthetic small intestinal scaffolds for improved studies of intestinal differentiation. Biotechnol Bioeng June 2014;111(6):1222e32. [163] Sung JH, Yu J, Luo D, Shuler ML, March JC. Microscale 3-D hydrogel scaffold for biomimetic gastrointestinal (GI) tract model. Lab Chip February 7, 2011;11(3):389e92. [164] Costello CM, Sorna RM, Goh Y-L, Cengic I, Jain NK, March JC. 3-D intestinal scaffolds for evaluating the therapeutic potential of probiotics. Mol Pharm July 7, 2014;11(7):2030e9. [165] Atala A, Bauer SB, Soker S, Yoo JJ, Retik AB. Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet April 15, 2006; 367(9518):1241e6.

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65 Extracorporeal Renal Replacement Christopher J. Pino1, H. David Humes1,2 1

Innovative Biotherapies, Ann Arbor, MI, United States; 2University of Michigan, Ann Arbor, MI, United States

INTRODUCTION The kidneys are vital organs that have critical excretory and regulatory roles in maintaining water and electrolyte homeostasis in the body while removing lowemolecular weight (MW) toxins. Often overlooked, the kidneys also provide an essential endocrine function because they secrete important signaling molecules that interact with the cardiovascular, skeletal, and immune systems. The kidney was the first organ to be successfully transplanted and is one of the most common and effective transplantation surgeries performed in the world. However, global shortages of donor kidneys have proven to be a major barrier to treating kidney disease, leaving millions of chronic kidney disease patients without total functional replacement. In the United States, 636,905 patients were listed as having end-stage renal disease (ESRD) at the end of 2012, as reported by the 2014 US Renal Data System (USRDS) database. According to the USRDS report, over 450,000 patients were receiving maintenance dialysis, 186,303 patients had a functioning kidney transplant, and 88,638 ESRD patients died during 2012 [1]. Adjusted rates of all-cause mortality are 6.1e7.8 times greater for dialysis patients than for individuals in the age-matched Medicare population [1]. The incidence rate of ESRD has plateaued and slightly decreased since 2009, with an adjusted incidence rate of 353 per million per year in 2012 [1]. The financial cost of ESRD is immense; spending for ESRD patients increased 3.2% to $28.6 billion in 2012, accounting for 5.6% of the Medicare budget [1]. A prior USRDS report estimated the cost of $54,900 per hemodialysis (HD) patient per year and $46,121 per peritoneal dialysis (PD) patient per year [2]. In contrast, transplant patients cost an average of $17,227 per patient per year [2]. The higher cost of maintenance dialysis compared with transplantation does not translate into better results; annual mortality for patients listed for transplant and awaiting a kidney is 6.3%, compared with only 3.8% for patients listed for transplant who received a kidney [2]. Although organ transplantation provides the best prognosis for survival, demand vastly outweighs the availability of donated organs. In 2014, the active waiting list was 2.8 times larger than the supply of donor kidneys [3]. Extracorporeal therapies to replace kidney function at least partially include HD, hemofiltration (HF), hemodiafiltration (HDF), and PD. These therapies replace the blood filtration function of the kidney by removing waste products and excess electrolytes via artificial or natural semipermeable membranes in extracorporeal circuits. These methodologies all address water and electrolyte balance as partial functional replacement of the kidney but are prone to various complications such as acidosis, volume overload, and uremic syndromes that accompany renal failure [4]. Conventional therapies fail to provide the lost endocrine functions of the kidney, and thus the metabolic, endocrine, and immunological roles of the functioning kidney are potential mechanisms for the difference in survival for patients who receive kidney transplants.

REQUIREMENTS OF A RENAL REPLACEMENT DEVICE In the kidney, HF is accomplished by nephrons that generate ultrafiltrate (UF) from blood via glomeruli. Each glomerulus, a tuft of capillaries supported by a basement membrane and specialized epithelial cells called podocytes, is an efficient filter with a molecular weight cutoff (MWCO) of around 65 kDa. This allows for small molecules including small MW toxins such as urea to be removed from the bloodstream, whereas the cellular component of Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00065-5

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blood and larger proteins including albumin are retained. UF is passed into the renal proximal tubule portion of the nephron, a hollow tube of cells surrounded by capillaries, which receives the filtrate from the glomerulus and accomplishes the bulk of reclamation of salt, water, glucose, small proteins, amino acids, glutathione (GSH), and other substances. A renal replacement device thus must recapitulate the filtration function of the nephron. However, in addition to its role in watereelectrolyte balance, the tubule also performs metabolic functions, including excretion of acid as ammonia and hydroxylation of 25-hydroxy-vitamin-D3 among other metabolic and regulatory functions, which are not supplemented in conventional renal replacement therapies (RRT). The kidney is responsible for the secretion of hormones that are critical in maintaining hemodynamics (renin, angiotensin II, prostaglandins, nitric oxide, endothelin, and bradykinin), red blood cell production via erythropoietin, and bone metabolism (1,25-[OH]2evitamin D3 or calcitriol) [5]. Free-radical scavenging and GSH-metabolizing enzymes are synthesized by the kidney, which also provide gluconeogenic and ammoniagenic capabilities [6,7]. Catabolism of low -molecular weight proteins, including multiple peptide hormones, cytokines, and growth factors, is also accomplished by the kidney [8]. The kidney also has a less-recognized immunoregulatory function. Mammalian renal proximal tubule cells (RPTC) are immunologically active. RPTCs are antigen-presenting cells [9] that have costimulatory molecules [10] that synthesize and process a variety of inflammatory cytokines [10,11]. Long overlooked, the kidney is an important immunomodulatory organ. RRT has poor patient prognosis owing to the acute tubular necrosis (ATN) that results in the loss of the kidney’s immunoregulatory function. Loss of immunoregulation results in a propensity to develop systemic inflammatory response syndrome (SIRS), sepsis, multiple organ dysfunction (MOD), and a high risk for death because of systemic immunologic or inflammatory imbalance. The endocrine and immunologic roles of kidney cells are also important in maintaining kidney and other vital organs’ health under stress conditions [12] and are not provided by conventional RRT.

DEVICES USED IN CONVENTIONAL RENAL REPLACEMENT THERAPY Water and solute homeostasis function of the kidney can be practically replaced by various dialysis or filtration methods including HD, HF, HDF, and PD. In these processes, synthetic membranes are used to remove water and solutes selectively based on diffusive or convective transport. In HD, in which solute removal from the blood is driven by a concentration gradient across a membrane, the process is dominated by diffusion. However, solute transfer also occurs to a lesser extent, convectively by a process of ultrafiltration, in which water and solutes move across the membrane [13]. This process is leveraged further in HF, in which pores in semipermeable membranes allow water and solutes across the membrane independent of the concentration gradient, and predominantly on the hydraulic pressure gradient across the membrane, which mimics the filtration function of the nephron. HDF combines HD and HF processes, in which both diffusion and hydraulic pressure are used as driving forces. In PD, similar to HD, solute removal is mainly driven by diffusion across a synthetic membrane; however, instead of solutes being removed directly from blood, solutes are removed from the peritoneal fluid of the patient, which avoids complications of blood access. PD has a higher safety profile, which enables dialysis treatments at home, overnight dialysis, and continuous ambulatory dialysis. In HD, blood exits the patient’s body through a catheter, which is delivered to a dialyzer via tubing and a blood pump (Fig. 65.1A). These dialyzers are commonly filled with hollow fiber membranes formed into a bundle. Within a hollow fiber dialyzer, blood flows through the lumen, the interior of each hollow fiber. Solutes in the blood interact with the fiber’s semipermeable membrane, which allows for low-MW solutes such as urea and creatinine to pass through the porous fiber wall into the extracapillary space (ECS) of the dialyzer for removal, whereas blood cells and critical large proteins such as albumin and immunoglobulin are retained in the blood. Dialysate, an aqueous solution administered during dialysis, flows opposite the blood on the other side of the semipermeable membrane in the ECS. The dialysate has various concentrations of solutes that dictate the driving force to remove, retain, or add to solute concentrations in the blood. Movement of the solute is directed from high to low concentration so that toxins are removed and critical electrolytes in the blood are maintained. In HF, blood is filtered using a pressure difference to drive water and solutes across a semipermeable membrane and into the filtrate compartment. Like dialyzers, hemofilters are routinely filled with hollow fibers. Positive pressure can be applied to the hemofilter’s blood compartment (lumen side of the hemofilter) or a negative pressure can be applied to the filtrate compartment (ECS), creating a transmembrane pressure. No dialysate solution is used in HF; rather, the filtrate is collected and discarded and ultrapure replacement solution is administered to the patient to maintain wateresolute balance, as shown compared with dialysis in Fig. 65.1B.

ADVANCEMENTS IN CONVENTIONAL RENAL REPLACEMENT THERAPY DEVICES

FIGURE 65.1

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Diagrams of hemodialysis (A) and conventional venovenous hemofiltration (B) methods. UF, ultrafiltrate.

Membrane materials used in dialyzers and hemofilters are diverse, ranging from metals and ceramics to polymers. Critical to the choice of materials for dialysis membranes, biocompatibility with blood is a major concern, as are cost and manufacturing. Synthetic polymers have been the dominant membrane materials used for separation processes owing to combination of these factors. The most common polymers in the manufacturing of synthetic membranes are nondegradable ones, including cellulose derivatives such as cellulose acetate, nitrates, and esters, polysulfones, and their ether-containing derivatives polyethersulfones and polyarylethersulfone, polyacrylonitrile (PAN) derivatives, polyamides, polyimides, polyolefins such as polyethylene and polypropylene, polyvinylchloride, and fluorinated polymers such as polytetrafluoroethylene, and polyvinylidine fluoride.

ADVANCEMENTS IN CONVENTIONAL RENAL REPLACEMENT THERAPY DEVICES Besides improvements in biocompatibility, synthetic membranes have been developed to better reproduce the physiologic process of glomerular ultrafiltration. In an attempt to recapitulate glomerular ultrafiltration and removal of “middle molecules,” synthetic membranes with larger pore sizes and high water permeability have been developed. These “high-flux” membranes are commonly prepared with hydrophobic base materials, along with various hydrophilic components. These membranes can be mass produced, with wide ranges of hydraulic permeabilities for various treatments [14]. Pore size, therefore, is a critical issue in hemofilter design. Conventional polymeric membranes are typically cast or spun from polymers in solution, resulting in a heterogeneous dispersion of pore sizes and geometries within fibers or sheets, with ill-defined MWCO. New techniques using a nanoscale spinning process to control pore size, resulting in an increased uniformity of pores with desired values [15]. Based upon this technology, the main focus of recent membrane development has been to increase pore size while sharpening the MWCO of high-flux membranes to maximize removal of low MW proteins. This direction is based upon the idea that removal of a distinct class of uremic toxins, such as b2-microglobulin and myoglobin, while minimizing the loss of albumin could improve treatment outcomes of patients with ESRD. Superflux or protein-leaking membranes, provide greater clearances for low MW proteins and small protein-bound solutes, but with significantly higher loss of albumin than high-flux dialysis membranes. Other determinants of dialyzer performance may include: fiber bundle configuration, spacing, and sterilization techniques [16]. Membranes typically used today are purely artificial and are used in the form of hollow fiber modules, commonly based on cellulose or synthetic polymers, including polysulfone, PAN, and polyamide, used as such or modified with a variety of agents [17e19]. However, research in biomaterials and tissue engineering holds promise for significant future improvements. Improvements in membrane biocompatibility have reduced inflammatory reactions, but

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off-the-shelf polymeric materials still induce a cascade of biochemical events leading to blood coagulation, thrombus formation, and often complement activation and inflammation. Systemic administration of anticoagulant such as heparin is required to solve these problems. However, the use of anticoagulants, may lead to additional complications such as hemorrhage, and thus researchers are trying to improve anticoagulation therapy by finding alternative antithrombotic agents such as hirudin [20] and methods of administration, such as the direct incorporation of bioactive agents onto the polymeric membrane itself [21], such as vitamin E or heparin [22]. Conventional membranes remove solutes only on the basis of molecular size, whereas the glomerulus also employs charge selectivity by the basement membrane. This nonselective nature of conventional membranes is another limitation to overcome. One possible solution to this problem is to grow cells on the hemofilter membrane to improve the selectivity of the membrane. Another approach is the design of “smart” membranes that have selective transport characteristics and contain a predetermined number and size of nonpassive nanoengineered pores that have specific interactions with solutes and solvents at the atomic level [23]. Despite improvements in membrane technologies allowing for more efficient dialysis, HF, or diafiltration therapies, all are suboptimal treatments for both acute kidney injury (AKI) and ESRD. These therapies do not meet the medical need to reduce mortality caused by AKI. Even in ESRD, the outcomes of patients receiving chronic dialysis therapy are still disappointing, with an annual mortality exceeding 25%, on average and a drastically shortened life expectancy of only 5 years [24]. Several groups have reported that the survival of critically ill patients with AKI could be improved by intensifying the dose of RRT [25e27], but the initial excitement generated by these reports has waned after studies demonstrated that the dialysis dose does not closely relate to outcomes [28e30].

RENAL ASSIST DEVICE: A MORE COMPLETE RENAL REPLACEMENT THERAPY Complete RRT should replace metabolic and endocrine functions of the kidney not supplied by conventional RRT. One such approach, called the renal assist device (RAD), used living cells supported on synthetic scaffolds [31]. In this bioengineered device, nondegradable scaffolding materials were employed to provide physical support for renal cells grown within the device. The RAD was constructed using renal tubule progenitor cells [32,33] cultured on semipermeable polymeric hollow fiber membranes (polysulfone) on which an extracellular matrix was layered to enhance the attachment and growth of the epithelial cells [34] (Fig. 65.2A). These porous hollow-fiber synthetic membranes not only provide the architectural scaffold for these cells and allow for delivery of nutrients to support the cells, they provide immunoprotection during therapy, preventing potential immunologic reactions in patients or compromise of the cell device by the patient’s immune system. With appropriate membranes, immunoprotection of cultured progenitor cells can be achieved concurrent with long-term functional performance as long as conditions support tubule cell viability [31]. In vitro experiments have tested transport and metabolic functions of the RAD using renal proximal tubule progenitor cells grown inside the lumen of hollow fiber devices with membrane surface areas of 97 cm2 to 0.7 m2, resulting in a device containing up to 108e109 cells [35]. The nonbiodegradability and pore size of the hollow fibers allowed the membranes to act as both scaffolds for the cells and an immunoprotective barrier. Confluent monolayer cells within the RAD exhibited morphological characteristics typical of differentiated tubule epithelia including apical microvilli, endocytic vesicles, and tight junctional complexes, as well as differentiated active transport properties, differentiated metabolic activities, and important endocrine processes [35]. RPTC cultured in the RAD were found to maintain their ability to synthesize and excrete ammonia, produce 1,25-(OH)2evitamin D3 (the active form of vitamin D), and, through metabolic degradation, remove GSH from the perfusate. Taken together, these results suggested that the RAD had the ability to replicate the major differentiated transport, metabolic, and secretory functions performed by the healthy renal proximal tubule. To assess the potential clinical translation of this technology, large-animal studies were undertaken using an RAD consisting of renal proximal tubule progenitor cells. In brief, dialysate from a conventional dialysis circuit can be passed into the RAD as a nutrient and oxygen source for the cells grown on the membrane. The membrane is both water- and solute-permeable, allowing for differentiated vectorial transport and metabolic and endocrine activity provided by the cells, and processed dialysate is discarded. Immunoprotection of cultured progenitor cells is achieved concurrent with long-term functional performance as long as conditions support tubule cell viability. In studies with acutely uremic dogs after bilateral nephrectomies, the RAD demonstrated replacement of filtration, transport, metabolic, and endocrine functions of the kidney [35e37]. Animals were treated with either an RAD or a sham control cartridge daily for either 7 or 9 h for 3 successive days or for 24 h continuously. Fluid and small solutes, including blood urea nitrogen (BUN), creatinine, and electrolytes, were adequately controlled in both groups, but

RENAL ASSIST DEVICE THERAPY OF ACUTE KIDNEY INJURY CAUSED BY SEPSIS

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FIGURE 65.2 Depiction of renal assist device (RAD) (A), selective cytopheretic device (B), and bioartificial renal epithelial cell system

(C) devices.

potassium and BUN levels were more easily controlled by RAD treatment. Furthermore, active reabsorption of Kþ, HCO3  , and glucose and excretion of ammonia were accomplished only in RAD treatments. Glutathione reclamation from UF exceeded 50% in the RAD. Finally, uremic animals receiving cell therapy attained normal 1,25-(OH)2e vitamin D3 levels, whereas sham treatment resulted in a further decline from the already low plasma levels. Thus, these experiments clearly showed that the combination of an HF cartridge and an RAD in an extracorporeal circuit successfully replaced filtration, transport, and metabolic and endocrinological functions of the kidney in acutely uremic dogs [35e37]. Additional animal experiments in a porcine model investigated the impact of treatment with this device on the high mortality of sepsis complicated by AKI [38].

RENAL ASSIST DEVICE THERAPY OF ACUTE KIDNEY INJURY CAUSED BY SEPSIS To examine the impact of cell therapy on the course of sepsis complicated by renal failure [38,39], a porcine model of septic shock was developed from the previous work [40,41]. Purpose-bred pigs were anesthetized and administered an intraperitoneal dose of bacteria, causing shock and renal failure. An hour later, continuous venovenous hemofiltration (CVVH) was initiated with either cell or sham RAD. Urine output and mean arterial pressure (MAP) declined within the first few hours after insult. Cell-treated animals survived 9.0  0.83 h versus 5.1  0.4 h (P  .005) for sham-treated animals. Serum cytokines were similar between the two groups, with the striking exception of interleukin (IL)-6 and interferon (IFN)-gamma. Treatment with the cell RAD resulted in significantly lower plasma levels of both IL-6 (P  .04) and INF-gamma (P  .02) throughout the experimental time course compared with sham RAD exposure. This controlled trial of cell therapy of renal failure in a realistic animal model of sepsis had several findings not immediately expected from a priori assumptions regarding renal function. Previously, although renal failure was strongly associated with a poor outcome in hospitalized patients, and chronic renal failure was associated with specific defects in humoral and cellular immunity, a direct immunomodulatory effect of

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the kidney had not been accepted. In this trial, clear differences in survival and clear differences in a serum cytokine associated with mortality in sepsis were found between groups. The increased mortality in renal failure appears to be not attributable to inadequate solute clearance, but may arise from other bioactivity of the kidney. Studies in both dogs and pigs have demonstrated that RAD treatment in a bioartificial kidney circuit improves cardiovascular performance associated with changes in cytokine profiles and confers a significant survival advantage in septic or endotoxin animal models [38,39,42].

CLINICAL EXPERIENCE WITH A RENAL ASSIST DEVICE TO TREAT ACUTE KIDNEY INJURY Encouraging preclinical data led to US Food and Drug Administration (FDA) approval for an investigational new drug and Phase I/II clinical trials. The first human clinical study of the bioartificial kidney containing human cells was carried out in 10 intensive care unit (ICU) patients with AKI receiving CVVH [43]. This study demonstrated that the RAD can be used safely for up to 24 h. Cardiovascular stability was maintained and increased native renal function, as determined by elevated urine outputs, was temporally correlated with RAD treatment. All patients were critically ill with AKI and multiple organ failure (MOF), with predicted hospital mortality rates between 80% and 95%. Six of the ten treated patients survived past 30 days, with mortality reduced to 40%. The human renal tubule cells contained in the RAD demonstrated differentiated metabolic and endocrinologic activity in this ex vivo situation, including GSH degradation and endocrinologic conversion of 25-OHevitamin D3 to 1,25-(OH)2evitamin D3. Plasma cytokine levels suggest that RAD therapy produces dynamic and individualized responses in patients, depending on their unique pathophysiologic conditions. For the subset of patients who had excessive proinflammatory levels, RAD treatment resulted in significant declines in granulocyte-colony stimulating factor (G-CSF), IL-6, IL-10, and especially the IL-6/IL-10 ratio, suggesting a greater decline in IL-6 relative to IL-10 levels and a less proinflammatory state. These favorable Phase I/II trial results led to a randomized, controlled, open-label Phase II trial conducted at 12 clinical sites in the United States [30]. A total of 58 patients with ARF requiring CVVH in the ICU were randomized (2:1) to receive CVVH plus RAD (n ¼ 40) or CVVH alone (n ¼ 18). Despite the critical nature and life-threatening illnesses of the patients enrolled in this study, the addition of the RAD to CVVH resulted in a substantial clinical impact on survival compared with the conventional CVVH treatment group. RAD treatment for up to 72 h promoted a statistically significant survival advantage over 180 d of follow-up in ICU patients with AKI and demonstrated an acceptable safety profile. Cox proportional hazards models suggested that the risk for death was approximately 50% of that observed in the CVVH-alone group. A follow-up Phase IIb study to evaluate a commercial manufacturing process was not completed owing to difficulties with the manufacturing process and clinical study design, which identified two key missing components for successful commercialization of cell-based therapy: (1) a reliable cell source to manufacture thousands of cell devices, and (2) a cost-effective storage and distribution process for cell devices.

IMMUNOMODULATORY EFFECT OF THE RENAL ASSIST DEVICE As described earlier, RAD treatment altered systemic circulating cytokine levels in animal and human experiments. In endotoxin-challenged and gram-negative peritonitis uremic dog models, plasma levels of IL-10 were significantly higher in RAD-treated animals [38,39]. The role of IL-10 in regulating immune response continues to be elucidated, but data suggest that IL-10 levels influence outcome in endotoxin shock and gram-negative sepsis. Several reports have demonstrated that administration of recombinant IL-10 is protective against gram-negative septic shock in murine sepsis models [44,45]. Another study in a similar model demonstrated that administration of antibodies to IL-10 was associated with higher mortality [46]. The mechanism underlying the link between proximal tubule function and IL-10 levels remains to be detailed, but preliminary data suggest that renal production of IL-6 induces liver production of IL-10 [47]. In gram-negative septic pigs without nephrectomy, RAD treatment significantly reduced plasma circulating levels of IL-6 and INF-gamma [42]. The difference in IL-6 concentrations is especially noteworthy because the plasma elevations of this proinflammatory cytokine have been directly correlated to outcome in patients with SIRS [48]. The lower concentration of plasma INF-gamma may be important owing to its central role in the inflammatory response. INF-gamma stimulates B-cell antibody production, enhances polymorpholeukocyte phagocytosis, and activates monocytes and macrophages to release proinflammatory cytokines

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[49e51]. Excessive rates of INF-gamma production by natural killer cells have correlated with progression to lethal endotoxin shock in mice [52]. Further support for an immunomodulatory role of renal tubule cells has been suggested in the Phase I/II clinical trial of the RAD containing human renal tubule cells [53]. Patients treated in this study had a wide spectrum of plasma cytokine levels. The subset of patients who presented with very high plasma cytokine levels and who were treated for an adequate period showed that RAD treatment resulted in significant reductions in G-CSF, IL-6, and IL-10 levels. The greater relative reduction in the IL-6/IL-10 ratio suggests that renal tubule cell therapy may rebalance the excessive proinflammatory response with the concurrent antiinflammatory response. These results are consistent with an immunomodulatory role for the RAD in patients with ATN and MOF. To evaluate the RAD’s influence further on local inflammation in tissue and distant organ dysfunction, especially in the lungs, a study compared bronchoalveolar lavage (BAL) fluid from cell-RADetreated and noncell, shamtreated groups in a pig model with septic shock with AKI [54]. The levels of total protein in BAL were significantly higher in sham control animals than in the RAD group (143  11 compared with 78  10 mg/mL, respectively; P > .05). Proinflammatory cytokines, including IL-6 and IL-8, were markedly elevated in the control group. These results demonstrate an important role for renal epithelial cells (RECs) in ameliorating multiorgan injury in sepsis by influencing microvascular injury and the local proinflammatory response. A more promising direction to improve the outcome of AKI is to better understand and interrupt the pathophysiologic processes that are activated in AKI, resulting in distant MOD and eventually death. AKI results in a profound inflammatory response state resulting in microvascular dysfunction in distant organs [55,56]. Leukocyte activation has a central role in these acute inflammatory states. Disruption of the activation process of circulating leukocytes may limit microvascular damage and MOD [57]. The RAD appears to influence systemic leukocyte activation and the balance of inflammatory cytokines and may alter the proinflammatory state of AKI and ultimately improve morbidity and mortality. Our group has developed a novel synthetic membrane embedded in an extracorporeal device to bind and inhibit circulating leukocytes. This selective cytopheretic device (SCD) is an acellular device that mimics many of the immunomodulatory aspects of the RAD. The SCD improved septic shock survival times in preclinical animal models [58] and improved the survival outcome of ICU patients with MOF in a prospective, single-arm, single-center study [59], a prospective, single-arm, multicenter US study [60], and a randomized, multicenter clinical trial in which SCDtherapy patients were maintained within recommended ionized calcium (riCa) range [61].

SELECTIVE CYTOPHERETIC DEVICE SCD therapy is an acellular device-based, extracorporeal blood treatment in which a patient’s blood is passed through a hemofilter followed by an SCD setup in series (Fig. 65.2B) while regional citrate anticoagulation is administered. The proprietary circuit [62] was designed to combine small solute clearance by the HF portion of the circuit, whereas the SCD portion provides immunomodulation via a loweshear stress (SS) environment for blood cells passing by the fibers within the SCD. SS encountered by blood cells within a hemofilter is greater than arterial SS in the body (>30 dyn/cm2), whereas SS encountered within the SCD approximates capillary SS ( 24 cm) substitution in a porcine model. BJU Int 2000;85:894e8. [133] Aurora A, Roe JL, Corona BT, Walters TJ. An acellular biologic scaffold does not regenerate appreciable de novo muscle tissue in rat models of volumetric muscle loss injury. Biomaterials 2015;67:393e407. [134] Kajbafzadeh AM, Khorramirouz R, Sabetkish S, Ataei Talebi M, Akbarzadeh A, Keihani S. In vivo regeneration of bladder muscular wall using decellularized colon matrix: an experimental study. Pediatr Surg Int 2016;32:615e22. [135] Roth CC, Mondalek FG, Kibar Y, Ashley RA, Bell CH, Califano JA, Madihally SV, Frimberger D, Lin HK, Kropp BP. Bladder regeneration in a canine model using hyaluronic acid-poly(lactic-co-glycolic-acid) nanoparticle modified porcine small intestinal submucosa. BJU Int 2011; 108:148e55. [136] Wishnow KI, Johnson DE, Grignon DJ, Cromeens DM, Ayala AG. Regeneration of the canine urinary bladder mucosa after complete surgical denudation. J Urol 1989;141:1476e9.

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73 Therapeutic Applications: Tissue Engineering of Skin Fiona M. Wood University of Western Australia, Perth, Australia

INTRODUCTION There has been an explosion of activity in the field of skin tissue engineering (TE) specifically focused on skin repair and regeneration. A simple online search of the key words “tissue engineering” and “skin,” 2007e17, reveals 51,842 publications, 41,780 of which are peer reviewed. Progress has been made in the areas of biomaterials for scaffold generation, the reliable source of cells, and novel approaches to assembly such as three-dimensional (3D) printing. However, skin is much more than a two-layered waterproof construct that provides microbiological protection. Comprehensive skin functioning involves maintaining homeostasis with intimate connectivity with the immune neuroendocrine systems. There is an intimate relationship with the underlying fat of the hypodermis and skin adnexal structures are functionally active [1]. Therefore, a fully integrated vascular and neural network is essential for cellular and neural signaling and driving cutaneous responses to changes in both the host and external environments [2]. The skin adnexal structures remain elusive to large-scale replication although our body of knowledge is building. Furthermore, the challenges of developing a vascularized construct integrating small vessels are such that the rapid or immediate restoration of skin function remains an ambitious goal [3]. Each of us is a self-organizing mass of multiple cell types. From fertilization of the embryo onward, our tissue structures develop until an adult morphology is achieved. At that point, our capacity for self-organization is directed to maintaining that morphology in the face of daily insults and the processes of aging. When a given insult overwhelms our capacity to repair by regeneration, the result is scar repair [4]. We know that tissues retain the variable ability to heal by regeneration [5]. With respect to the skin, in all but trivial injuries the capacity for regeneration is exceeded, triggering cellular mechanisms that result in scar formation. The mechanism of pathology or injury also has an impact on the outcome; for example, burn injury is notorious for developing aggressive scars that compromise the individual functionally, cosmetically, and psychologically [6]. It has been well-described that to heal a wound, we need a source of cells capable of differentiating into the given tissue type and an extracellular matrix (ECM) capable of supporting cell migration, proliferation, and differentiation [7]. We and others have spent considerable resources researching these areas to improve the speed and quality of wound healing to reduce the scar [8]. The question we ask is: “How can we harness the technology of TE of skin to provide a controlled repair to restore the original morphology?” TE was initially coined by Vacanti in 1990 and can be defined as the application of engineering principles to biological systems [9]. The healing of skin has been the subject of writings as far back as ancient times with the attempts to stimulate healing, protect the surface while healing, and even replace the skin surface [10]. Explorations into using TE principles in skin repair are based on a significant history. With the increasing knowledge and understanding of the skin structure and function along with the developing TE techniques, the question remains, can we provide a regenerative repair avoiding scar? We know that the skin provides the barrier to the external environment as a dynamic, complex, 3D structure made of cells from all embryological layers. Integral to its functions are the vessels and nerves within the tissue construct. Furthermore, skin is specialized over different body sites, adapting to local functional needs demonstrated by the

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macroscopic differences seen between areas, e.g., the eyelid compared with the palm of the hand. Notable body site differences are seen in the distribution and character of the hypodermis and adipose tissue with an influence on the function and repair potential. The varying capacity of skin to respond to injury is seen with the greater scar potential of sternal and deltoid areas [11]. Thus, as we go forward in developing TE solutions to skin, it is timely to take stock of the skin’s functions and interactions. How will skin’s changes over the different body sites influence TE needs? Collaboration with specialists in bioinformatics will be essential to understanding the implications of changes in genes, genetic expression, and epigenetics with the resulting phenotypic expression we need to guide the cells and matrix interactions into the tissue constructs [12]. In the therapeutic use of TE, an understanding of the etiology of the skin defect and pathophysiology of the patient will identify what needs to be replicated, rebuilt, and replaced. Implementation into clinical practice hinges on the TE being problem-driven, providing a practical, timely, costeffective solution to the clinical problem [13]. TE technologies have been used to produce skin constructs used to test a range of topical preparations and explore toxicology. However, from a clinical viewpoint, TE of skin has developed in response to the clinical need for, e.g., skin repair in major burns [14], chronic ulcers [15], and giant nevi [16]. It is clear that the current reference standard of skin loss, skin grafting, will always leave a scar [17]. The development of TE creates the opportunity to tailor the repair to the defect with the understanding that one solution will not fit all. TE offers innovative solutions providing a spectrum of clinical solutions including: • strategies to facilitate wound healing in situ by introducing bioactive agents ranging from biomaterials to cellbased therapies, and • the development of laboratory-based, multilayered tissue constructs including multiple cell types tailored to a specific skin defect. The development of TE of skin is intimately linked to the vision of facilitating scar-free healing. It has broad implications for after trauma or surgery and in fibrosing pathologies in which the common outcome is a functional compromise owing to the distortion of normal tissue architecture. In the developed world, it is estimated that 100 million patients acquire scars each year as a result of 55 million elective operations and 25 million operations after trauma. Within this number, it is estimated that there are approximately 11 million keloid scars and four million burn scars, 70% of which occur in children [18]. There is a clear opportunity to affect survival and quality of life by engaging in TE solutions. We are living in a time when science and technology are advancing at an exponential rate. Harnessing the power of that science and technology into clinical practice presents an ever-increasing challenge. We are all aware of the latest breakthroughs holding promise to improve the quality of life, such as genetic engineering facilitating DNA manipulation to improve health. However, the growth of knowledge is possible in controlled systems in which experiments can be designed to investigate a single variable; here lies the greatest challenge: for example, in burn injury research. The design of clinical trials is dogged by the complexities of assessing the extent of injury, the individual’s response, and the availability of validated outcome measures. Clinical practice is a fusion of experience and knowledge with the development of medical subspecialization directed toward a targeted problem-solving approach and has facilitated great advances in clinical care. However, this should not be at the expense of broad general knowledge gaining insight into potential links and facilitating cross-fertilization. It is essential to link the tissue engineer with the clinical specialist to ensure that the opportunities, risks, and benefits of TE skin are understood, to facilitate appropriate clinical trials. By collaboration among disciplines, there are real opportunities for improvements in clinical care translating to improved outcomes for patients [19]. The road from the bench to the bedside is long and navigates the areas of regulation, commercialization, reimbursement, and clinical trial design, to name a few. Also, translational research itself is an area in need of research and audit. The investigation of drivers and barriers to the implementation of TE skin solutions is an area of increasing effort. It is vital to learn from history and understand the current situation to continue developing innovative TE solutions but also innovative solutions within health systems, to move timely translation into clinical practice. To put the challenges and opportunities into context, this chapter will: • explore the functions of the skin and injury responses we need to understand to harness the TE technology effectively, • identify the needs that could be supplied by TE strategies,

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• discuss available technologies, • demonstrate TE skin solutions in clinical practice, and • look to future challenges in areas with a need for further development.

DEVELOPMENT, ANATOMY, AND FUNCTION OF SKIN Skin is commonly described as a multilayered physical barrier composed chiefly of the surface cellular epidermis and relatively acellular dermis. Exploring the development, anatomy, and functions of skin demonstrates its complexity and will guide our TE endeavors. There is no magician’s mantle to compare with the skin in its diverse roles of waterproof, overcoat, sunshade, suit of armour and refrigerator. Sensitive to the touch of a feather, to temperature and pain, withstanding the wear and tear of three score years and ten and executing its own running repairs. [20]

During development the initial covering of the embryo is the periderm, which is thought to have a barrier function expressing tight junctions; it interacts with the amniotic fluid, which has surface microvilli. Periderm cells express keratins, which are also associated with migratory epidermal cells seen in wound healing. During the first 6 weeks, the epidermis becomes two-layered, with the outer periderm and the inner developing epidermis. During this time there is no dermis, but a subepidermal cellular layer with deposition of basement membrane type IV collagen and laminin is seen by 5 weeks. By 8 weeks, there is evidence of vascular development. During the embryonic fetal transition stage at 9e10 weeks, several changes are seen: • epidermal cells express keratins 1 and 10, which are associated with differentiation; • maturation of the basement membrane occurs with the development of cell adhesion and expression of integrins a6 and b4; • rapid deposition of dermal matrix takes place; • migration of melanocytes begins from the neural crest; • Langerhans’ cells are detected originating in the fetal thymus and bone marrow; and • Merkel cells are seen initially in the epidermis and subsequently in the dermis. The early fetal period from 11 to 14 weeks sees the development of the hair follicles within the skin. The skin adnexal structures continue to develop and mature into the midfetal period at 15e20 weeks. As the fetus grows in the late period, 20e40 weeks, acceleration of stratum corneum can be identified in specific regions, including the palms, soles, face, and scalp. There is a close association between keratinization and hair follicle development, with the stratum corneum developing initially in the perifollicular regions. The mature stratum corneum is a structure that develops in the late fetal period: a combination of the terminally differentiated keratinocytes forming a cornified envelope and lipid extrusion from the abundant lamellar bodies of the granular layer keratinocytes. Embryologically, the neural tissue and the epidermis are derived from ectoderm. By the end of the fourth week of embryonic development, the neural ectoderm has separated from the surface ectoderm, forming the neural tube. Under normal conditions, the nerve endings will never be exposed to the external environment. The skin forms the interface and has developed as the interactive responsive surface. In the following weeks, mesoblastic cells from the neural crest migrate into the skin as melanoblasts and the early nerve fibers develop as the vasculature migrates into mesodermal elements destined to become dermis. A close association is seen in development with the skin, developing as a tactile organ providing feedback information from the surface to the developing central nervous system (CNS). An understanding of neural plasticity of the CNS and the peripheral nerve field underpins the self-organization principle. The co-development is put forward as an explanation of the coevolution of the human CNS and the skin as an adaptive dynamic interface. In the fully developed skin, there are cells from all three embryological layers in a complex framework of ECM. There are functions common to all skin areas, but there has also been adaptation to the specific functions of given body sites that are seen even at the early fetal stages. The mature epidermis is composed primarily of keratinocytes arising from a layer of basal cells situated on the basement membrane. As the keratinocytes differentiate, they form a stratified squamous epithelium. As the cells undergo terminal differentiation, they lose their nuclei and form a highly cross-linked protein-based layer of keratin. The basal cells are in intimate contact with terminal dendrites: synapse-like structures have been described between nerve endings and keratinocytes. The melanocytes are situated in the basal layer with the melanosomes being

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transferred to the differentiating keratinocytes giving the color of the skin owing to pigment load. The cells linking to the immune system, Langerhans’ giant cells and dendritic cells, are also present in the epidermis. The epidermis is specialized to the body sites, most noticeably with the thickened cornified layer of the sole and palm [21]. The dermis is attached to the epidermis at the dermal epidermal junction by the basement membrane morphologically arranged as the rete pegs, which are exaggerated in the glabrous skin areas. The dermis is mainly connective tissue, predominantly collagen, with elastin seen in the superficial papillary dermis. The fibroblast is the cell that produces the ECM, which is specialized over differing body sites; areas such as the groin and axilla are thin and more elastic than the thicker, more rigid dermis of the back. Cells of hematopoietic origin such as lymphocytes and macrophages migrate into the dermis as the vascular bed becomes established and are involved in immune surveillance. The neural and vascular networks maintain the skin and facilitate the functions of the dynamic interactive skin interface [22]. The investigation of wound healing, in parallel with an understanding of skin development and functions, gives us the opportunity to develop innovative methods further to apply to TE. The skin has developed specifically in relation to the multiple functions it performs. As an active organ, it is responsive to changes in the external and internal environment, pivotal in maintaining the body’s homeostasis. Our knowledge of skin functions is still growing and includes: • • • • • • • • • •

a semipermeable barrier, overcoat, and suit of armor; thermoregulation and refrigerator; antibacterial and waterproof; UV protection and sunshade; sensory receptor that is sensitive to the touch of a feather, temperature, and pain; self-regenerating, withstanding the wear and tear of a lifetime; capable of rapid repair, executing its own running repairs; immune modulation; psychological interaction; and vitamin production.

The loss of skin integrity can result in severe morbidity and even mortality. The body needs a barrier against the atmosphere to maintain homeostasis. The production of the stratum corneum can be mimicked in tissue culture by exposing a sheet of keratinocytes in culture to an aireliquid interface [23], but it requires maturation in situ to develop the “smart material” of the enucleated cell bodies fully and the extracellular lipoproteins moisturized by vitamin Eeproducing sebum. The keratinocytes produce surface proteins that are antimicrobial in the first line of defense against colonizing bacteria on the skin surface [24]. The expression of these proteins changes as the keratinocyte is stressed, as in wounding or culturing, such that protection from microbial invasion is highlighted. The multiple and specific sensory inputs to the skin are pivotal in regulating the body’s temperature and immune responses and psychological responses via neural and neuroendocrine control systems. Animal studies have highlighted the sensory role of the skin in normal development; touch is associated with growth potential of the internal solid organs [21]. The skin is also profoundly influenced by the pathophysiology of the individual, with cutaneous changes in anatomy and physiology being linked to many disease processes. With this expanding body of knowledge with respect to the skin, we engage the fields of TE to provide solutions to repair and replace skin defects to maintain function and avoid scarring. To do so, we also need a working knowledge of the skin response to injury and tissue loss: the processes of wound healing [25]. Briefly, wound healing in skin is a complex series of cascading events that has been described in three overlapping stages from the initial inflammation to tissue formation and subsequent tissue remodeling [26]. The initial response is clot formation to achieve hemostasis. The activation of platelets releases the contents of their a-granules, resulting in activation of the clotting cascade and the release of adhesive proteins forming the matrix of the clot, e.g., fibrin and chemotactic factors and growth factors into the wounded area. The coagulation pathway links to activation of the complement pathway’s facilitation of the recruitment of neutrophils needed to facilitate the inflammatory response by removing cellular debris and microorganisms. The ingrowth of new vessels as granulation tissue is initiated and the keratinocytes at the wound edge mobilize to commence reepithelialization. Macrophages migrate to the wound, releasing multiple protein growth factors as the wound response progresses to the repair phase. Both hemopoietic and mesenchymal stem cells (MSCs) from the

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circulation are attracted into the wound with fibroblasts producing ECM, some as myofibroblasts associated with wound contraction. There is an increased interest in the interactions between the keratinocytes and underlying fibroblasts as the matrix is remodeled and the new basement membrane develops. As the keratinocyte is exposed to collagen, it secretes collagenase and as the basement membrane integrity is restored, the cells revert to their normal phenotype in the situation in which healing is achieved without scarring. Molecular cues after inflammation are associated with positional information and assist in driving the celleECM interactions. In situations of more extensive tissue damage, the fibroblast transforms to a scar phenotype with the production of disordered collagen, resulting in an ECM framework that differs chemically, architecturally, and mechanically from the native skin construct. The interactions between the ECM and the fibroblasts respond to the changes as healing progresses from tissue repair to remodeling. The initial migratory fibroblast transitions to the profibrotic phenotype, producing ECM proteins. In the remodeling phase, the cell number within the dermis or scar reduces as the cells undergo apoptosis. The construct is dynamic and responsive to cellular, chemical, and mechanical cues. Knowledge of the wound-healing progression over time allows the TE skin solution to be clinically integrated into the process and may be directed toward a number of target strategies: the control of cells in growth, the genetic manipulation of cells to express a given phenotype, seeding of the retained dermis with cells from the dermal epidermal junction, removal of the full thickness of the area of compromised skin, and replacement with a specifically tailored TE construct.

POTENTIAL PREREQUISITE REQUIREMENTS FOR TISSUE ENGINEERED SKIN SOLUTIONS The skin matures from the softness of the newborn to the skin in old age with the loss of elasticity and reduced potential for repair [27]. We believe that the young heal rapidly but scar aggressively, in contrast to the elderly, who heal more slowly and scar less [28]. Regeneration of the skin without functional or aesthetic deficit, rather than enhanced repair, remains the ultimate goal of wound-healing therapies [29]. However, the degree of scarring and the quality of the repair depend highly on the time taken to heal, with faster healing correlating to improved outcome [8]. The availability of the TE skin for timely use is a key factor balancing the issues of allograft to autograft and biological to nonbiological solutions. The differences in wound-healing responses with age, the condition of the patient, and the etiology of the defect will have implications when harvesting donor tissue for TE and will potentially influence the choice of TE technique [30]. The skin surface is continually replaced under normal conditions and the morphology is retained over the years with changes seen as a result of injury, pathology, and aging. The capacity to regenerate and self-organize becomes overwhelmed in all but trivial injuries such that the repair forms a scar that all too often is debilitating, both physically and psychologically. The traditional approach to reducing the time to healing a skin defect has been to graft skin. Full-thickness skin grafts (FTSG) will give the best scar result but appropriate donor sites to match with recipient sites can lead to skin mismatch with retention of the donor site characteristics, as seen when an FTSG from the groin is used to release a contracture on the palm of the hand. The donor site availability for split-thickness skin grafting (SSG) may be limited in size when large body surface areas are compromised as in burn injury or giant nevi. The area of cover of a given donor site can be increased by meshing or expanding, as in the Meek technique [31]. The expansion of the SSG is associated with small areas of wound healing by secondary intention and a poorer scar outcome. It is the desire to eliminate or at least reduce the scar by reducing donor site morbidity that has driven TE of skin over the past several decades. Essential factors to achieving healing are: • a source of cell’s capable of differentiating into the tissue and • an ECM capable of supporting the cells. With an increased understanding of the skin physiology and interactions with the internal and external environment, we need to also consider: • the 3D spacial information of the area under repair and • feedback from the surface to facilitate self-organization. The ideal needs for a TE skin replacement continue to be debated and are related to the clinical indications for use. Variations in skin with body site and age dictate that a tailored solution is required. Furthermore, the mechanism of

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skin loss has a profound impact on the local and systemic response; when comparing surgical excision with a burn injury, it is seen that the burn is associated with an extreme inflammatory response [32]. The interventions and timing of the interventions are important considerations when preparing the wound for the TE construct. In our group, in treating acute burn injury and scarring, we have been guided by the following basic requirements [33]: • • • • • • • •

rapidly available autologous site-matched reliable wound adherence minimal donor site morbidity clinically manageable improved quality of scar affordable

The list could be debated and expanded according to clinical needs and opportunities such as the use of allograft MSCs with their demonstrated immune privilege [34]. However, the improvement of outcome by clinically manageable affordable solutions affords a solid benchmark. For TE to be successful it is clear that an in-depth working knowledge of the biology of the tissue is essential. There are a number of cell types within the skin, each with specific and often interrelated functions, some established within the construct, and others transitory from the circulation; their origin and maintenance of clonal capacity remain elusive. It is fundamental to the success of TE to have an understanding of essential information about how the cells relate to the other cells and the ECM of the skin to develop into a mature tissue construct [35]. The ECM scaffold is integral to tissue integrity. We know that the physical shape, mechanical properties, and chemical composition of the environment of a cell influence its phenotype. The predominate dermal cell responsible for ECM production, the fibroblast, changes its phenotype after injury and repair, as demonstrated in the epigenetic modification. The stiffness of the ECM is linked to the fibroblast phenotype, and an understanding of the mechanobiology is an opportunity to guide appropriate tissue regeneration. With the complexity of multiple factors at play, hierarchical mathematics will be useful such that the field of bioinformatics may assist in refining design in the future [36]. In designing of the TE solution, an understanding of this relationship will lead to increased clinical success guiding phenotypic expression. From an engineering perspective, there are technologies that will facilitate innovative clinical solutions such as the advanced modeling and fabrication for scaffold manufacture. Bioprinting has demonstrated the capacity to use the advancing knowledge of biomaterials, natural and synthetic, combined with the cells to produce scaffold cell constructs of increasing complexity. The materials for use need to [37]: • • • • • •

be biocompatible; have properties compatible with mechanical aspects of bioprinting, such as extrusion; facilitate appropriate tissue-specific cell viability and function; maintain structural integrity and mechanical properties after printing; be responsive to chemical and cellular signaling; and integrate into the host tissue environment with vascular and neural connectivity.

The development of bioreactors to maintain viability and expand cell numbers associated with scaling up in an appropriately regulated laboratory is essential for cell cultureebased techniques [23]. With the advent of bioprinting, it is essential to consider cell capabilities that ideally [37]: • have the capacity for cell proliferation generating adequate cell numbers in a timely fashion, depending on the pathology being treated; • are nonimmunogenic or autologous; • are multipotential; • are responsive to the environment; and • have a capacity to integrate without fibrosis. The increased use of MSCs as immune-privileged cells in a range of pathologies has led to the exploration of them as a source of cells in the TE field meeting these criteria [34]. The increasingly reliable generation of patient-specific induced pluripotential stem cells (iPSCs) provides an alternative source of cells with the capacity to be plastic down a range of lineages, affording potential for future differentiation into specialized cells of the skin adnexal structures [38]. As a source of cells for genetic engineering, treatment of cutaneous pathology becomes a possibility. Cells from

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a range of sources can be guided by the introduction of signaling molecules to facilitate activity such as migration, proliferation, and differentiation, bringing us back to having an in-depth working knowledge and even expanding that knowledge base from embryology to wound healing to introduce the appropriate triggers. The cells can also be manipulated by physical changes such as hypoxia to stimulate responses or be guided by introducing magnetic nanoparticles to allow manipulation by external forces such as a magnetic field [39]. Putting the ECM and cells into proximity of a 3D structure is the beginning of the process that needs to be guided either in the laboratory or in situ in the wound to develop into the tissue construct. It is clear from the experiences of the past several decades that many of the proposed TE skin solutions are disruptive technologies. It is essential that the TE skin not be designed for only a clinical problem and that it should be reproducible and reliable, but also that it should be linked to an education and training program so that its full clinical potential is realized [19].

CURRENT TISSUE ENGINEERING SKIN TECHNOLOGIES Engineering principles have been applied to skin for many years with the development of medical devices to harvest SSG accurately and mesh the skin to allow expansion. The practice of tissue expansion is a well-established surgical tool for the development of skin by subcutaneous insertion of inflatable devices in vivo, which can be serially enlarged with the resulting development of the skin as demonstrated by cell proliferation [40]. The tissue-expanded skin has all layers and the complete characteristics of the donor site, including retention of adnexal structures and functional innervation [41]. There is increasing interest in the concept of tissue expansion in vitro with a full-thickness skin biopsy put under tension in a bioreactor system to maintain its viability and facilitate cell replication resulting in tissue growth [42]. However, these solutions are limited by time, area, and in some cases, donor site and scar outcome. The need remains to provide both rapid large surface area cover and complex site-specific specialized skin repair, with the ultimate goal of healing by regeneration, not scar repair. The initial approach to TE skin was to separate the layers and consider the epidermis and dermis to be separate problems. Work by Green in 1970 focused on the culture of keratinocytes into cell sheets suitable for grafting [43]. The solution to large surface area skin repair was to harvest cells from an uninjured donor site and to undertake laboratorybased tissue expansion. The resulting sheets of cells could be used to close the wound as would a traditional SSG. However, there were problems with the time taken to culture in the laboratory, fragility, adherence to the wound surface, and durability over time, because only the epidermis was replaced; in addition, cost was problematic [44]. In trying to solve some of these problems, there has been a development in the area of subconfluent cell transfer on a number of cell culture surfaces [45] in addition to the delivery of cells in suspension as an aerosol [46]. Subconfluent cells have more reliable adherence and are available in a shorter time frame, from 3 weeks for sheets to 5 days for subconfluent cultures [47]. The process of harvesting cells from the dermal epidermal junction by enzymatic and physical dissociation has been used for immediate delivery of a noncultured cell population to the wound [48]. The cells are a mixed cell population in the same ratio as that seen in the normal skin construct, because there has been no selection of cell populations seen when culturing. The maintenance of the melanocytes enables the development of appropriate pigmentation [49]. The cells adhere, migrate, and proliferate across the wound surface and then differentiate and self-organize into a mature epidermis. The scar outcome is intimately linked to the underlying wound bed, which will be discussed in a later section. The development of a suitable dermal scaffold is also the focus of the TE field [50]. Topographical features are known to influence cell behavior through a phenomenon known as “contact guidance,” and alteration in the size of the surface detail can elicit different cell responses [51]. Running parallel to studies on the epidermal cell culture was work by Yannas and Burke on dermal replacement, which culminated in the first commercially available product, Integra [52]. The concept of tissue-guided regeneration within an architectural framework is in clinical practice [53]. Underpinning research on the composition and construction demonstrated the importance of considering both aspects: a combination of bovine collagen coated with glycosaminoglycan but with a pore size of less than 60 mm or greater than 100 mm, resulting in disordered granulation tissue, with the optimal pore size resulting in the migrating cells expressing a reticular dermal fibroblast phenotype. The main drawback with Integra is that it addresses only the dermal aspect, with the outer layer on silicone acting as a pseud-epidermis for the period of vascularization, usually 3 weeks before a second surgical procedure is needed to repair the epidermis [54]. The epidermal repair is with a thin SSG, which may be meshed to cover a larger area than the donor site with epidermal cells to speed the time to

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healing and reduce the mesh scar pattern [55]. The two-stage problem has been addressed by trying to reduce the time to vascularization or by developing constructs that can be used with an SSG at the same procedure. Apligraft, Matriderm, and Pelnec are dermal templates marketed to provide appropriate topography and matrix properties to promote cell migration into the wound, improve healing, and reduce scarring as a one-stage procedure with SSG [56]. An alternative is to use our knowledge of healing as we see cells migrating from areas of the skin adnexal structures in the dermis to form the new epidermal layer. Introduced cells harvested from the dermal epidermal junction seeded into Integra will migrate and organize into a new epidermis with an established dermal epidermal junction within 3 weeks [57]. The use of Integra is a clear demonstration of the potential of tissue-guided regeneration with the expression of a cell phenotype guided by the morphology and chemistry of the matrix [58]. It is well-established that cells change their phenotype in response to changes in their environment [59]. The development of suitable technologies to generate an optimal environment for wound healing is important to enhancing cell response to tissue injury, reducing the time to heal and improving the outcome. Our knowledge of ECMecell interactions is increasing with the recognition of the cell signaling by nanoscale structures on cell surfaces with roles in attachment and cell migration [60]. Developments in nanotechnology have opened up possibilities in TE to improve scaffold design, but relatively little is known about how changes in topography at the nanoscale affect cell behavior [61]. The scaffolds can be manufactured to address specific skin functions: to protect against injury from loss of fluid and proteins, enable the removal of exudates, inhibit exogenous microorganism invasion, and improve the aesthetic appearance of the wound site [62]. Scaffolds are generally matrices of synthetic and/or natural polymers fabricated by various techniques including solvent-casting, gas foaming, electrospinning, phase separation, freeze-drying, melt molding, and solid free-form fabrication [63]. Critical to their performance is reproducibility, with control over pore size and the distribution of pores, removal of residual toxic organic solvents, and the control of the inflammatory and immune responses owing to polymer degradation and the associated byproducts [64]. The printing technology could conceivably be used directly in wound bed preparation, providing initial protection and a blueprint for cell migration and guided tissue regeneration [65]. Our group has explored anodic aluminium oxide (AAO) as a potential scaffold or template in TE [66]. Selforganized oxide growth under controlled conditions generates a densely packed hexagonal array of uniform-size nanopores aligned perpendicular to the surface of the AAO film. The size of the pores can be nanoengineered by manipulating the anodization time and voltage, the anodizing electrolyte, and/or the time of postchemical etching. Aluminum oxide is well-known for its biocompatibility in the human body; it is inert, stable, and nonreactive, which makes it suitable for TE applications. Engineering of surfaces to manipulate healing is a rapidly expanding area; interactive dressing systems is in widespread clinical use. With the realization of the impact of surface topography and chemistry on cell expression and developing nanoengineering techniques such as electrospinning and electrospraying, there is increasing interest in smart surface technology skin healing [67]. Biocompatible polymeric selfassembling nanofiber constructs have the advantage of a large surface area that can be linked to bioactive compounds. The release of the bioactive compounds can be controlled by intrinsic factors such as in a hydrogel, release kinetics, or extrinsic release triggers [67]. Exciting advances have been made in the area of nanocubes, nanocages, and nanorods as primary candidates to be studied for the phototherapeutic release of bioactive agents [68]. Clinically, the results of single cytokine applications have been disappointing and study of natural healing processes is a result of complex cascade interactions over time [68]. It is unsurprising that we cannot achieve with a single cytokine administration what is the result of complex celleECM interactions. The aim of advancing technology is to mimic the structure and function of the ECM with the ability to adapt over time to the changing environment of the healing wound [69]. TE technologies are also used to test the potential impact of agents on the skin, such as in the cosmetic industry. The degree of complexity of laboratory-based testing systems vary from 3D-printed constructs at an aireliquid interface to produce a differentiated epidermal layer to techniques such as the “scar in a jar” system for testing biological activity of drugs targeting specific pathways in fibrosis [70]. Bringing together the scaffold and cellular components has been successful with the development of multilayered constructs seeded with multiple cell types [71]. Clinical series have been presented demonstrating a soft supple skin but with the persistent problem of poor color match [72]. An understanding of the interdependence of the cells of the dermal epidermal junction may well be important to the development of skin constructs with the appropriate melanocyte function [73]. The main drawback in the clinical use of complex laboratory-based constructs is the time taken in the laboratory [74]. However, the potential to use such a technology in timed reconstructive surgery as opposed to acute trauma is beneficial with the ability to tailor the skin construct to the planned defect. The 3D distribution of cells within the wound has been addressed by the innovative use of “inkjet printing” technology, with the cells “printed,” controlled by the shape and depth of the defect [75]. The cell type within the system can be

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changed with the fibroblasts being laid down before the keratinocytes [76]. ECM and other proteins can also be introduced into the system, such as hair-based keratins, chemical-processing substrates from biological origins to develop innovative scaffold solutions to enhance cell performance within the constructs. The source of the cells and the guidance along a path of differentiation are critical elements for a functional skin construct. The use of iPSCs offers the potential for differentiation into multiple cell types, including fibroblasts, keratinocytes, and melanocytes [77]. Guidance of cell differentiation and ECM production may hold the answer to replicating the adnexal structures. The iPSCs also have potential in gene-based therapies in treating skin disorders such as epidermolysis bullosa [78]. The use of the cells as therapeutic delivery agents may broaden therapeutic opportunities. The time to availability of a TE technique is a key driver to their clinical employment. The use of allograft materials allows an “off-the-shelf” approach, whereas autograft material may take time in the expansion phase [79]. The use of scaffolds alone can provide an advanced wound management system to facilitate healing and can replace a tissue defect guiding repair as it is replaced or provide a permanent solution. The use of cells alone can also provide a surface epithelium that can modulate the underlying healing, form a mosaic of cells of intrinsic and extrinsic origin guiding repair, or provide a permanent surface [80]. The combination of the two elements can provide an advanced skin repair solution, but because the differentiation of the construct is more advanced, the time taken in the processes is prolonged. Consideration of the time taken has led to an investigation of the construct being used in the immature form differentiating in situ. It is clear that TE provides a range of innovative solutions that are useful for many wound and skin replacement areas. Effective use of the technology hinges on the clinician understanding the wound preparation and the aim of the repair. The wound assessment drives the initial clinical decisions directing management in terms of resuscitation, tissue salvage, and infection control, and then planning the repair. An understanding of the range of TE solutions is essential to appropriate clinical implementation. Consideration of the hypodermis in the resulting functional unit adds a further degree of complexity and opportunity to advance the quality of the tissue construct [81]. Fat grafting to improve the contour of the hypodermis has been shown to be associated with modulation of the cutaneous scarring. The cellular mechanism of the interaction is of interest when considering the opportunity for systemic, regional, or local solutions for a given clinical problem.

TISSUE ENGINEERING SKIN SOLUTIONS IN CLINICAL PRACTICE The following clinical case of an extensive 65% total body surface area burn injury associated with multiple fractures and pneumothorax is used to demonstrate the decision making and options available in current practice. The initial stabilization and resuscitation is life-saving, along with attention to infection control using Acticoat, a nanocrystalline silver dressing. Once stable, surgical debridement is planned to excise the areas of skin that cannot be salvaged. Although the timing of debridement is vigorously debated with respect to controlling ongoing inflammation and improving outcome potential, an important element in decision making is what is available to cover the debrided wounds, either as a temporary measure or as a permanent replacement. A full-thickness wound requires dermal and epidermal repair for the optimal outcome. In largeesurface area wounds, standard SSG is not possible in one procedure owing to restricted donor sites; the SSG can be meshed to achieve healing in a larger area but healing of the interstices by secondary intention often leaves an unsightly meshed pattern scar. Cadaver allograft has been widely used but provides only a temporary solution and will require serial cover as the donor sites heal. A combination of allograft dermis and sheet cultured epithelial autograft (CEA) has been reported to result in a composite repair with retention of the dermal elements. The CEA sheets may take 21 days to culture, or more timely availability can be achieved using preconfluent cultures on carriers or delivered as suspension. The dermis can be replaced with a number of off-the-shelf products such as Matriderm, Pelnec, or Apligraft. The TE technique most widely reported to date remains Integra. The use of composite TE skin could be considered to augment the second stage of the repair with Integra, because it takes time in laboratory preparation. In this case, Integra was used to reconstruct the dermis in the areas of full-thickness excision, as seen in Fig. 73.1, in which the Integra is held in place with a combination of staples and dressings to facilitate vascularization and integration of the construct. In areas where dermis could be salvaged, cells were harvested from the dermal epidermal junction of uninjured skin using a ReCell kit with Biobrane as a dressing. The main drawback with Integra is the period of vascularization of 3 weeks before proceeding to repair of the epidermis. The epidermis was repaired using a thin meshed SSG 1e3 combined with a noncultured cell suspension from the ReCell kit to reduce the meshed pattern scar. The healing is seen in Fig. 73.2 with the mash pattern fading well in the upper compared with the lower

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FIGURE 73.1 Postsurgical debridement of full-thickness burn over 65% of the total body surface area (TBSA) and replacement with Integra dermal scaffold held in place with elastic netting; area of a partial-thickness burn on the abdomen treated with autologous cells under a protective dressing; and uninjured skin donor site on the left flank.

FIGURE 73.2 Post-second procedure to repair the epidermal layer using a combination of meshed split-thickness graft with a noncultured autologous cell spray.

abdomen. At the time of scar maturity in Fig. 73.3, the scar situation demonstrates the issues faced by the patients [82]: • • • • •

contour deflect caused by the removed subcutaneous fat layer, persistent mesh pattern in the lower scar, mismatch of pigmentation, contracture bands distorting the anatomy, and the repair resulting in a scar.

There has clearly been progress; the development of TE techniques is intimately linked to the advancing survival and quality of scar in patients with major burn injuries. However, there is a clear need to continue to develop TE with the aim of total 3D soft tissue and skin replacement [83].

THE FUTURE The vision of scarless healing has led to the exploration of regeneration and the interplay among genes, cells, and tissues [84]. Pluripotential stem cells are present within each individual; the drivers of the cells down a regenerative path are an as yet unknown but exciting area of research with promise for the future [85]. The introduction of

THE FUTURE

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FIGURE 73.3 Two years after injury, before planned reconstructive surgery.

allograft stem cells such as MSCs may provide an alternative source of cell regeneration [86]. Understanding that every intervention from the time of injury influences the scar worn for life has pushed research in a multitude of directions. TE of skin is an exciting area that already has made a significant clinical impact [87]. However, we are far from the routine provision of technologies and strategies to provide site-matched, fully functional skin [88]. Great progress has been made in the areas of dermal templates and cell-based therapies and bringing the two elements together in skin constructs. The clinical implementation of TE skin solutions, along with the use of interactive surface dressing systems, has improved outcomes. The task is far from completed: • Skin adnexal structures are elusive [89]. • Timely availability of TE skin remains problematic [3]. In addition, more work is needed in: • • • • • • •

harnessing the explosion in smart material technology [90]; developing vascular constructs, the small vessel problem [91]; understanding the drivers of tissue-guided regeneration [92]; understanding the concept of self-organization and the bioinformatics behind morphology [93]; investigating the impact of both reinnervation and neural plasticity and its role in scarless healing [94]; understanding the barriers to clinical translation [95]; and developing regulatory pathways for novel solutions to ensure safe but timely availability [96].

At the high-tech end of the spectrum, we may consider bringing together laser surface imaging linked to fabrication to build bioreactors with the shape of the defect, with the “smart” cytokine-loaded scaffold materials tailored to the correct 3D shape. Cells of the appropriate body site could then be introduced into the scaffold by cell-printing methods in the individualized bioreactor and the flow of tissue culture medium used to induce small-vessel formation. At the time of transplant, the application of external techniques such as infrared could control the release of biologically active molecules from the “smart” scaffold surface to ensure, for instance, reinnervation and restoration of function with the capacity to integrate into the body. Understanding of linking local tissue replacement with systemic integration is essential to facilitate the linking partially cellular constructs with a 3D framework architecture facilitating secondary cell migration and ingrowth. An alternative to the individualized bioreactor could be the wound itself, prepared by “smart” surfaces and directly seeded with cells. With the understanding of the drivers to self-organization, they could be used to enhance in situ tissue-guided regeneration.

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CONCLUSION Over the past several decades, the concept of TE has been an area of active research, investigating innovative solutions. Multiple combinations of 3D engineered scaffolds exist with a functional cell load to produce tissue over time, which is the fourth dimension of skin repair fundamental to the clinical selection of technique. The elements required for tissue repair are: • a source of cells capable of differentiation into the lost tissue, • an architectural framework for cells to migrate into and express the appropriate phenotype, • 3D spatial information of the damaged tissue and the relationship to the surrounding viable uninjured tissue interface, and • a feedback mechanism to guide self-organization. In Jul. 2005, the Medical Journal of Australia published a vision of clinical care in a number of disciplines in 50 years’ time: Assessment is key in understanding the extent of injury. Debridement is focused on tissue salvage. Reconstruction balances repair with regeneration. Investigation of multimodality, multiscale characterisation, including confocal microscopy and synchrotron technology will quantify assessment. Debridement using autolytic inflammatory control techniques with image guided physical methods will ensure the vital tissue frameworks are retained. Tissue guided regeneration afforded by self-assembly nano-particles will provide the framework to guide cells to express the appropriate phenotype in reconstruction. To solve the clinical problem a multi-disciplinary scientific approach is needed to ensure the quality of the scar is worth the pain of survival.

Many of the technologies highlighted are available, but the significant need for clinical translation remains to move along the innovation pathway to ensure safe implementation into health care systems. Progress requires collaboration at all stages from basic science to clinical trial design and health economics, driven by improved clinical outcomes. Translation of new technologies into health systems requires the rigor of a research framework to identify and measure the impact of innovation in communication and education. Close working relationships between basic research and clinical service delivery are essential. Furthermore, the scientific and clinical advances need to be in line with regulation programs and linked to commercial interest to ensure their widespread availability. If the solution to the problem is scarless healing by a regenerative repair process and the aim is to improve the outcome from injury by restoring function, the future must blend the long-term vision with incremental shortterm improvements. There has been great progress in TE of skin, and it is an exciting area that offers tangible clinical solutions with an enormous potential for further improvement. The range of potential solutions that have developed is clear recognition of the complexity of the problem and the unique requirements relating to body site, patient, and pathology and the extent of the skin needing to be replaced, repaired, or regenerated. The challenge we face is to capitalize on that tradition and link to the opportunities afforded by unprecedented growth in science and technology, to ensure the quality of the scar outcome is worth the pain of survival.

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74 Regenerative Medicine Approaches for Engineering a Human Hair Follicle Gail K. Naughton Histogen, Inc., San Diego, CA, United States

INTRODUCTION The hair shaft consists of terminally differentiated keratinocytes, which are tightly compacted to form the hair shaft. Hair shafts are made by the hair follicle, a complex miniorgan that is formed primarily during fetal and perinatal skin development [1e3]. Hair follicle formation involves tightly coordinated interactions between the ectoderm and mesoderm with ectodermal stem cells giving rise to all epithelial components of mesoderm cells forming the follicular dermal papilla, as well as the connective tissue sheath [4]. It is composed of an outer root sheath (ORS), inner root sheath, and hair shaft. Under normal homeostasis the hair follicle undergoes a cycling, which is characterized by the formation of a new hair and shedding of an old hair through the growth cycle (anagen), apoptotic regression (catagen), and a resting phase (telogen) (Fig. 74.1). A very unique characteristic of the mammalian hair follicle is its lifelong recapitulation of its molecular embryogenesis each time it enters into anagen. Dermal papilla cells (DPs) are located at the base of the follicle and are surrounded by epithelial matrix cells. DPs support the proliferation as well as the differentiation of epithelial matrix cells and activate bulge stem cells during the transition from telogen to anagen. The bulge is part of the ORS that is close to the sebaceous gland and the interfollicular epidermis. Bulge cells give rise to all of the epithelial cell types during anagen, as well as the sebaceous gland and overlying epidermis. In anagen, cells at the base of the hair follicle proliferate and migrate up the follicle to form a new hair. Simultaneously, the bulge cells give rise to progeny. Robust activity during anagen requires an increase in nutrients and the phase is therefore also associated with new blood vessel formation to facilitate a more rapid supply of these nutrients. Much has been learned about the regeneration of skin and skin appendages through wound-healing studies and cell- and tissue-based treatments. Some of the earliest approved products in regenerative medicine were cell-based three-dimensional engineered tissues for the treatment of acute and chronic wounds [5]. These products taught us the importance of naturally secreted growth factors and matrix proteins in inducing healthy granulation tissue, angiogenesis, and reepithelialization. Unfortunately, none of the products induced the formation of critical skin appendages, including hair follicles and sweat glands. Anecdotal evidence in wound care reported cases with badly debrided skin healed with the formation of new hairs. [5a] It had been previously thought that mammalian hair follicles only form during fetal development [6] and that the loss of hair after birth was permanent. Reports that hair follicles could be developed de novo after wounding were published decades ago in mice [7], rabbits [8], and humans [9] but these were not accepted due to the lack of evidence of follicular neogenesis. In 2007 a pivotal paper was published showing that hair follicles formed de novo after wounding in genetically normal adult mice [10]. Evidence was presented that nascent follicles arose from epithelial cells outside of the stem cell niche and suggested that wound epidermal cells can assume a hair follicle stem cell phenotype. Overexpression of the Wnt ligand in the epidermis increased the number of new follicles, while inhibition of Wnt signaling completely stopped the folliculogenesis. The regenerated follicles produced new hairs and cycled up to three times during the 90 days following wounding, supporting the presence of functional stem cells in the newly formed follicle. The de novo formation of hair follicles in the adult mammal mimics embryogenesis at the molecular level and raises the possibility

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FIGURE 74.1 The normal hair follicle cycle.

of forming new follicles from preexisting hair follicle stem cells in the scalp or from controlled wounding when used in combination with Wnts and other cell stimulatory molecules. A number of preclinical and clinical studies have been performed using cells and cell- and tissue-derived materials to induce stem cell proliferation in the dying follicle or create new follicles in the scalp. The onset of hair loss is characterized by a much shorter than normal anagen (growth phase) and a reduction in hair density [4]. Given this mechanism, the development of a successful hair loss treatment needs to address two key processes: the transformation of vellus hairs into terminal hairs by stimulating and extending the length of the anagen phase, and the induction and maintenance of the anagen across the entire area effected by miniaturized follicles that result in increased hair density. The number of patients with alopecia has been on the rise and has been attributable to hormones and genetic factors [11], as well as autoimmune diseases, chemical and heat stress, medicines, and psychological stresses [12,13]. Androgenetic alopecia (AGA) is the most common form of hair loss in men during which terminal hairs turn into vellus hairs under the influence of dihydrotestosterone (DHT). Hair follicular miniaturization occurs as a consequence of progressively shorter anagen phases. To date the treatments of alopecia have focused on either blocking the DHT or inducing greater blood flow to help support anagen. These treatments need to be used daily, and within 2 weeks of stopping treatment there is a mass shedding of hair. What is needed is a treatment that will reactivate the miniaturized follicle and cause normal and sustained anagen. Research suggests that dormant hair follicles in human scalp affected by alopecia have a largely intact bulge stem cell population and a significantly reduced population of hair germ progenitors [14]. The hair germ progenitors become activated at the very onset of anagen, while the bulge stem cells become activated shortly after the actual initiation of anagen [15]. It is therefore conceivable that the reactivation of anagen in alopecia-affected telogen (resting) hair follicles may occur by the delivery of threshold levels of growth factors to induce bulge cells as opposed to triggering hair germ cells. The use of such a growth factor therapy has been studied through a variety of approaches, including autologous platelet-derived growth factors (PDGFs), growth factors from adipose cells grown under hypoxic conditions, growth factors from hypoxia-induced multipotent stem cells (HIMSCs), and the use of autologous stem cells to act as growth factor delivery units.

USE OF AUTOLOGOUS GROWTH FACTORS IN HAIR FOLLICLE REGENERATION Autologous growth factors have been associated with the promotion of tissue regeneration in a number of medical applications. Early work with platelet-rich plasma (PRP) showed excellent generation of granulation tissue and subsequent reepithelialization of hard-to-heal diabetic ulcers. Autologous PRP is derived from centrifuging whole blood to collect a pellet of platelets. Alpha granules in activated platelets release numerous growth factors, including epidermal growth factor (EGF), basic fibroblast growth factor (bFGF), platelet-derived growth factor (PDGF-BB), transforming growth factor beta 1 (TGF-B1), and vascular endothelial growth factor (VEGF). Although initial studies

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with PRP focused on the treatment of diabetic ulcers, PRP is routinely injected into the skin for wrinkle indications, it is injected into joints to reduce pain and inflammation, and some findings suggest that PRP may also significantly support hair follicle restoration. One such study involved the treatment of 19AGA patients with plasma rich in growth factors (PRGFs) [16]. Five PRGF injections were administered into the mid-dermal region and phototrichograms were performed at baseline and at 1 year following the last set of injections to assess follicle density/diameter and the terminal/vellus hair ratio. After PRGF therapy, mean hair density/diameter increased as did the terminal/ vellus hair ratio. An improvement was also noted in epidermal thickness, perifollicular neoangiogenesis, terminal/ miniaturized hair ratio, and cell proliferation. Although this was an uncontrolled clinical trial, the data do support a potential therapeutic effect of the autologous growth factors on hair follicle regeneration. Currently, a number of practitioners are looking at factors that can affect the ability of PRGFs to induce terminal hairs, including the age of the patient, vitamin D deficiencies, and general health.

USE OF ADIPOSE-DERIVED STEM CELLS AND THEIR CONDITIONED MEDIUM FOR HAIR GROWTH Adipose-derived stem cells (ADSCs) are mesenchymal stem cells (MSCs) found within the stromalevascular fraction of subcutaneous adipose tissue [17]. ADSCs self-renew, display a multilineage developmental plasticity, and have been used in a variety of tissue repair and regeneration clinical studies. It has been reported that ADSC transplantation promotes hair growth in animal experiments [18]. This is not surprising since hair regeneration depends on cellecell interactions, as well as external signals surrounding the follicle. Communication between adipose tissue and the hair follicles is key and adipose cells play an important role in the progression of the hair cycle [19]. Canine ADSCs were isolated and cell differentiation into dermal papilla-like tissues (DPLTs) was induced by culturing them in a dermal papilla-forming medium. The trichogenicity (hair forming potential) of the differentiated cells (DPLTs) was assessed and athymic mice were treated either with control or the differentiated cells. DPLTs were seen to have a compact aggregated structure and they secreted a complex extracellular matrix component, as well as versican and alkaline phosphatase, two dermal papilla-specific proteins. New hair fibers were seen 15e20 days postinjection, with the greatest number of hairs seen in the treatment group. Regenerated hairs in the control group were predominantly found in the periphery of the wound and were characteristically short and thin. By contrast, the DPLTs-treated mice had new hairs throughout the treatment area, as well as a significantly higher amount of sebaceous glands at days 10 and 20 and increased vascularization. These data suggest that engineered canine DPLTs demonstrate characteristics of dermal papillae and have a positive effect on hair regeneration. A phase 1/2 trial is currently under way in the United States to test the ability of autologous ADSCs to promote hair growth in men with male pattern baldness. A critical function of ADSCs is the secretion of growth factors that activate neighboring cells and mediate diverse skin-regenerative effects, including wound healing, antioxidant protection, and antiwrinkling [20]. ADSCs also secrete regenerative growth factors that participate in hair morphogenesis and regeneration. The hair growth promoting effect of ADSC-conditioned media (ADSC-CM) was studied by subcutaneous injection of the CM into C3H/NeH mice [21]. The CM contains a variety of growth factors, including VEGF, hepatocyte growth factor (HGF), insulin-like growth factor (IGF), and PDGF. The researchers discovered that culturing ADSCs under hypoxic conditions significantly improved their proliferative and self-renewal capacity and enhanced their growth factor secretion. The hypoxia-derived CM (HCM) induced the enhanced proliferation of human keratinocytes, as well as human follicle dermal papilla cells in vitro as compared to the CM from ADSCs grown under normoxic conditions (NCM). Based on these data, the induction of hair growth in 7-week-old C3H/HeN female mice (n ¼ 21) treated with HCM was assessed. Mice received three subcutaneous injections of 100 mL NCM or HCM. Injections were given at days 0, 3, and 6 and skin darkening as an early indicator for hair follicle formation was monitored for a 12-week period of time. After 8 weeks, dark spots and hair regeneration were seen in the areas of injection with complete hair growth seen by 12 weeks in the HCM-treated animals. Dark spots and hair regrowth were not seen in the NCMtreated mice. Clinical pilot studies have been performed to study the hair growth effect of ADSC-CM on both male and female patients [22]. In women with female pattern hair loss, ADSC-CM was applied to the scalp after microneedling, the use of ultrafine needles to interrupt the skin barrier. After 12 weeks there was a significant increase in both hair thickness and hair density, and no adverse reactions were reported. The researchers also studied the effect of intradermal injections in both men and women with hair loss. Twenty-two patients received six sets of injections every 3e5 weeks and this treatment regimen resulted in an increase in hair numbers after week 12. No large, controlled, double-blinded clinical studies have been conducted to date and research is being done to enhance the hair regeneration potential of ADSC-CM.

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TISSUE-DERIVED MATERIALS FOR HAIR REGENERATION Human placental extract (Hominis placenta [HP]) has been used for decades in China and Korea as a therapeutic agent for liver regeneration and endocrine abnormalities. HP extract has been shown to have antiinflammatory, wound-healing, and antioxidant effects in the clinical setting [23]. Initial studies with cow placenta showed an increase in hair density in the mouse model [24] and this led to an investigation into whether HP treatment could influence hair growth in mice. Fifteen 7-week-old male C57BL/6 mice were divided into control, minoxidil, or HP-treated groups. Animals were depilated to synchronize hair follicle growth to the anagen phase. Animals received topical treatment of the controls or treatment once daily for 15 days, after which time measurements were taken for hair growth and thickness by dermoscopy. Immunohistochemistry was also performed on skin sections from all animals to assess BrdU incorporation and DNA synthesis. Results demonstrated an increase in hair number, hair shaft diameter, and density in the HP-treated animals as compared to the control animals. In addition, high levels of BrdU incorporation were observed in the dermal papilla, ORS, and connective tissue of the HP-treated animals. Proliferating cell nuclear antigen was highly expressed in the lower regions of the hair matrix, indicating induced DNA synthesis in the treatment group. Fibroblast growth factor 7 (FGF-7), a critical growth factor in the stimulation of anagen, was seen to have an increased expression in the outer layer of the hair shaft and surface epithelium in the HP-treated group. Placenta extracts have been utilized in a variety of clinical indications, including wound healing where the growth factors, amino acids, and angiogenic-promoting agents have proven to be beneficial. Hair follicle cycling has a great deal in common with the processes involved in wound healing and HP continues to be assessed as an anagen-promoting stimulus.

ADDITIONAL STUDIES ON SECRETED GROWTH FACTORS AND HAIR GROWTH Additional preclinical studies have been performed to build on the research showing the importance of Wnt proteins and other growth factors in the stimulation of follicular stem cells to induce sustainable anagen. In one such study, MSCs were engineered to overexpress Wnt1a [25]. Wnt-CM was injected into the murine model discussed earlier and resulted in accelerated hair follicle progression from telogen to anagen and enhanced alkaline phosphatase expression in the DP region. Of great significance is the fact that hair induction-related genes were upregulated even when they had been first selectively downregulated by treatment with DHT. This study clearly demonstrated that MSC-produced Wnt1a promoted the ability of the DP cells to induce hair cycling and regeneration in a physiological manner in a model simulating male pattern baldness. In another study, neural stem cell (NSC) extract was used to stimulate the hair follicle niche to induce hair growth in the mouse model. Expression levels of multiple growth factors and signaling factors were studied and it was reported that NSC extract enhanced hair growth by activating hair follicle niches via the coregulation of TGF-B and bone morphogenetic protein (BMP) signaling pathways in the telogen phase. The activation and differentiation of intrafollicular hair follicle stem cells, extrafollicular DPCs, and matrix cells was also seen both in vitro and in vivo. Key growth factors, including HGF, IGF-1, keratinocyte growth factor (KGF), and VEGF, were also increased to enhance hair growth stimulation. This study clearly demonstrated the ability of NSC extract to promote hair growth through the direct regulation of hair follicle niches via the TGF-B and BMP signaling pathways, as well as the induction of critical growth factors [26].

SIMULATING THE EMBRYONIC ENVIRONMENT In the mid-1990s it was reported that scarless healing occurs in the embryonic/fetal environment even after extensive surgery [27]. The embryonic environment is known to be characterized by rapid cell growth and cell differentiation, and embryonic tissue develops under low oxygen (hypoxic) conditions, and, unlike adult tissue, it contains predominantly collagens III and V and low levels of TGF-B and bFGF. Ezashi et al. [28] have shown that mimicking the embryonic condition of hypoxia enhances the growth of human embryonic stem cells (hESCs) and both hESCs and adult stem cells show improved maintenance of their undifferentiated state under low oxygen conditions [28e30]. To assess the effect of embryonic conditions on normal cells, neonatal fibroblasts were seeded onto dextran beads and grown in suspension in controlled closed bioreactors under 5% oxygen. Under these cell culture conditions, over 5000 genes were differentially expressed as compared to identical growth conditions with normal oxygen. A large number of the upregulated genes were associated with multipotent stem cells, including SOX 2, Oct4,

SIMULATING THE EMBRYONIC ENVIRONMENT

FIGURE 74.2

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Expression of hair follicle-related stem cell markers in neonatal cells grown under simulated embryonic conditions.

FIGURE 74.3 Injection of hypoxia-induced multipotent stem cell (HIMSC)-conditioned media subcutaneously into the C-57bl/6 mouse model induced an increased density of hair follicles at the site of injections.

NANOG, and KLF4 genes [31]. The expression of stem cell-associated proteins, including Nodal, Brachyury, Nestin, and Oct4, was also seen. It was also noted that genes were upregulated and surface markers were expressed, which are normally associated with follicular stem cells, including Lhx2, SOX 21, Nestin, NFATc1, and Krt 15 (Fig. 74.2). Because of this finding, CM from the HIMSCs was utilized to study whether the composition could induce hair growth in the C57Bl/6 mouse model. Intradermal injections of the CM induced new hair growth in the treated mice and an increased number of hair follicles and hair follicle stem cells were noted in the treatment area (Fig. 74.3). No such hair induction was seen in the unconditioned media controls. The CM was analyzed and shown to contain high levels of growth factors associated with the induction and increased endurance of anagen, including HGF, FGF-7 or KGF, VEGF-A, follistatin, angiogenin, and placental growth factor (PlGF-1 or PlGF). Given this composition, a proof-of-concept clinical study was performed to assess the ability of the HIMSC-CM, called hair-stimulating complex (HSC), to induce hair growth in men with male pattern baldness [32]. The clinical study was a double-blind, placebo-controlled, randomized, single site trial and was designed to evaluate safety of the HSC product and assess efficacy in stimulating hair growth. All 26 subjects tolerated the single, intradermal injection of HSC procedures well, and no signs of an adverse reaction were reported. Histopathological evaluation of the treatment site biopsies taken at 22 and 52 weeks posttreatment revealed no abnormal morphology, hamartomas, or other pathological responses. Treatment and control regions received temporary tattoos so that photographs at each follow-up time point could be taken in exactly the same region. Trichoscan image analysis of HSC-treated sites at 12 and 52 weeks showed significant improvements in hair growth over the placebo. At the initial 12-week evaluation period, HSC-treated sites demonstrated an increase in hair shaft thickness (6.3%  2.5% vs. 0.63%  2.1%; P ¼ .046), thickness density (12.8%  4.5% vs. 0.2%  2.9%; P ¼ .028), and terminal hair density (20.6  4.9% vs. 4.4  4.9%; P ¼ .029). At the 1-year time point a statistically significant increase in total hair count (P ¼ .032) continued to be seen. The fact that hair counts continued to increase over the 1year period even after only one injection at day 0 supported the hypothesis that HSC stimulated a prolonged and sustainable anagen, a result that had not been reported in previous clinical hair growth studies. An example of such growth is seen in Fig. 74.4 where trichoscan data demonstrated a 123% increase in terminal hairs over the 52-week period. To assess the efficacy and safety of an increased number of injections and a second dosing at week 6, a phase 1/2 study was conducted in Manila with Investigational New Drug (IND) approval by the Food and Drug Administration Philippines. The clinical study was a double-blind, randomized, two center trial in 56 subjects. All subjects tolerated the eight 0.1 cc intradermal injections at baseline and 6 weeks well, and no signs of an adverse reaction were reported. Blood and urine samples taken before and after each injection set showed no liver, kidney, or bone marrow toxicity. Trichoscan image analyses of treated and control sites were taken at baseline, 12, 24, 36, and 48 weeks. At the

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FIGURE 74.4 One-time injection of hair stimulating complex (HSC) yielded significant new terminal hair growth in treated areas at 12 weeks

and 1 year.

12-week time point, significant improvements in total (P ¼ .0013), terminal (P ¼ .0135), and vellus (P ¼ .033) hair growth over baseline were seen, as was an increase in cumulative thickness density (P ¼ 026). The primary efficacy endpoint of increased terminal hair at 12 weeks was met, with a 19.5% increase seen, a 49.5% increase over the same endpoint in our proof-of-concept trial. In addition, unlike currently approved products, HSC induced hair growth in the temporal recession, as well as vertex and mid-scalp regions, and was highly effective in men over 40 years of age, as well as in men in their 20s and 30s. At the 48-week time point there continued to be a significant increase in total hairs over baseline (P ¼ .028). These results clearly demonstrate the safety and efficacy of intradermal injections of HSC in subjects with AGA and the benefit of a second treatment at 6 weeks to capture preanagen follicles that were too early in the cycle at week 0. An additional clinical study was conducted on five men and five women in the United States under a physician IND. This “safe passage” study involved 20 (0.1 cc) injections at baseline, 20 at week 6, and 40 injections at week 12 and week 40. The injections were well tolerated and no serious adverse effects were reported. Some subjects reported an itching at the injection sites, which was resolved within 30 min. An increase in hair growth, as demonstrated by global photography, was seen in all patients, which included men with male pattern baldness, as well as women with hair loss associated with stress, chemical damage, and perimenopausal hormonal changes. Fig. 74.5A and B show results in two of the 10 patients treated. To be able to do phase 2 dosing studies in men and to remove nongrowth factor-related impurities from the CM, HSC production was modified to include capture of the growth factors by a heparin sepharose column and the use of various filters to remove DNA. The six key growth factors in HSC that are associated with hair growth are measured by enzyme-linked immunosorbent assay before product release and sterility, bioburden, and a series of viral testing are performed. Cell potency assays are performed to assess the activity of follistatin, FGF-7/KGF, and the angiogenic factors in the product. The following is a summary of the proposed mechanism of action of HSC in the induction of hair growth. A number of growth factors work synergistically to induce and maintain angiogenesis in the hair follicle. The hair follicle undergoes cyclic expansion and regression, which requires changing demands for vascular support, with perifollicular vascularization being seen during anagen and decreased vascularization associated with catagen and telogen. VEGF helps to mediate the induction of angiogenesis to provide the increased nutritional needs of hair follicles during rapid cell division in anagen [33]. In addition to increasing vascularization, VEGF induces the division of DPCs, which is a key step in the induction of hair follicle formation. Angiogenin is a soluble factor normally secreted by DP cells [34]. Angiogenin is a potent angiogenic factor that also promotes the stimulation of DP and ORS keratinocytes and the subsequent elongation of the hair follicle. HGF is a multifunctional polypeptide that acts as a mitogen, motogen, and morphogen. In addition to promoting angiogenesis, it has been shown to stimulate follicle growth of human hair and follicle elongation by stimulating DP

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FIGURE 74.5 (A) Escalating doses of hair-stimulating complex (HSC) every 6 weeks induced cosmetically relevant new hair formation in a subject with male pattern baldness. (B) HSC induced new vellus hair formation within 18 weeks in the temporal recession region of a 38-year-old female with stress-induced alopecia.

cells and hair bulb-derived keratinocytes [35,36]. PlGF is a member of the VEGF subfamily and is important in promoting angiogenesis and vasculogenesis, particularly during embryonic development. PlGF has been shown to play a role in the promotion of hair growth by accelerating the growth of hair follicle cells and prolonging anagen hair growth [37]. The mechanism of action involves preventing hair follicle cell death by increasing levels of phosphorylated extracellular signal-related kinase and cyclin D1. FGF-7/KGF is synthesized by skin fibroblasts and DP cells and has a mitogenic effect on skin keratinocytes. It is thought to be the most potent growth factor reported to date [38]. The KGF receptor is found on ORS cells and addition of KGF to these cells in organ cultures has significantly increased hair follicle cell proliferation [39]. In vivo the FGF-7/KGF secreted by DP cells causes both rapid proliferation of ORS cells and the resulting keratinocytes to migrate into the hair follicle shaft and form a new vellus hair or convert vellus hairs to terminal hairs [40]. Follistatin is found in all human tissues and organs and has the primary function of binding and bioneutralizing members of the TGF superfamily, with particular affinity for activin. It is an antagonist to both activin and BMP, both of which are involved in maintaining a slow cycle of stem cell proliferation in resting hair follicles [41,42]. Follistatin has been shown to be an important regulator of cell proliferation, differentiation, and apoptosis in hair follicle initiation and hair cycling. In AGA, androgens induce TGF-B1, which inhibits growth of the dermal papilla cells and keratinocytes [43]. Adding follistatin to dermal papilla cells helps to reverse this blockage, further strengthening the role of this activin antagonist in hair growth [44]. Naturally occurring growth factors work synergistically to promote proliferation, differentiation, and angiogenesis. The composition of growth factors in HSC works together to induce angiogenesis of the hair follicle for adequate nutrient support, as well as stimulating DP cells and ORS cells to proliferate to induce hair follicle elongation, cell migration into the hair shaft, and subsequent hair growth. Given the early clinical results with HSC, a study was performed in Japan to see if it could be effective as a holding solution for follicular units implanted in hair transplant procedures [45]. Saline is the most common solution used to

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preserve hair follicles for the period of time between removal and implantation. Surgeons have come to accept a large degree of hair loss, known as effluvium, from the transplants, and patients need to be patient for the months it takes to allow the transplanted hair to recover from shock and enter into the anagen cycle. As expected in the weeks following transplant, the investigator reported that transplanted hairs held in saline were lost to effluvium at the 6-week follow-up. Of the hairs held in the human cell conditioned media (hCCM) solution, however, most remained intact at this follow-up time point, reflecting not only the more hospitable environment created by the hCCM but potentially also the increased wound-healing capabilities of the material. Fifty-five percent of the follicles soaked in saline for up to 2 h during the transplant procedure showed hair loss, whereas only 10% of the follicles soaked in hCCM resulted in showing the effluvium response.

BIOENGINEERING A HUMAN HAIR FOLLICLE A major challenge facing stem cell biologists is the bioengineering of an entire functional organ. Since the hair follicle is a miniorgan it is an ideal candidate to utilize as a goal to be the first organ to be engineered. Formation of such a follicle would require simulating the series of molecular and cellular events that will recreate follicle morphogenesis. The steps will progress from the formation of dermal condensations, to epithelial invagination, to DP formation at the base of the follicle, to the arrangement of the various cell types and their morphogenesis, to the differentiation of hair shaft components, and then to actually demonstrate the creation of a hair. Creation of a functional hair follicle would not only be a major milestone in regenerative medicine, it would allow the creation of tissue-engineered skin for burn patients that would provide the critical functions of skin appendages in addition to providing a novel therapy for people with alopecia. Early work studying how hair follicles are formed during early development resulted in the reconstitution of pelage hair follicles when a mixture of prepared cell suspensions from embryonic or neonatal mice was grafted onto the dorsal surface within a silicon chamber [46]. Zheng et al. [47] demonstrated that dissociated neonatal mouse epidermal and dermal cells could be reconstituted to form hair follicles. The de novo follicles formed in a sequence similar to that seen in embryonic development, showed follicular layering, formed a bulge region, as well as a sebaceous gland, and demonstrated appropriate biochemical differentiation markers and cycling. Isolated K15 enhanced green fluorescent protein-positive bulge cells showed stem cell properties and created new follicles. An additional important finding was that the ratio of epithelial and dermal cells injected affected the efficiency of the hair formation. The study showed that 5000 dermal cells and 2500 epidermal cells were necessary to produce a single follicle. Ebama et al. [1] used a similar technique to attempt to create a functional human follicle. Cografting of neonatal foreskin human keratinocytes with murine DP-enriched cells produced structures resembling hair follicles with multiple epidermal cell layers and a well-keratinized inner region. The follicular structures expressed hair keratin markers, as well as markers associated with developmental stages, but lacked normal hair. These results were also repeated with adult keratinocytes. Neither the enriched DP population nor the epidermal fraction alone formed any follicles, further elucidating the importance of the epithelialemesenchymal interactions in hair follicle regeneration. Building on the research with the mouse and human-dissociated cells to form new follicles, Sriwirlyanont and colleagues worked to produce functional follicles in engineered skin [48]. To date the growth of hair appendages in tissue-engineered skin substitutes has not been accomplished in a reproducible manner predominantly due to the lack of trichogenicity in postnatal cells. In this study a chimera of human neonatal-cultured keratinocytes and murine-cultured DP cells was grafted orthotopically to full thickness wounds on athymic mice. Noncultured dissociated murine skin cells were used as the positive control and human cultured keratinocytes and fibroblasts without DP cells were the negative control. Neonatal murine-only skin substitutes produced external hairs with sebaceous glands, while chimeric skin substitutes formed hairs without the gland. The latter was associated with the upregulation of the LEF-1 hair-related gene and the downregulation of stearoyl-CoA desaturase, a sebaceous gland marker. No visible hairs or glands were formed with human-only skin substitutes. Results of this study illustrated that sebaceous gland formation is not a prerequisite for hair growth in tissue-engineered skin substitutes. The regeneration of a human hair using epidermal and dermal components in ex vivo or mouse models has not been demonstrated to date. Research by Rahmani et al. [49] has resulted in identifying a cell type that has furthered progress in human hair regeneration and may lead to generating new hair in the human scalp. To assess how DP cells are maintained in healthy follicles an in vivo fate mapping of adult hair follicle dermal sheath (DS) cells was performed. It was seen that a subset of DS cells is retained following each hair cycle and that these cells exhibit a self-renewal characteristic and are key in repopulating the DS and DP with new cells. Hair regrowth was retarded when the DS cells

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were ablated and hair type specification was altered, supporting the hypothesis that the DS cells function to modulate normal DP function. This work identified a bipotent stem cell within the adult hair follicle mesenchyme and supported the approach of using the DS cells in the clinical setting. A phase 1/2a study utilizing an autologous cell-based treatment for AGA has been completed by RepliCel Life Sciences and has demonstrated safety and effectiveness. Subjects with mild-to-moderate AGA as characterized by the Ludwig scale (female subjects) or Norwood Hamilton scale (male subjects) were enrolled in the study. After blood sample testing substantiated their general health, scalp biopsies were taken and sent to a cyclic guanosine monophosphate-compliant facility. Biopsies were dissociated and treated to isolate dermal sheath cup (DSC) cells, which were then expanded in culture. The 19 subjects returned to the clinic after the DSC cells were grown to numbers sufficient for transplantation. The process for creating and implanting the DSCs is shown in Fig. 74.6. Subjects were their own control and received medium alone (placebo) or replicated DSC cells in medium into areas of their scalp that were randomized by a predetermined blind. The primary protocol of the study was to assess local treatment sites for any abnormalities at the 6-month time point. The secondary endpoints at 6 and 24 months assessed both safety and efficacy. At the 6-month interim point the data collected showed no serious adverse reactions in any of the subjects. Thirty percent of the subjects experienced a burning sensation at the site of treatment, which was resolved within 24e48 h. At 6 months, several efficacy endpoints were successfully met as well. Significantly more patients demonstrated a greater than 5% increase in hair density at the DSC-treated areas, with density increases ranging from 5% to 19.6%. Increases were seen in total hair density, as well as vellus and terminal hair density. Additional clinical studies are currently under way in Japan and are being conducted by Shiseido Company, a world leader in aesthetic products. The clinical data to date support the importance of DS cells in hair growth, particularly with the cosmetically relevant increase in hair density. Similar autologous cell-based studies have been performed with a combination of epithelial and dermal cells or dermal cells alone by both Intercytex Group plc and Aderans Research. Trials have consistently shown safety in all patients and signals of efficacy in a number of treated subjects. In autologous cell-based trials a number of questions still remain to be answered. These include the mechanism of action and whether the injected cells actually develop a new follicle, become part of an existing follicle, or stimulate hair growth by releasing growth factors and exosomes that then increase growth in the targeted hair follicle. The

FIGURE 74.6 Collecting and processing dermal sheath fibroblasts for implantation for hair regrowth.

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number of cells necessary to be utilized to stimulate hair growth is key, since the ratio of new hair follicles to donor follicles must be great enough to produce a commercially feasible product. Being able to measure and retain trichogenicity of the implanted cells is essential, as well as being able to understand whether a regenerating system in vivo has the ability to attract other cells from the body, as is seen with the recruitment of bone marrow cells in the process of normal wound healing, which should offer a real benefit in new organ formation.

SUMMARY The ultimate goal in organ replacement regenerative medicine is the ability to create fully functional organs that have been damaged or destroyed by injury, aging, or disease. The hair follicle is an ideal miniorgan, which has been targeted to achieve this goal since it reforms itself throughout life through key interactions between epithelial cells and inductive cells from the mesenchyme. Whether success will be achieved by creating the follicle in vitro in complex three-dimensional culture systems or injecting a combination of cells and growth factors into the scalp for remodeling in vivo is still unknown. Studying follicular neogenesis in the embryonic environment has led to tremendous advances in understanding the complex signaling pathways and epithelialemesenchymal interactions necessary to stimulate human hair growth. Research by several laboratories has underscored the powerful inductive ability of the follicular dermal cells. Although current cellular implant approaches have focused on autologous dissociated cells from the scalp, the potential of an allogeneic approach is very feasible. Dermal hair follicle tissue has successfully been transplanted between individuals without any sign of rejection [50]. In addition, it is believed that the hair follicle is an immune privileged location that does not express the major histocompatibility complex class 1 antigen that is associated with the rejection process [51]. To make progress in either autologous or allogeneic expanded cells for transplantation we must make progress with developing solutions for maintaining trichogenicity of the cells in culture. A very intriguing proposition is the utilization of iPS cells or other stem cells in bioengineering a human hair follicle. A number of stem cell types, including neural [52] and bone marrow-derived MSCs [53], have shown the capacity to form skin and hair when injected into a blastocyst. A product designed to engineer new hair follicles, consisting of inductive dermal cells and competent epithelial cells, is likely to be the first organ system to be successful in the clinic.

References [1] Ehama R, Ishimatsu-Tsujy Y, Iriyama S, Ideta R, Soma T, Yano K, et al. Hair follicle regeneration using grafted rodent and human cells. J Invest Dermatol 2007;127:2106e15. [2] Plikus MV, Sundberg JP, Chuong C-M. Mouse skin ectodermal organs. In: Fox J, Barthold S, Davisson M, Newcomer C, Quimby F, Smith A, editors. The mouse in biomedical research. 2vol. 3. Amsterdam: Academic Press; 2007. [3] Fuchs E. Scratching the surface of skin development. Nature 2007;445:834e42. [4] Plikus MV, Chuong CM. Complex hair cycle domain patterns and regenerative hair waves in living rodents. J Invest Dermatol 2008;128:1071e80. [5] Naughton G, Mansbridge J, Gentzkow G. A metabolically active human dermal replacement for the treatment of diabetic foot ulcers. Artif Organs 1997;21(11):1203e10. [5a] Ito M, Cotsarelis G. Is the Hair Follicle Necessary for Normal Wound Healing? J Invest Dermatol 2008;128(5):1059e61. [6] Schmidt-Ulrich R, Paus R. Molecular principles of hair follicle induction and morphogenesis. Bioessays 2005;27:247e61. [7] Lacassagne A, Latarjet R. Action of methylcholanthrene on certain scars on the skin in mice. Can Res 1946;6:183e8. [8] Billingham RE, Russell PS. Incomplete wound contraction and the phenomenon of hair neogenesis in rabbit skin. Nature 1956;177:791e2. [9] Kligman AM, Strauss JS. The formation of vellus hair follicles from human adult epidermis. J Invest Dermatol 1956;27:19e23. [10] Ito M, Yang Z, Andl T, Cui C, Kim N, Miller S, Cotsarelis G. Wnt-dependent de novo hair follicle regeneration in adult mouse skin after wounding. Nature 2007;447:316e21. [11] Ellis JA, Stebbing M, Harrap SB. Polymorphism of the androgen receptor gene is associated with male pattern baldness. J Invest Dermatol 2001;116(3):452e5. [12] Hadshiew IM, Foitzik K, Arck PC, Paus R. Burden of hair loss: stress and the underestimated psychosocial impact of telogen effluvium and androgenetic alopecia. J Invest Dermatol 2004;123(3):455e7. [13] Gilhar A, Keren A, Shemer A, d’Ovidio R, Ullmann Y, Paus R. Autoimmune disease induction in a healthy human organ: a humanized mouse model of alopecia areata. J Invest Dermatol 2013;133(3):844e7. [14] Garza LA, Yang CC, Zhao T, Blatt HB, Lee M, He H, Stanton DC, Carrasco L, Spiegel JH, Tobias JW, Cotsarelis G. Bald scalp in men with androgenetic alopecia retains hair follicle stem cells but lacks CD200-rich and CD43-positive hair follicle progenitor cells. J Clin Invest 2011;121:613e22. [15] Greco V, Chen T, Rendl M, Schober M, Pasolli HA, Stokes N, Dela Cruz-Racelis J, Fuchs E. A two step mechanism for stem cell activation during hair regeneration. Cell Stem Cell 2009;4:155e60. [16] Anitua E, Fino A, Martinez N, Orive G, Berridl D. The effect of plasma rich in growth factors on pattern hair loss: a pilot study. Dermatol Surg 2017;43(5):658e70.

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[17] Zuk PA, Zhu M, Mizuno H, Huang J, Futrell JW, Katz AJ, Benhaim P, Lorenz HP, Hedrick MH. Multilineage cell from human adipose tissue: implications for cell-based therapies. Tissue Eng 2001;7:211e28. [18] Lee A, Bae S, Lee SH, Kweon OK, Kim WH. Hair growth promoting effect of dermal papilla like tissues from canine adipose-derived mesenchymal stem cells through vascular endothelial growth factor. J Vet Med Sci 2016;78(12):1811e8. [19] Chen CC, Plikus MV, Tang PC, Widelitz RB, Chuong CM. The modulatable stem cell niche: tissue interactions during hair and feather follicle regeneration. J Mol Biol 2016;428(7):1423e40. [20] Kim WS, Park BS, Sung JH, Yang JM, Park SB, Lee GY, Kim KJ, Whang KK, Kang SH, Park BS, Sung JH. Wound healing effect of adiposederived stem cells: a critical role of secretory factors on human fibroblasts. J Dermatol Sci 2007;48:15e24. [21] Park BS, Kim WS, Choi JS, Kim HK, Won WH, Ohkubo F, Fukuoka H. Hair growth stimulated by conditioned medium of adipose derived stem cells is enhanced by hypoxia: evidence of increased growth factor secretion. Biomed Res 2010;31(1):27e34. [22] Fukuoka H, Suga H. Hair regeneration treatment using adipose derived stem cell conditioned medium: follow up with trichograms. Eplasty 2015;15:e10. [23] Seo TB, Han IS, Yoon JH, Seol IC, Kim YS, Jo HK. Growth promoting activity of Hominis Placenta extract on regenerating sciatic nerve. Acta Pharmacol Sin 2006;27(1):50e8. [24] Zhang D, Lijuan G, Jingjie L, Zheng L, Wang C, Wang Z. Cow placenta extract promotes murine hair growth through enhancing the insulinlike growth factor-1. Indian J Dermatol 2011;56(1):14e8. [25] Dong L, Hao H, Xia L, Liu J, Ti D, Tong C, Hou Q, Han Q, Zhao Y, Liu H, Fu X, Han W. Treatment of MSCs with Wnt1a-conditioned medium activates DP cells and promotes hair follicle regrowth. Sci Rep 2014;4:5432e40. [26] Hwang I, Choi KA, Park HS, Hang-Soo J, Hyesun K, Jeong-Ok S, Seol K, Kwon HJ, Park IH, Hong S. Neural stem cells restore hair growth through activation of the hair follicle niche. Cell Transplant 2016;25(8):1439e51. [27] Adzick NS, Lorenz HP. Cells, matrix, growth factors, and the surgeon. The biology of scarless fetal wound repair. Ann Surg 1994;220:10e8. [28] Ezashi T, Das P, Roberts RM. Low O2 tension and the prevention of differentiation of hES cells. Proc Natl Acad Sci USA 2005;102:4783e8. [29] Chakravarthy MV, Spangenburg EE, Booth FW. Culture in low levels of oxygen enhances in vitro proliferation potential of satellite cells from old skeletal muscles. Cell Mol Life Sci 2001;58:1150e8. [30] Covello KL, Kehler J, Yu H, et al. HIF-2alpha regulates Oct-4: effects of hypoxia on stem cell function, embryonic development, and tumor growth. Gene Dev 2006;20:557e70. [31] Pinney E, Zimber M, Schenone A, Montes-Camacho M, Ziegler F, Naughton GK. Human embryonic-like ECM (hECM) stimulates proliferation and differentiation in stem cells while killing cancer cells. IJSC 2011;4(1):70e5. [32] Watt FM, Jensen KB. Epidermal stem cell diversity and quiescence. EMBO Mol Med 2009;1:260e7. [33] Li W, et al. VEGF induces proliferation of human hair follicle dermal papilla cells through VEGFR-2 mediated activation of ERK. Exp Cell Res 2012;318(14):1633e40. [34] Zhou N, Fan W, Li M. Angiogenin is expressed in human dermal papilla cells and stimulates hair growth. Arch Dermatol Res 2008;301(2): 139e49. [35] Shimaoka S, et al. Hepatocyte growth factor/scatter factor expressed in follicular papilla cells stimulates human hair growth in vitro. J Cell Physiol 1995;165:333e8. [36] Jindo T, Imai R, Tsuboi R, Ogawa H. Hepatocyte growth factor/scatter factor stimulates hair growth of mouse vibrissae in organ culture. J Invest Dermatol 1994;103(3):306e9. [37] Yoon S, et al. A role of placental growth factor in hair growth. J Dermatol Sci May 2014;74(2):125e34. [38] Guo L, Degenstein L, Fuchs E. Keratinocyte growth factor is required for hair development but not for wound healing. Genes Dev 1996;10: 165e75. [39] Yang K, Brown LF, Detmar M. J Clin Invest February 15, 2001;107(4):4098e417. [40] Jang JH. Stimulation of human hair growth by recombinant keratinocyte growth factor. Biotechnol Lett June 2005;27(11):749e52. [41] McDowall M, Edwards NM, Jahoda CAB, Hynd PI. The role of activins and follistatin in skin and hair follicle development and function. Cytokine Growth Factor Rev 2008;19(5):415e26. [42] Inui S, Fukuzato Y, Nakajima T, Yoskikawa K, Itami S. Androgen-inducible TGF-B1 from balding dermal papilla cells inhibits epithelial cell growth: a clue to understand paradoxical effects of androgen on human hair growth. FASEB J 2002;16:1967e9. [43] Nakamura M, et al. Control of pelage hair follicle development and cycling by complex interactions between follistatin and activin. FASEB J 2003;17:497e9. [44] Rendl M, Polak I, Fuchs E. BMP signaling in dermal papilla cells is required for the hair follicle-inductive properties. Genes Dev 2008;22: 543e57. [45] Sadick N, Zimber MP, Cooley J, Yagyu K, Naughton GK. Embryonic-like secreted proteins enhance hair follicular unit viability and posttransplant healing. In: ISHRS annual meeting review; 2011. [46] Lichti U, Weinberg WC, Goodman L, Ledbetter S, Doolet T, Morgan D, Yuspa SH. In vivo regulation of murine hair growth: insights from grafting define cell populations onto nude mice. J Invest Dermatol 1993;101:124Se9S. [47] Zheng Y, Du X, Wang W, Boucher M, Parimoo S, Stenn K. Organogenesis from dissociated cells: generation of mature cycling hair dollicles from skin-derived cells. J Invest Dermatol 2005;124:867e76. [48] Sriwirlyanont P, Lynch K, McFarland K, Supp D, Boyce S. Characterization of hair follicle development in engineered skin substitutes. PLoS One 2013;8(6):1e10. [49] Rahmani W, Abbasi S, Hagner A, Raharjo E, Kumar R, Hotta A, Magness S, Metzger D, Biernaskie J. Hair follicle dermal stem cells regenerate the dermal sheath, repopulate the dermal papilla, and modulate hair type. Dev Cell 2014;31:543e58. [50] Reynolds AJ, Lawrence C, Cserhalmi-Friedman PB, Christiano AM, Jahoda CA. Trans-gender induction of hair follicles. Nature 1999;402: 33e4. [51] Paus R, Eichmuller S, Hofmann U, Czarnetzki BM, Robinson P. Expression of classical and nonclassical MHC class 1 antigens in murine hair follicles. Br J Dermatol 1994;131:177e83.

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[52] Clarke DL, Johansson CB, Wilbertz J, Veress B, Nilsson E, Karlstrom H, Lendahl U, Frisen J. Generalized potential of adult human stem cells. Science 2000;288:1660e3. [53] Jiang Y, Jahagirdar BN, Reinhardt RL, Schwartz RE, Keene CD, Ortiz-Gonzalez XR, Reyes M, Lenvik T, Lund T, Blackstad M. Pluripotency of mesenchymal stem cells derived from adult marrow. Nature 2002;418:41e9.

Further Reading Blanpain C, Fuchs E. Epidermal homeostasis: a balancing act of stem cells in the skin. Nat Rev Mol Cell Biol 2009;10:207e17.

C H A P T E R

75 US Stem Cell Research Policy Josephine Johnston, Rachel L. Zacharias The Hastings Center, Garrison, NY, United States

INTRODUCTION Since James A. Thomson and colleagues reported the isolation of pluripotent stem cells from human embryos in 1998 [1], stem cell research has received considerable public and policy attention. Because the isolation of embryonic stem cells involves destroying human embryos, some groups and individuals have opposed, or raised concerns about, some or all of the research on moral grounds. The opposition and concern have influenced a number of policies and laws at state, federal, and international levels. From 2001 until 2009, much embryonic stem cell research was ineligible for federal funding in the United States. In 2009, federal rules were relaxed somewhat, although federal funds still may not be used to create embryonic stem cells or to support research on cells taken from embryos created for research purposes, including by cloning. Despite these restrictions, embryonic stem cell research progressed into the 21st century in the United States using monies supplied by individual donors, charitable organizations, and states. After briefly discussing ethical and policy issues in adult and fetal stem cell research, this chapter will survey the current policy issues in embryonic and induced pluripotent stem cell (iPSc) research, beginning with federal and state funding policies, which will be compared to regulation strategies adopted in other nations active in the research, before considering oversight, commercialization, and ethical issues that arise as the research and technology move forward.

SOURCES OF STEM CELLS Stem cells are special kinds of cells that can regenerate themselves and make new, more specialized cells. For the purposes of an ethics and policy discussion, stem cells can be divided into four kinds based on the source of the cells: adult stem cells, fetal stem cells, iPScs, and embryonic stem cells. In terms of ethics, politics, policy, and law, much depends on the source of the cells.

Adult Cells “Adult” stem cells are derived from the somatic cells of adults and children (frequently taken from an umbilical cord). Although obtaining adult stem cells can raise some ethical issues, they are similar to those raised in other kinds of research involving human subjects, the most important of which is the requirement for free and informed consent [2]. Because the ability of competent adults to consent to research enjoys wide acceptance, somatic stem cell research has not been a major focus of ethical or political debate. It does, however, enter public consciousness as a less controversial, proposed alternative to embryonic stem cell research. Opponents of embryonic stem cell research cite clinical success of adult stem cells as evidence that research using embryonic cells is unnecessary [3,4]. However, many adult stem cell scientists insist on the importance of pursuing research on stem cells from both sources [5,6].

Fetal Cells Pluripotent stem cells have also been extracted from the primordial reproductive tissue of aborted fetuses [7]. Any source of stem cells that relies on women undergoing elective terminations is likely to be controversial in the United Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00075-8

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States simply because of the ongoing debate over the morality and legality of abortion. Nevertheless, researchers have used tissue from aborted fetuses since as early as the 1930s and federal funds are currently available for this kind of stem cell research. Although research involving human fetuses has been regulated in the United States since allegations of experiments on in and ex utero fetuses emerged in the early 1970s [8] (and around the same time as the Supreme Court’s famous 1973 decision on the legality of abortion in Roe v. Wade [9]), it seems that these regulations do not apply to most stem cell research using fetal cells. Human subject regulations, promulgated in 1975, require extra protections where federally funded research involves pregnant women, fetuses, and human in vitro fertilization (IVF). Among other things, these regulations forbid researchers from having any involvement in a woman’s decision to terminate her pregnancy or from offering gestational carriers financial inducements [10]. Although it has always been clear that these regulations applied to in utero fetal research, there has been some uncertainty as to whether the regulations apply also to research that, like stem cell research, uses cadaveric fetal tissue [11]. In March 2002, the Office for Human Research Protections issued a guidance for research involving fetal stem cells in which it states that the research will only be subject to federal human subject protections where it involves “a living individual,” which would exclude fetal stem cell research unless the formerly pregnant woman is involved in the research [11]. Likewise, the 1993 federal legislation on fetal tissue transplantation research likely does not apply to stem cell research using fetal cells unless the research also involves transplanting the cells into humans. Similar to the restrictions imposed by the federal human subject regulations, this legislation stipulates that no alteration in the timing, method, or procedures used to terminate the pregnancy should be made solely for the purposes of obtaining the tissue, nor can tissue be paid for or taken without the consent of the gestational carrier [12]. This legislation reappeared in public discourse in 2015, when the reproductive health care provider Planned Parenthood was accused of illegally profiting from fetal tissue donated to research [13]. Under the 1993 law, centers such as the accused California clinic can be reimbursed for the cost of obtaining, storing, and processing fetal tissue, but cannot be paid for the donation itself. Although 12 state inquiries found no evidence of illegal activity on behalf of Planned Parenthood, the organization stopped obtaining reimbursements for fetal tissue in October 2015 [14]. Even though both the federal regulations and the fetal tissue transplantation legislation likely do not apply to most fetal stem cell research, practices similar to those required by these laws (and lessons learned from cases that arise from their interpretation) will likely be considered by Institutional Review Boards (IRBs) reviewing fetal stem cell research. Various states also have laws affecting fetal stem cell research, which will apply to all researchers in those states regardless of their funding source. For example, five states ban research involving aborted fetuses and 12 states ban paying for fetal remains [15]. Despite fetal stem cell research’s intimate connections with fetal tissue research and the controversial practice of induced abortion, this kind of stem cell research has seldom been the subject of public debate.

Embryos Instead, debate has consistently focused on research that uses stem cells extracted from 4- to 7-day-old human embryos, referred to broadly as human embryonic stem cells (hEScs). Because any single cell in the very early human embryo can develop into a whole fetus, it is thought that embryonic stem cells have the potential to develop into almost any cell type and repair damaged or diseased parts of the human body. The therapeutic potential of embryonic stem cells therefore is thought to be enormous, but so is the moral peril because extracting these stem cells generally necessitates destroying the embryo. For this reason, the research has been vigorously opposed by many individuals and groups, including (but not limited to) those who consider the early embryo to be a person or, if not a full person, an entity of such sufficient moral significance that it should not be created for, or destroyed in, research. Embryonic stem cells can be derived from embryos created by two mechanisms: IVF or cloning. IVF embryos are created through the combination of an egg and sperm cell in a lab. Embryos produced through IVF may have been produced for reproductive uses but then determined no longer to be clinically necessary (so-called “spare” or “surplus” embryos), or they may have been produced specifically for research use. This difference in the reasons for creating the embryos is worth understanding. Creating embryos for research use is often opposed on the grounds that it is wrong to create human life for the purposes of destroying it [16]. Research use of spare embryos, on the other hand, has received more support on the ground that the embryos would be destroyed anyway. Whether the intention of the original creator of the embryos is sufficient reason for permitting research on spare embryos but not on embryos created for research use has been questioned [17,18]. Even if this moral issue can be

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resolved, spare embryos may not be satisfactory as the sole source of embryos for stem cell research because there may not be sufficient numbers available to meet demand. A 2003 survey of US fertility clinics reports that of the more than 400,000 embryos in frozen storage in the United States, only 2.8% have been donated to research [19]. Other researchers have argued that these frozen embryos will not be genetically diverse enough for therapeutic purposes [20]. Whether for these reasons or others, the US National Academies, which issued guidelines for embryonic stem cell research in 2004 (discussed in more detail later), proceeded on the basis that researchers could use spare embryos and embryos created for research use, including by cloning. The arguments against creating life purely for research use are also used to argue against using cloning techniques to generate human embryos. Currently, cloned embryos are created through a technology called somatic cell nuclear transfer (SCNT) (SCNT was used in the cloning of Dolly the Sheep in 1997 [21]). In ethics and policy discussions, cloning is usually broken down into two types based on the end goal of the cloning: therapeutic or research cloning and reproductive cloning. Therapeutic cloning uses SCNT to create cloned cells that will only be used in vitro, including embryos or somatic cells from which hEScs or iPScs can be derived, respectively. In therapeutic cloning, cloned embryos are not implanted for reproduction, and are discarded after stem cells have been derived [22]. Reproductive cloning is the creation of a cloned embryo through SCNT with the goal of creating a new living person. While therapeutic cloning destroys embryos, reproductive cloning would, if successful, result in the birth of a cloned human. Therapeutic and reproductive cloning are treated differently in policy in the United States and internationally.

Induced Pluripotent Stem Cells In 2002, Catherine Verfaillie and colleagues announced that they had isolated multipotent adult progenitor cells from bone marrow [23]. Then, in 2006, scientists in Japan and the United States announced the creation of pluripotent stem cells from adult skin cells [24,25]. These iPScs are human somatic cells reprogrammed to develop into nearly every human cell type, and are believed to be functionally very similar or identical to embryonic stem cells. Many consider these somatic cell-based lines to be an ethically simpler alternative to embryonic stem cells that could satisfy the same scientific and clinical needs. In the past decade, iPSc technologies have contributed considerably to research on human diseases, drug screening, and, with the introduction of gene editing, offer a virtually unlimited supply of human cell lines to research efforts [26]. Debates continue regarding the scientific differences between hESc and iPSc lines, and the significance (if any) of the ways in which they vary [27,28]. Due to the self-renewing quality of both cell types, hESc and iPSc may introduce a risk of neoplasm and tumor formation if the cells are not engineered properly. As such, hEScs were only approved in clinical trials in 2009, and iPScs have yet to be used clinically [29,30]. While embryonic stem cells frequently garner ethical attention surrounding the destruction of embryos, pluripotent stem cells have raised separate ethical considerations, mainly stemming from their creation by human cloning technologies. At the time of writing, research remains active in both the embryonic stem cell and iPSc domains. While future developments could shift research toward exclusive use of iPScs, past and current state, federal, and international policies have centered primarily on embryonic stem cell research.

UNITED STATES FEDERAL AND STATE STEM CELL POLICY History of US Stem Cell Law and Policy Federal policies affecting human stem cell research include those targeted specifically at the research, or aspects of it, and more general policy that has implications for stem cell research. In both cases, most of these policies affect hESc research. An example of policy that predates the derivation of hEScs is federal regulation of research involving human embryos, which began in the 1970s. Following the 1973 decision of the Supreme Court in Roe v. Wade, the United States Department of Health, Education and Welfare (DHEW) (a predecessor to the Department of Health and Human Services [HHS]) placed a moratorium on research with living human embryos. In 1974 Congress followed suit, creating their own moratorium on federal funding for research on embryos and embryonic tissue. While the DHEW moratorium was lifted in 1975, a de facto Congressional ban remained, and was translated into legislation in 1995 with the passage of the DickeyeWicker Amendment [31,32]. The amendment specifies that no HHS (which includes the National Institutes of Health [NIH]) funds can be used for the creation or destruction of a human embryo for research

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purposes. Since 1996, this policy has appeared as a rider to the annual appropriation acts for the Departments of Labor, Education, and Health and Human Services. Since then, no research on human embryos themselves, including the derivation of stem cells and the creation of human embryos through cloning, has been supported with federal funds. Instead, funds from states and private sources have been used to create embryonic stem cell lines. Once those lines have been created, federal funding has been available for research on them (see later). While the terms of this funding have varied by administration, Presidents Clinton, Bush, and Obama formulated policies to allow for federal funding of research on hEScs, within the bounds of the DickeyeWicker Amendment. Unless Congress removes this provision from HHS appropriations, or until embryonic stem cells (or their equivalent) can be extracted without destroying or harming human embryos [33], federal funding cannot be made available for the critical derivation step in the research. Following the announcement of derivation of hEScs in 1998, President Clinton developed a policy to permit use of federal funds to study embryonic stem cells without contravening the DickeyeWicker Amendment [2]. Under the policy, federal funds could be used for research on embryonic stem cells after they had been extracted from the embryo, but the funds could not be used for the crucial extraction step in the process. In opposing President Clinton’s policy, some members of Congress complained that although it did not pay for embryo destruction, it nevertheless encouraged research that required the destruction of human embryos [2]. The basic reasoning in President Clinton’s policydthat extraction of stem cells from embryos is a separate step from research on those cells once extracteddhas persisted in the policies of Presidents Bush and Obama since, and was tested in a federal court in 2010. However, President Clinton’s policy itself was not able to be implemented before his presidency ended, and the federal policy was altered dramatically in 2001, when President George W. Bush announced that he would limit federal funding to the study of embryonic stem cells that were already in existence at the time that his policy was announced (9 p.m. Eastern Standard Time on August 9, 2001). Research on newly derived cell lines would not be permitted with federal money. Other conditions included that the cells must have been extracted from “spare” embryos that were donated with the informed consent of the donors, to whom no financial inducements may have been offered [34]. To facilitate research on existing embryonic stem cells that met these criteria, the NIH set up a stem cell registry shortly after the policy was announced listing embryonic stem cell lines eligible for use in federally funded research and available for shipping. In his 2001 address, President Bush stated that more than 60 embryonic stem cell lines met his criteria, but for a variety of reasons only 22 lines were consistently listed in the registry as available for shipping. In explaining his policy, President Bush explicitly referred both to scientists’ beliefs in the enormous therapeutic potential of embryonic stem cell research and to his own belief in the value of embryonic life. Of embryonic stem cell research, he noted: “At its core, this issue forces us to confront fundamental questions about the beginnings of life and the ends of science. It lies at a difficult moral intersection, juxtaposing the need to protect life in all its phases with the prospect of saving and improving life in all its stages.” He called himself “a strong supporter of science and technology” but also noted that he believes “that human life is a sacred gift from our Creator” [35]. If President Bush intended his policy as a compromise between the value of scientific research and the value of embryonic life, it was one that left many scientists, disease groups, and others unsatisfied. Substantial criticism was directed at the 2001 policy. Those who oppose any research in which human embryos are destroyed argued that the federal funding restrictions needed to be supplemented with a nationwide ban on creating embryos for use in research, including by cloning [2,36]. A more vocal opposition called the 2001 policy overly restrictive. Their arguments included that the President overvalued embryonic life (that it is not more important than research aimed at treating disease) [2], that the policy was arbitrary and inconsistent [2], that the cell lines in the NIH registry were of poor quality and inappropriate for long-term use [37], and that the policy harmed American science by encouraging scientists to focus on research for which federal funding is available or to move to other countries to conduct their research [2]. During the Bush presidency, stem cell research was an important and in many cases partisan political issue. Federal policy was debated during the 2004 presidential campaign [38] and was a major issue at the Democratic National Convention, where Ronald Regan Jr., son of former Republican President Ronald Regan, gave a speech calling for greater government support of embryonic stem cell research. State policies to fund hESc research ineligible for federal support (described later in the chapter) were in some cases motivated by disagreement with President Bush’s 2001 policy. Other political activity, however, was bipartisan, including two 2004 lettersdone signed by over 200 congressional representatives and the other signed by nearly 60 senatorsdasking President Bush to relax federal funding restrictions on embryonic stem cell research. Similar bipartisan support was expressed for the Stem Cell Research Enhancement Act, introduced in 2005 and 2007, which would have provided federal support for hESc research using embryos created for IVF, no longer clinically necessary and donated with informed consent [39e41]. Both attempts

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were passed by the House and Senate, but vetoed by President Bush (an override was unsuccessful in 2005 and not attempted in 2007). Republican politicians who supported changes in federal policy included Senator Orrin Hatch, who mixed opposition to abortion with support for embryonic stem cell research provided it used only spare embryos [42]. Although the Stem Cell Research Enhancement Act was never passed, its language and motivations contributed to a shift in federal policy at the commencement of the Obama administration.

Current US Stem Cell Research Policy Federal Policy At the time of writing, US policy for embryonic stem cell research is contained in President Obama’s March 9, 2009 Executive Order [43], the NIH Guidelines on Human Stem Cell Research [44], and any applicable state laws [45]. The 2009 Executive Order and NIH guidelines establish the criteria for embryonic stem cell research that uses federal funds. They do not apply to research conducted using nonfederal monies, such as might be supplied by an individual donor, a company, a charitable organization, or a state government. Nevertheless, because much basic biomedical research is traditionally funded by the federal government, the federal funding policies have a significant impact on embryonic stem cell research in the United States. Executive Order 13505 does little more than revoke the policy announced in 2001 by President George W. Bush, state that the NIH may fund research using embryonic stem cells, and request that they issue new guidance for such research. The particulars of the new federal funding policy are to be found in the NIH Guidelines on Human Stem Cell Research, released in 2009. Under these guidelines, federal funding may be used to study embryonic stem cells derived from embryos “that were created using in vitro fertilization for reproductive purposes and were no longer needed for this purpose” (sometimes referred to as spare, surplus, or leftover embryos) and that were donated by the individuals who sought reproductive treatment. The donors must have given their voluntary written consent for the embryos to be used in research, and documentation must be available to show, among other things, that the donors received no cash or in-kind payment in exchange for making the donation and that there was a clear separation between the decision to create the embryos for reproductive purposes and the decision to donate. The guidelines also establish a registry of cell lines that meet federal funding criteria and describe procedures for establishing eligibility for funding of lines that are not already on the NIH registry. Because of the continued applicability of the Dickeye Wicker Amendment, federal funding is not available for the derivation of stem cells from embryos or research on cells taken from embryos that were created for research purposes, including by cloning [44]. Although considerably less controversial, President Obama’s Executive Order and the NIH guidelines that followed have also been criticized for being too permissive by those opposed to embryonic stem cell research and too restrictive by those who believe federal funding should be available for research on stem cell-extracted embryos created for research purposes, including by cloning [46]. Overall, stem cell policy under President Obama has received little political attention. The most notable opposition came by way of the courts, just months after the announcement of Obama’s executive action. A suit was brought against the Obama administration (specifically, HHS Secretary Kathleen Sebelius) by two adult stem cell researchers, James Sherley and Theresa Deisher, who argued that federal policy violated the spirit of the DickeyeWicker Amendment by funding research that relies on the destruction of human embryos. The case succeeded initially in the federal district court for the District of Columbia [47]. But after a series of appeals, the US Court of Appeals for the District of Columbia Circuit overturned the decision in 2012, citing no wrongdoing or violation of DickeyeWicker in the NIH’s interpretation of the guidelines [42]. In 2013, the Supreme Court refused to hear the case, decidedly upholding NIH’s 2009 guidelines and the limited federal funding of hESc research [48] (Figs. 75.1 and 75.2). State Policy and Private Funding In response to President Bush’s 2001 hESc research funding policy, advocates lobbied both state governments and private parties to provide funds for the research. The funding helped propel the research forward. Harvard University’s Douglas Melton and colleagues announced in 2004 that they had derived 17 new embryonic stem cell lines with support provided by the Juvenile Diabetes Research Foundation, the Howard Hughes Medical Institute, and Harvard University [49]. According to a 2005 special report in Scientific American, about $200 million of private money was spent on US stem cell research annually during the Bush administration [50]. We could find no reports of whether or how private funding of hESc research has changed since 2009. However, private foundations continue to contribute heavily to hESc research despite the increased federal funding under the Obama administration [51].

1314

75. US STEM CELL RESEARCH POLICY

FIGURE 75.1 Federal funding for embryonic versus nonembryonic stem cell research (FY2003eFY2017). Changes in federal funding for

embryonic and nonembryonic human and nonhuman stem cell research between the years 2003 and 2017. The red line denotes the shift from the Bush administration to the Obama administration, at which time US federal stem cell research and funding policies changed significantly.

FIGURE 75.2 Federal funding for stem cell research by type (FY2012eFY2017). Federal funding allocations for stem cell research by type

throughout the Obama administration (postfunding policy shift, 2012 through 2017). Data courtesy of National Institutes of Health Research Portfolio Online Reporting Tools. Estimates of funding for various research, condition, and disease categories (RCDC); February 10, 2016. https://report.nih.gov/ categorical_spending.aspx.

In response to President Bush’s hESc research policy, states also moved to fund stem cell research, as outlined in Table 75.1. In 2004, New Jersey became the first state to announce funding for local stem cell researchers, creating the Stem Cell Institute of New Jersey and the New Jersey Stem Cell Research Grant Program [52]. This program, as well as one in Illinois, has since ended, but funding continues to be available in other states, including New York, Connecticut, Maryland, Massachusetts, and Wisconsin [53].

1315

UNITED STATES FEDERAL AND STATE STEM CELL POLICY

TABLE 75.1

Human Embryonic Stem Cell Research Laws and Funding, by State

State Alabama Alaska Arizona Arkansas California Colorado Connecticut Delaware Florida Georgia Hawaii Idaho Illinois Indiana Iowa Kansas Kentucky Louisiana Maine Maryland Massachusetts Michigan Minnesota Mississippi Missouri Montana Nebraska Nevada New Hampshire New Jersey New Mexico New York North Carolina North Dakota Ohio Oklahoma Oregon Pennsylvania Rhode Island South Carolina South Dakota Tennessee Texas Utah Vermont Virginia Washington West Virginia Wisconsin Wyoming

hESc Research Legal

Funding for hESc or cloning research

Derivation from Excess IVF Embryos

Cloning (SCNT) Legal Therapeutic / Research

Reproductive

Cloning: R and T

a

Cloning: R

*

Program Ended Cloning: R and T

*

Cloning: R and T

Cloning: R and T b

Cloning: R hESc Derivation, Cloningc d

Program Ended

e

f

hESc, Human embryonic stem cell; IVF, in vitro fertilization; SCNT, somatic cell nuclear transfer.

Key: Green: Permitted by Law Red: Prohibited by Law Grey: Not considered in law - implicitly permissible, not governed by regulations or guidance * Adult and fetal stem cell research are permitted, embryonic is not

Continued

1316 TABLE 75.1

75. US STEM CELL RESEARCH POLICY

Human Embryonic Stem Cell Research Laws and Funding by Statedcont’d

State Specific Notes: a Amendments both allocating and prohibiting state funding for stem cell research have been introduced in Florida, neither have passed into law. b Minnesota law prohibits use of living human conceptus (any human organism from fertilization through the first 265 days thereafter) for any type of scientific, laboratory research or other experimentation. It is unclear whether embryonic stem cells fall under this definition: while the publically funded University of Minnesota Stem Cell Institute is permitted to use embryonic stem cell lines (from embryos created for IVF and not necessary for reproduction), no grants for hESc research have been issued in the state. c Nebraska limits embryonic stem cell research with state funds; restrictions apply to state healthcare funds provided by tobacco settlement dollars. No state facilities or funds can be used to destroy or create an embryo for the purpose of research (no derivation of stem cells or SCNT). Thus, derivation of hESc and use of SCNT can be conducted with public funds only with cell lines developed elsewhere. Both hESc derivation and SCNT are permitted using private funds. d A 2005 veto indirectly allows Nevada state funds to be used for research on human embryonic stem cell lines that existed before August 2001 (in accordance with the Bush Policy). e In 2015, the Oklahoma House passed a bill criminalizing embryonic stem cell research. The bill did not progress through the Senate and was not passed into law. State funding is allocated only to non-embryonic stem cell research, but funding is not explicitly forbidden by law. f Like Minnesota, Pennsylvania law prohibits “knowingly performing nontherapeutic experimentation upon any unborn child [.from fertilization until birth].” However, it does not explicitly ban embryonic stem cell research, which has been conducted at the University of Pennsylvania. References: Embryonic and Fetal Research Laws. National Conference of State Legislatures. 2016. Available from: http://www.ncsl.org/research/health/embryonic-and-fetal-researchlaws.aspx Georgia Institute of Technology. State Funding Boosts Stem Cell research in California, Other States. Cell Stem Cell. 2015. Available from: http://phys.org/news/201502-state-funding-boosts-stem-cell.html Johnson, JA., Williams, ED. Stem Cell Research: State Initiatives. CRS Report for Congress. Available from: https://stemcells.nih.gov/staticresources/research/GW-StateFunding.pdf Laws Related to Human Cloning. Americans United for Life. 2012. Available from: http://www.aul.org/wp-content/uploads/2012/04/bioethics-maps.pdf Protection of Human Life Act of 2015. (Oklahoma) Available from: https://legiscan.com/OK/bill/HB1379/2016 State Laws on Human Cloning. The New Atlantis. 2015. Available from: http://www.thenewatlantis.com/docLib/20150825_TNA46Appendix.pdf Stem Cell Research: A Science, American Stem Cell Research, State Cloning Legislation. OpenStax College. Available from: https://www.quizover.com/course/section/ states-with-bans-on-reproductive-and-therapeutic-cloning-by-openstax Stem Cell Research at the Crossroads of Religion and Politics. Pew Research Center. 2008. Available from: http://www.pewforum.org/2008/07/17/stem-cell-research-atthe-crossroads-of-religion-and-politics/. U.S. Stem Cell Policy Map Info. The New York Stem Cell Foundation. Available from: https://nyscf.org/scmapus Vestal, C. States Applaud New Stem Cell Funding. Pew Charitable Trusts. 2009. Available from: http://www.pewtrusts.org/en/research-and-analysis/blogs/stateline/ 2009/03/11/states-applaud-new-stem-cell-funding.

Thus far, however, the largest state initiative has been in California [54]. In November 2004, voters in California supported a proposition to allocate $3 billion over 10 years to embryonic stem cell research. The initiative, known as Proposition 71, authorized the state of California to sell $3 billion in general obligation bonds to provide funding for stem cell research and research facilities in California. Under the proposition, the funds have been distributed as grants and loans to California-based institutions by the Californian Institute for Regenerative Medicine (CIRM, established 2006), which also established regulatory standards for the research [55]. Critics of the initiative called it fiscally irresponsible given the state’s economy and other health and research needs, and argued that the institute as structured lacked accountability [55]. Nevertheless, 59% of voters approved the measure. In 2015, California launched CIRM 2.0, described as a “radical overhaul” of the agency’s operations, systems, and programs with an emphasis on speed, partnerships, and patient preference. In 2015, the state provided $135 million to 47 grants and 252 programs supporting stem cell research [56].

STEM CELL RESEARCH GUIDELINES The National Academies of Science The very limited federally funded hESc research program combined with increased interest in hESc research by states and private funders created a situation in which research was being funded by a variety of funders around the United States without the benefit of national rules or guidelines. In 2004, issues surrounding creation, use, and donation of hEScs were taken up by a panel convened by the US National Academies of Science (NAS), a private, nonprofit organization whose mission is to advise the nation on issues in science, engineering, and medicine. While they do not carry the force of the law, in general the National Academies’ reports and recommendations are very influential and its Guidelines for Human Embryonic Stem Cell Research [57] received significant media attention when they were first released in April 2005. The 2005 guidelines (and their 2007 and 2008 revisions) were instrumental templates for state legislatures and institutions as they developed their own funding and research policies.

STEM CELL RESEARCH GUIDELINES

1317

They were adopted, at least in part, by most major US research institutions [58]. While the guidelines are not the first document to speak to the conduct of embryonic stem cell research in the United States, at the time they were formulated it was not clear whether any of the previous guidance applied to contemporary hESc research because previous guidance was formulated under President Clinton but never implemented and effectively revoked by President Bush, or was simply too restrictive to meaningfully guide institutions and researchers that had made the decision to move forward with the research [59]. The guidelines were therefore received as filling a policy vacuum. President Obama’s 2009 Executive Order and the subsequent guidelines released by the NIH overrode the NAS guidelines to a certain extent. In 2010, the National Academies released a significantly amended set of guidelines, intended to be interpreted in tandem with the NIH’s regulations [58]. The committee noted that non-NIH guidelines (such as its own) would continue to be important in the areas not eligible for federal funding, and thus not governed by the NIH. Some of these areas include the derivation of stem cell lines (ineligible for federal funding under DickeyeWicker), research on hESc lines that were derived from embryos produced from some source other than IVF for reproductive purposes (including IVF for research purposes and SCNT cloning methods), and the wider range of experiments on chimeras. In the areas in which the NIH guidelines do overlap, NAS recommended that the NIH guidelines supersede its own. Where research falls outside of the scope of the federal guidelines, NAS generally recommends that the research is “permissible only after additional review and approval.” This research includes generation of hEScs by whatever means and research involving the introduction of hEScs into animals other than humans and primates at any stage of development. NAS continued to recommend that research involving in vitro culture of any intact human embryo for longer than 14 days, and the introduction of hEScs into nonhuman primate blastocysts or non-hEScs into human blastocysts should not be permitted at this time. The committee that drafted the original NAS guidelines was asked to consider the use and derivation of stem cells from embryos originally created during fertility treatment, embryos created using donated eggs and sperm, and cloned embryos. It therefore did not engage in the debate over whether it is morally permissible to destroy human embryos in research, as is required for derivation of embryonic stem cells. It also did not consider whether there is a moral difference between research that uses spare embryos and research that uses embryos created specifically for research purposes, including by cloning. The acceptability of these sources of stem cells was assumed. Overall, the NAS committee recommended banning very little scientific activity. Instead, it recommended institutional review of protocols, oversight of the involvement of egg, sperm, and embryo donors, establishment of stem cell banks, and documentation of research activity. Two recommendations were particularly significant. First, the committee recommended that much embryonic stem cell research be subject to a mixture of local and national oversight. Local oversight would occur at each institution engaged in embryonic stem cell research, which would establish an Embryonic Stem Cell Research Oversight (ESCRO) committee to oversee all issues related to the derivation and use of embryonic stem cells, review all proposals for scientific merit, maintain records of research that takes place at the institution, including registries of new cell lines, and educate investigators. As noted by the 2010 amendment, many stem cell institutions created ESCRO committees. While the 2009 NIH guidelines did not require such a board, the National Academies maintains the importance of ESCRO committees in consulting and training roles at all institutions conducting stem cell research. ESCRO committees remain especially important at centers conducting nonfederally funded hESc research, which is not subject to the same federal oversight or regulation. In many cases, local IRBs provide additional oversight. Even though much embryonic stem cell research does not strictly speaking need to go before an IRB, the NAS committee recommended that the procurement of egg, sperm, and embryos should always be reviewed by an IRB, regardless of the applicability of federal regulations, and that IRBs never waive the requirement for informed consent from a person donating cells, eggs, sperm, or embryos to research, even where the federal human subject research regulations provide for such a waiver. The other significant set of recommendations in the 2005 NAS guidelines addressed the involvement of egg, sperm, and embryo donors. In line with much guidance, law, and regulation around the world, the NAS guidelines recommended requiring consent from embryo donors for research use of those embryos. They went beyond previous US guidance, however, by extending this requirement to egg and sperm donors, including when the embryo was originally created for fertility purposes. At the time of the recommendation, gamete donors were asked to consent to reproductive use of their gametes, but were not generally asked to consent also to subsequent research use of any embryos their gametes were used to produce [60]. This NAS committee noted that this requirement might rule out the use of some embryos already created for fertility purposes that are now in frozen storage. The most recent iteration of the guidelines specifies that “written agreement at the time of gamete donation that one potential use of the blastocysts and/or morulae is embryo research will constitute sufficient consent,” meaning that consent specifically to hESc research is not needed [58]. On the issue of compensating egg, sperm, and embryo donors, the NAS guidelines noted the arguments in favor of compensation: paying egg and sperm donors is routine in the US fertility context, and many Americans

1318

75. US STEM CELL RESEARCH POLICY

participating in other kinds of research are offered financial inducements to secure their participation. They acknowledged that arguments for compensating egg donors are particularly strong: “the invasiveness and risks of the procedure suggest that financial remuneration is most deserved, but at the same time there is a greater likelihood of enticing potential donors to do something that poses some risk to themselves” [58]. These arguments notwithstanding, the National Academies committee followed previous US guidelines and guidelines and laws from many other nations in recommending that egg donors be reimbursed only for “direct expenses,” and that no payment whatsoever be offered to sperm or embryo donors. They did allow, however, for reimbursement of fertility clinics for costs, including professional services, associated with obtaining consent and collecting eggs, sperm, or embryos. In addition to cash payment, the NAS guidelines recommended against compensation in kind. Donors are not to receive any benefit from their donation, including “personal medical benefit” (excepting autologous transplantation, where the donor receives stem cells derived from his or her donation). This rule would prevent a kind of egg- or embryo-sharing arrangement whereby women or couples receive cheaper or free fertility treatment in exchange for donating a portion of their eggs or embryos to stem cell research. Similar arrangements exist in the fertility context, where women or couples receive a discount if they donate some of their eggs to others undergoing fertility treatment. This arrangement would help make fertility treatment available at a lower cost, but it would also more quickly exhaust the woman’s or couple’s supply of eggs or embryos, thereby possibly reducing their chances of achieving pregnancy [61]. In 2008, NAS’s Human Embryonic Stem Cell Research Advisory Committee revisited the issue of donor compensation. The committee retained its prohibition on compensating embryo donors and discussed the arguments for and against compensating gamete donors. The guidelines were amended to make explicit that “actual lost wages” qualify as direct expenses for which gamete donors may be compensated, but additional payments are still not permitted [62].

International Society for Stem Cell Research In 2006, the International Society for Stem Cell Research (ISSCR) released its first set of guidelines for embryonic stem cell research [63]. ISSCR is the world’s largest professional organization for stem cell scientists, and is not tied to specific countries or national policy agendas. The guidelines were revised and expanded in 2016 [64]. As can be seen in Tables 75.2 and 75.3, in many ways the updated ISSCR guidelines are very similar to those of the National Academies. Like NAS’s recommendation to form institutional oversight committees, ISSCR recommends formation of Embryo Research Oversight committees. However, under the ISSCR guidelines, these committees could operate at not only an institutional level, but could also provide local, regional, national, and international oversight. The ISSCR guidelines recommend similar standards for informed consent and the reimbursement and payment for embryos and other biological materials, and, notably, agree with the NAS that informed consent must be collected from all gamete donors. The scope of technologies and research uses covered by the ISSCR guidelines is, like in the NAS guidelines, quite broad. Both sets of guidelines address (and by extension apply to) many more kinds of research than are covered by the NIH guidelines, including derivation of cells from research embryos. Nonetheless, ISSCR takes at times a more conservative approach to novel technologies, research types, and sources of embryonic stem cell lines than the NAS guidelines. For instance, their guidelines propose a more robust review for the derivation of new stem cell lines and for the use of cells derived from research embryos [64]. Likely because they are at the time of writing the most recently published set of guidelines, the ISSCR’s recommendations address the broadest scope of technologies, including gene editing of nuclear genomes of human sperm, eggs, or embryos, mitochondrial replacement therapy, and the creation and use of human totipotent cells [64]. The potential of these technologies has expanded considerably in the years since the publication of the NIH guidelines and NAS’s final 2010 report. While not specific to the American policy context, the ISSCR’s recommendations provide critical insights into many developing uses of embryos and embryonic stem cell lines.

INTERNATIONAL COMPARISONS Biomedical research is an international undertaking and stem cell research is no exception. The governments of nations active in embryonic stem cell research employ a variety of regulation strategies, most of which include some restrictions on and oversight of the research while still allowing it to move forward. No single regulatory approach has prevailed internationally, although some patterns have emerged. For instance, many nations use national legislation to regulate stem cell research, often requiring oversight by a national stem cell research committee

1319

INTERNATIONAL COMPARISONS

TABLE 75.2

Federal and Professional Guidelines: Scope of Guideline Application NIH1

Research Using Previously Derived hEScs: Created using IVF for reproduction; no longer clinically necessary* Created by Somatic Cell Nuclear Transfer (SCNT) Created by parthenogenesis Created using IVF specifically for research purposes Generation and Derivation of new hES cell lines Use of hEScs or induced pluripotent stem cells to create chimeras: The introduction of hES or iPS cells into non-human primates at any stage of embryonic, fetal or postnatal development The introduction of hES or iPS cells into animals other than humans or primates at any stage of embryonic, fetal or postnatal development The introduction of hES or iPS cells into non-human primate blastocysts or of any embryonic stem cells into human blastocysts Allowing any animal to breed that has hES or iPS cells that could contribute to the germ line Research using SNCT for reproductive cloning Research involving human totipotent cells Research involving in vitro culture of any intact human embryo for longer than fourteen days

NAS

ISSCR

2

2

* Eligible for federal funding.

hEScs, Human embryonic stem cells; iPScs, induced pluripotent stem cells; ISSCR, International Society for Stem Cell Research; IVF, in vitro fertilization; NAS, National Academies of Science; NIH, National Institutes of Health. Key: Green: Permitted after currently mandated reviews; all research under this designation must follow all accompanying guidelines pertaining to consent, reimbursement, banking, etc. Yellow: Permitted only after additional review and approval. Red: Not permitted by guidelines at this time. In the case of NIH, research pertaining to these areas is not eligible for federal funding. The other two guidelines are not legally binding, but institutions, states, and private funders are advised to consider guidelines when considering which projects to fund and approve. Gray: Not considered by guidance. Notes: a The NIH guidelines speak and apply to only that research which is federally funded. Thus areas of stem cell research that are not addressed in the NIH guidelines are neither eligible for federal funding nor controlled by the NIH’s guidance. b The NIH 2009 guidelines prohibited funding on the creation of humaneanimal chimeras in which human stem cells are introduced into nonhuman primate blastocysts. In 2015, the NIH created a moratorium on funding for the integration of human stem cells in all nonhuman vertebrate animals’ pregastrulation-stage embryos. The NIH issued a call for comments in 2016 to consider ending the moratorium and consider extending the scope of chimera research. For more, see the AnimaleHuman Chimeras” section at the end of this chapter.

TABLE 75.3

Federal and Professional Guidelines: Research Guidelines NIH

NAS

ISSCR

INFORMED CONSENT Donors must understand all options for embryos no longer needed for reproductive purposes

x

Decisions related to creation of human embryos for reproductive purposes must have been made free from influence of researchers proposing to derive or utilize hEScs in research

x

x

x

Donors must provide consent for embryo donation to research at the time of donation

x

x

x

Donors must provide consent for tissue and cell donation if at any time these materials are to be used for research involving the creation of human embryos

x

All gamete donors must provide consent for embryo donation During the consent process, donors must be informed: 1. That embryos would be used to derive hEScs and may be kept for many years 2. What would happen to embryos in the derivation of stem cells for research 3. That the research is not intended to provide direct medical benefit to the donor 4. That the results of research may have commercial potential, and the donor(s) would not receive financial or other benefits 5. Whether information that could identify the donor(s) would be available to researchers

x x

x

x

Continued

1320 TABLE 75.3

75. US STEM CELL RESEARCH POLICY

Federal and Professional Guidelines: Research Guidelinesdcont’d NIH

During the consent process, donors must be given: 1. The information that hEScs or iPScs may be used for research on transplantation, genetic manipulation, or the mixing of human and nonhuman cells in animal models 2. A summary of the risks involved to the donor 3. A statement as to whether donors wish to be contacted in the future to receive information obtained through studies of the cell lines, if identities are to be retained Donors must be informed that they can withdraw consent at any time until stem cells are derived

NAS

ISSCR

x

x

Donors must be informed of institutional policies on return of incidental findings

x

REIMBURSEMENT AND PAYMENT No payments, cash or in kind, can be offered for donated embryos

x

x

Research oversight bodies must authorize proposals to reimburse or compensate donors Compensation for egg donors should cover “direct costs” of time, effort, and inconvenience only

x x

x

x

Mechanisms should exist on institutional, regional, and/or national level(s) to oversee and approve novel technologies, methods, or cell types

x

x

IRBs should review the procurement of all gametes, morulae, blastocysts, or somatic cells for the purpose of generating new hESc or iPSc lines

x

IRBs should approve any processes pertaining to banking of hESc lines

x

OVERSIGHT

DERIVATION, BANKING, AND DISTRIBUTION Institutions engaged in obtaining and storing hESc lines should develop standards for: 1. Committee review and oversight of banking and withdrawals 2. Documentation requirements (donor consent and IRB approval forms, medical, clinical, and diagnostic information, infectious disease screening, culture conditions, and cell line characterization) 3. A secure system for privacy of donors, especially with identifiable information National and international repositories should accept deposits of newly derived stem cell lines and distribute them on an international scale

x

x

hESc, Human embryonic stem cell; iPSc, induced pluripotent stem cell; IRB, Institutional Review Board; ISSCR, International Society for Stem Cell Research; NAS, US National Academies of Science; NIH, National Institutes of Health.

or licensing authority. Substantively, bans on creating embryos by cloning are common, although not universal, as are bans on creating embryos by fertilization except as part of fertility treatment, which means that research in many, although certainly not all, nations is limited to surplus or spare embryos. The US approach of regulating federally funded research but leaving nonfederally funded research unregulated is highly unusual by international standards [65]. In the United Kingdom, comprehensive legislation governing all research and medical use of human gametes and human embryos has existed since 1990 [66]. A major feature of the legislation is that it institutes few bans, but requires that all collection and use of embryos and gametes be licensed and overseen by an independent body called the Human Fertilisation and Embryology Authority (HFEA). The existence of this legislation and the HFEA before hEScs were first isolated meant that the research already had a regulatory system into which it could immediately be slotted, obviating the need for significant new legislative action by the UK government in response to the research (although regulations were promulgated in 2001 to add three new purposes for which research on embryos is permitted, including increasing knowledge about serious disease) [67]. Some scholars have urged the United States to adopt a similar regulatory system, under which very few activities are banned outright but instead require a license and are subject to oversight [68]. In neighboring Canada, national funding guidelines similar to the policy formulated by President Clinton were released in 2002 by the Canadian Institutes of Health Research (CIHR) [69]. Significant terms of these guidelines include that all research using CIHR funds is subject to national oversight by the Stem Cell Oversight Committee, that all embryonic stem cell lines generated using CIHR funding must be recorded in a national electronic registry and made available to other Canadian academic researchers at a cost, and that stem cells can only be extracted from

SELECTED ETHICAL, LEGAL, SOCIAL, AND POLICY QUESTIONS OF STEM CELL RESEARCH

1321

spare embryos (embryos cannot be created for research purposes, including by cloning). In 2014, the CIHR guidelines were integrated into the 2nd Edition of Tri-Council Policy Statement: Ethical Conduct for Research Involving Humans in December 2014 [70]. The CIHR guidelines were followed in 2004 by the Assisted Human Reproduction Act, which, like the legislation in the United Kingdom, regulates much more than just embryonic stem cell research [71]. Several provisions of the Act, however, directly impact embryonic stem cell research. In particular, the Act prohibits creating a cloned embryo and creating an embryo for “any purpose other than creating a human being or improving or providing instruction in assisted reproduction procedures,” thereby limiting embryonic stem cell research in Canada to embryos originally created in the course of fertility treatment (although the legislation does not specify that such embryos be surplus to the fertility needs of the donors) [72]. The distinction that many policies draw between stem cell research involving embryos created for use in research (including by cloning) and research involving embryos originally created for use in fertility treatment also appears in Australia’s Research Involving Human Embryo’s Act 2002 and the Prohibition of Human Cloning Act 2002. Both Acts limit research to “excess ART embryos,” defined as embryos created for use in the assisted reproductive technology treatment that are now excess to the needs of the woman or couple for whom they were created [73]. The Prohibition of Human Cloning for Reproduction and the Regulation of Human Embryo Research Amendment Bill 2006 retained these conditions, but further specified laws on cloning, allowing for SCNT while prohibiting reproductive cloning [74]. Similar conditions are attached to research in other nations, including France, Denmark, Finland, Greece, the Netherlands, and Japan (a full breakdown of country policies can be seen in Table 75.4). South Korea, a country active in stem cell research, is in the unusual position of permitting the creation of cloned embryos for research into rare or incurable diseases, but prohibiting the creation of embryos for research by IVF. Under a 2004 law, revised in 2008, IVF may only be used to create embryos for reproductive purposes, which can later be donated to research if not used. Cloned embryos may only be created for use in research into the treatment of rare or incurable diseases [75]. In early 2006, cloning research led by Woo-Suk Hwang of Seoul National University was found to have included a number of questionable practices, including that some egg donors were members of the research team, some egg donors were paid to donate, and researchers accompanied some egg donors as they underwent the extraction procedure [76]. Following charges of embezzlement and falsifying results against the prominent stem cell researcher, the South Korean government halted approvals and funding for stem cell research until early 2009. Embryonic stem cell projects have been approved again, and since 2012 South Korea has reemerged as a leader in stem cell research, due both to a federal funding boost and considerable regulatory freedom [77]. In Israel, another country active in stem cell research, no legislation regulates the field, although there is a law against implanting a cloned embryo [78]. Therefore derivation of stem cells from embryos created for research use, including by cloning, is allowed (although to date there are no reports of Israeli scientists creating cloned embryos) [79]. Under Singapore’s 2004 legislation, research is permitted on spare embryos and embryos created for research, including by cloning [80]. With over 40 active stem cell research groups in the country limited by very few regulations, the country has been called “Asia’s Stem Cell Center.” In 2012, Singapore established a Bioethics Advisory Committee to oversee research, but provides very few legal or regulatory barriers [81].

SELECTED ETHICAL, LEGAL, SOCIAL, AND POLICY QUESTIONS OF STEM CELL RESEARCH As embryonic stem cell research moves forward in the United States, various ethical and policy questions arise, some of which have already been discussed. For example, should researchers create embryos in the laboratory by fertilization or cloning or should they only use spare embryos? Either way, they will need to interact with fertility clinics or with egg, sperm, or embryo donors, raising questions about how those interactions should be conducted. Should researches pay fertility clinics for procuring gametes and embryos for stem cell research? How should gamete and embryo donors be approached to donate, precisely whose consent should be required, and should the donors be compensated? Once researchers extract cell lines, are they obliged to make those lines available to other researchers? Should researchers patent new cell lines or new stem cell-related processes? If they do obtain patents, what practices should they follow in licensing those lines or processes? Should researchers be allowed to mix human cells and animal cells in the creation of chimeras or hybrids, and what could happen if they do so? Here we address several of these issues in some detail.

1322 TABLE 75.4

75. US STEM CELL RESEARCH POLICY

Human Embryonic Stem Cell Research Laws and Funding in Countries with Applicable Policies

Country Argentina Australia Austria Belgium Brazil Bulgaria Canada Chile China Columbia Costa Rica Croatia Cyprus Czech Republic Denmark Ecuador Estonia Finland France Germany Greece Hungary Iceland India Iran Ireland Israel Italy Japan Latvia Lithuania Mexico The Netherlands New Zealand Norway Poland Portugal Romania Russia Singapore Slovakia Slovenia South Africa South Korea Spain Sweden Switzerland Trinidad and Tobago Tunisia Turkey United Kingdom United States

hESc Research Legal

Imported hESC Lines

~

a

Sources of Stem Cells Excess IVF Embryos Created for Research Embryos (IVF or, if permitted SCNT)

Cloning (SCNT) Legal Therapeutic / Reproductive Research

*

b

* +

+

c

^ d e

+

f g

h

~

i

j +

k

l ^

m

hESc, Human embryonic stem cell; IVF, in vitro fertilization.

n

Key: Green: Permitted by Law or Guidance (some uses may be limited by rules or regulations) Red: Prohibited by Law (some uses may be exceptions, and allowed under specific conditions) Grey: Not considered in law e implicitly permissible, not governed by regulations or guidance

n

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TABLE 75.4

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Human Embryonic Stem Cell Research Laws and Funding in Countries With Applicable Policiesdcont’d

* The country doesn’t explicitly regulate stem cells, but one of the country’s laws allows surplus IVF embryos to be used for research purposes, suggesting that hESC research is implicitly permissible. + Banned unless medically, scientifically, and therapeutically vital, in particular for the specific embryo involved. w The country has no legislation on stem cell research, but the creation of embryonic stem cell lines and research on such lines is de facto prohibited. ^ Researchers must provide explicit justification for research or use of embryos/ESCs and establish that creation / use of embryos is essential for research. Country Specific Notes: a In Austria, ESC research is de facto prohibited because embryos and gametes cannot be donated for purposes other than assisted reproduction with heterosexual couples. However, the use of imported hEScs was not addressed explicitly by law and thus is implicitly permissible. While the derivation of embryonic stem cells from human embryos is currently prohibited, the Austrian Bioethics Commission has advocated for hESC derivation of surplus embryos. b The law in Chile is unspecific. A 2006 law defends the “protection of human life from the moment of conception,” and states that no human embryos can be destroyed to obtain stem cells, and that ES cell lines only can be used for “therapeutic diagnosis or scientific research”. c The creation of and research with hES cell lines is banned in Germany as a “matter of principle.” In 2008, Germany eliminated a past law that made it a criminal offence to use hES cells in research conducted outside German’s borders. Research on imported hES cell lines is allowed, but hES cells cannot be derived in Germany, and must have been imported before May 1, 2007. d ESC research has been permitted in Iran since 2002, when Supreme Leader Ayatollah Khamenei issued a ‘stem cell fatwa’ that declared that human embryo research was consistent with Shia tradition. Iranian scientists and bioethicists have since developed ethics guidelines for hES cell research. e Ireland has no legislation on embryo or stem cell research, but the Irish Medical Council has banned medical practitioners from creating embryos specifically for research. f Lithuania limits hESC and embryo research to non-interventional clinical observations only. g Mexico is currently debating an amendment that would ban all research on human embryos, thereby halting the creation of and any research on hESC lines. h In New Zealand, research can only be conducted on non-viable embryos (in contrast with viable excess IVF embryos, permitted by other countries). i Stem cell research is only allowed in Romania under official approvals. There is no regulation of IVF, research on embryos, or hESCs. j Embryonic stem cell research is permitted in Singapore under recommendations from the Bioethics Advisory Committee, which are adhered to by the scientific community. The creation of embryos through SCNT is permitted on a case-by-case basis to derive patient-specific cells. k South Korea permits SCNT as a source for hESCs (in addition to excess IVF embryos) despite the ban on production of non-reproduction embryos; SCNT can be used “for the purpose of conducting research aimed at curing rare or currently incurable diseases.” l Sweden permits therapeutic cloning if hereditary genetic traits remain unchanged. m Turkey allows for non-embryonic stem cell research when officially approved, no hESC research is permitted. n United States law and guidelines allow for SCNT and the derivation of hESCs at large, but neither are permitted as part of federally funded research. Some individual US states have laws addressing SCNT and hESC research. References: George RP, Landry DW, Co-Chairmen C. The stem cell debates: lessons for science and politics, Appendix E: Overview of International Human Embryonic Stem Cell Laws. The New Atlantis. 2012. Available from: http://www.thenewatlantis.com/publications/appendix-e-overview-of-international-humanembryonic-stem-cell-laws Human Stem Cell Research and Regenerative Medicine: Focus on European policy and scientific contributions. European Science Foundation.2013. Available from: http://archives.esf.org/fileadmin/Public_documents/Publications/HumanStemCellResearch.pdf Jones, D. G. Where does New Zealand stand on permitting research on human embryos? The New Zealand Medical Journal. 2014;127(1399):74-82. Mahalatchimy, A. Regulation of stem cell research in Europe. EuroStemCell. Available from: https://www.eurostemcell.org/regulation-stem-cell-research-europe National Legislation Concerning Human Reproductive and Therapeutic Cloning. United Nations Educational, Scientific and Cultural Organization. 2004. Available from: http://unesdoc.unesco.org/images/0013/001342/134277e.pdf Palma, V. et al. Stem Cell Research in Latin America: Update, Challenges, and Opportunities in a Priority Research Area. Regenerative Medicine. 2015; 10(6): 785-98. Reardon, S. Mexico proposal to ban human-embryo research would stifle science. Nature. 2016; 540: 180-181. Saniei, M. Human embryonic stem cell science and policy: The case of Iran. Social Science & Medicine. 2013; 98:345-350. Sithole, S. Stem Cell Research e The Regulatory Framework in South Africa. The South African Journal of Bioethics & Law. 2011; 4(2): 1-7. Turkmen, H. O. & Arda B. Ethical and legal aspects of stem cell practices in Turkey: where are we? Journal of Medical Ethics. 2008; 34(12): 833-37. Walters, L. Human Embryonic Stem Cell Research: An Intercultural Perspective. Kennedy Institute of Ethics Journal. 2004;14(1): 3-38. Wheat, K, Matthews, K. World Human Cloning Policies. Rice University. Available from: http://www.ruf.rice.edu/wneal/stemcell/World.pdf

Compensating Egg Donors The South Korean stem cell controversy over the procuring of eggs in South Korea for use in cloning research brought additional attention to egg donation. Findings related to the donation process raised concerns about whether the women who donated eggs to Hwang’s research did so completely voluntarily. Voluntariness is a core commitment of modern research ethics [82], which generally translates into requirements that no one is pressured to participate in research and that each participant is able to withdraw from the research at any time without endangering ongoing medical care [82]. Researchers usually avoid enrolling family members and employees because they might reasonably feel significant pressure to participate. The commitment to voluntary participation, and specifically the derived right to withdraw from the research at any time, could also be under threat if researchers physically accompany volunteers through procedures as Hwang and his colleagues apparently did.

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Concern to protect voluntariness is also a major factor motivating bans on compensating egg donors, the argument being that the need for money could compel participation, especially, though not exclusively, among the poor. Nevertheless, it is unclear whether the concern is sufficient to justify a complete ban on compensation, particularly in a US context. As the National Academies committee noted, many other research participants in the United States receive compensation in exchange for their involvement in research [57]. Such payment, particularly where modest, is said to be not only a necessary incentive, but also fair treatment of research subjects (after all, the researchers and their staff will be paid for the time and resources they contribute toward the research). The concern about voluntariness in the hESc context is likely heightened because egg donation is time-consuming, painful, and involves some risk. Egg donors are injected with drugs over weeks so that they superovulate (produce many eggs). The eggs are then removed from the woman by either inserting a hollow needle through her vagina or by laparoscopic surgery. Risks of the stimulation and egg-collection process include hot flashes, headaches, sleeplessness, mood alteration, ovarian hyperstimulation syndrome, nausea, vomiting, pain, bleeding, and infection. There is even a controversy over a possible danger of ovarian cancer from the medications and ovarian stimulation for IVF [83,84]. Could payment encourage some women to donate eggs even though donation might be painful and pose a risk to their health? The answer is yes. However, whether the risk and pain are unacceptable is another issue. In theory at least, if an IRB approves research involving egg donation, it has decided that the risk to donors is reasonable in relation to the importance of the knowledge that may reasonably be expected to result [85]. Compensation is not supposed to be offered to research participants to seduce them to take an unacceptable risk. But voluntariness may not be the only concern about paying egg donors in stem cell research. There is also some opposition to paying anyone for providing bodily materials (rather than solely for their time and effort), for example, as expressed in a federal law prohibiting payment for organ donation (although that law expressly does not apply to blood, sperm, or human eggs) [86]. The stance against the sale of bodily materials is well defended in scholarly circles. Bioethicist Thomas Murray argued nearly 30 years ago that all donations of body parts, whether for research or for clinical treatment, should be gifts and not sales [87]. Others, however, including law professor Lori Andrews, writing around the same time, countered that individuals “have the autonomy to treat their own (body) parts as property,” particularly those parts of the body that they can regenerate [88]. The debate about compensating egg, sperm, and embryo donors continues today. Beyond clarifying their definition of “direct reimbursable expenses” as “costs associated with travel, housing, child care, medical care, health insurance, and actual lost wages,” the NAS guidelines included a note on autonomous choice: To facilitate autonomous choice, decisions related to the creation of embryos for infertility treatment should be free of the influence of investigators who propose to derive or use hES cells in research. Whenever it is practicable, the attending physician responsible for the infertility treatment and the investigator deriving or proposing to use hES cells should not be the same person.

Although included in the final iteration of the guidelines, the recommendations leave the issue open for interpretation by researchers, and further questioning by the scientific and policy communities at large [58]. The only state that provides for compensation of egg donors is New York, which in 2009 announced that its funds could be used to offer women up to $10,000 for donation of eggs to hESc research.1

Commercialization and Access to Treatments Another argument against paid donation is that it adds to the costs of conducting research and thereby the price of eventual treatments. But this argument works best if the same spirit is adopted by the scientists, institutions, and companies involved in the research and in any eventual treatment. Indeed, a commitment to scientific progress and widely available treatments in stem cell research might entail a commitment by all those involved to, for example, banking and widely distributing new cell lines, participating in international collaboration, and adopting patenting and licensing practices designed to facilitate access (including possibly not patenting some discoveries at all) [89]. These concerns about secrecy, privacy, access, and commercialization in stem cell research mirror a larger debate in biomedical research in general [90].

1

The legal status of compensation limits like that announced by New York’s Empire State Stem Cell Board is currently unclear in light of a recent class action lawsuit filed by egg donors in opposition to fertility industry caps on payments (Kamakahi et al. v. American Society for Reproductive Medicine et al., Case No. 3:11-cv-01781-JCS, in the US District Court for the Northern District of California).

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In terms of access, particular emphasis is often put on patenting and licensing practices, which can have an enormous impact on progress made by other researchers, as well as on the availability of eventual treatments [91]. Indeed, a number of patents already attach to embryonic stem cell research [92], including a long-standing patent claiming a purified preparation of primate (including human) embryonic stem cells and a method for isolating them [93]. To establish its 2001 registry, the NIH negotiated with patent holders to issue licenses to noncommercial entities royalty free so that research could move forward. The NIH-negotiated licenses, however, explicitly cover only research and do not extend to commercial activities [94]. It is not yet known whether these same terms apply to research using lines in the new NIH Human Embryonic Stem Cell Registry. As a general matter, patents in biomedical research can restrict other researchers from using a specific method or material, creating monopoly-like conditions, and hampering the progress of research. These conditions could significantly raise prices and reduce equitable access to treatments. Access for research and treatment was among the issues motivating the lawsuit over gene patents held by Myriad Genetics that was resolved in the Supreme Court in 2013. While the reasoning in the Myriad case refers specifically to the ineligibility of DNA sequences for patents, some have predicted that the case’s vague precedent for “natural” biologic materials could impact stem cell patents [95,96]. If it does not, the patentability of stem cell research methods and future treatments could raise treatment prices and create significant disparities in treatment access. At the time of writing, the future of the Affordable Care Act and other insurance and drug pricing regulations is uncertain in the United States. Without suitable insurance and treatment cost protections, disparities in access to stem cell treatments may be even more pronounced. Patenting and licensing issues were to some extent anticipated in California’s Proposition 71, which included a provision requiring the establishment of standards in all grants and loans allowing the state to financially benefit from licenses, patents, and royalties and resulting from the research activities funded under the measure [55]. Regulations implementing these standards require that a portion of profits be returned to the state [97].

AnimaleHuman Chimeras To develop and test possible human clinical uses, human stem cells are being implanted into nonhuman animals at various stages of development. “Humaneanimal chimeras” are created by introducing human genetic material into nonhuman animal embryos. Human stem cells can be inserted into animal embryos, fetuses or postnatally, and will grow and develop into various organs and systems within the animal. The ethical and legal questions surrounding chimera animals depend on where and when human stem cells are introduced into the animal system: is the animal a primate or vertebrate that is genetically similar to humans; are stem cells inserted into the blastocyst, embryo, or later stage in the animal’s development; will the human cells impact an animal’s nervous system or enter its germ line for reproduction; and will the animal be brought to term or allowed to reproduce? Although guidelines vary as to how they address the creation of animalehuman chimera (refer to Tables 75.2 and 75.3), they are all attentive to several similar factors. First, whether or not the animal into which the human cells are introduced is a close genetic relative of humans is considered to be important both clinically and ethically. From the perspective of clinical validity, a primate or other vertebrate animal is a more useful test model than an animal with dramatically different genetic and biological structures from humans. However, from an ethical perspective, some commentators have raised concerns about the “humanization” of animals. If an animal species is a close genetic neighbor to humans, it is thought that human stem cells introduced early in development could be sufficiently integrated into the animal in ways that could be “morally humanizing”dthat is, that move the animal’s moral status toward human beings thereby creating a novel creature with a novel, and unclear, moral status [98]. This concern about moral status is also behind two other important issues about animalehuman chimeras: the location and number of the human cells within the nonhuman animal, and the stage of development during which stem cells are introduced in animals. If stem cells are introduced in a postnatal animal, they are unlikely to become integrated into systems beyond the organ or tissues at which they were targeted [98]. However, if human stem cells are transplanted into animal embryos (single celled) or blastocysts (between 50 and 100 cells), they are considerably more likely to become integrated into many if not all of the animal’s biological systems, including the animal’s central nervous or reproductive systems. Could human stem cells in the brain of a pig, for instance, “humanize” it [98]? Could human cells enter the animal’s reproductive system causing it to produce human sperm or eggs? In response to the latter concern, prohibition on breeding of any animals into which human stem cells have been transplanted is widespread [44,54,64].

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The NIH 2009 guidelines specifically stated that research introducing embryonic or iPScs into any animals at any stage of development could not be covered by federal funding [44]. In 2015, the NIH extended the federal funding ban on chimera research, announcing a moratorium on funding for research introducing human pluripotent cells into all nonhuman vertebrate animal pregastrulation embryos [99]. After a public comment and workshop period, NIH released a call for public comments in September 2016 [100] on a proposal to expand the prohibition on introducing hEScs or iPScs into nonhuman primate embryos to include preblastocyst-stage in addition, to blastocyststage embryos. Within its proposed changes, the NIH proposed to potentially permit (as decided by the steering committee) funding for research integrating human stem cells into nonhuman vertebrate embryos up through the end of the gastrulation stage (lifting the 2015 moratorium), as well as research in which human stem cells are introduced into postgastrulation, nonhuman, nonrodent mammals, which may present a “substantial functional modification to the animal brain by human cells” [98]. This final provision, if supported by federal funding, could result in researchers being able to create animals that present the aforementioned risk of “humanization.” Such a concern has raised several questions. For instance, what addition of human materials, if anything, would constitute an “elevation of moral status” in a research animal such that it could no longer be used for research? Furthermore, a concern about humanization begs a more fundamental question about our own humanness, moral status, and consciousness. Some have argued that the interactions and impact of human material in chimeras will not reach a point at which these questions truly become necessary, and that instead, issues surrounding animal welfare and the potential harms that we may have on research animals are more important at present [98] (Currently, many research centers mandate that Institutional Animal Care and Use Committees oversee animalehuman chimera research and other stem cell transplants occurring in nonhuman animals.) Nonetheless, as stem cell technologies and research move forward, questions about research that blurs the line between humans and nonhuman animals will likely remain at the forefront of ethical and policy decision-making.

CONCLUSION As this brief overview shows, stem cell research, and in particular hESc research, has led to the development of numerous policies. In the United States, as well as in some other nations, these policies have responded to the controversial nature of research involving the destruction of human embryos. The US federal funding policy of 2001, which limited the use of federal funds in embryonic stem cell research, generated a lot of attention, including significant criticism. In response to the 2001 policy’s limits, funds for embryonic stem cell research were provided by private donors, charitable organizations, and some states. In 2009, President Obama repealed the 2001 policy. Nevertheless, federal funding, while now more widely available, is still limited to research on cells derived from spare embryos and cannot be used to derive new cell lines. As research on embryonic stem cells taken from a variety of sources moves forward in the United States and internationally, some of it without funding or guidelines, a range of policies are developing that respond to a number of important issues, including local and national oversight, the role of donors in the research, the consequences of commercial interests, the creation of animalehuman chimeras, and the extension of stem cell research technology using new technologies such as gene editing. At the time of writing, it is still unclear if and how the Trump administration will approach the ethical issues associated with embryonic stem cell research and funding. Simultaneously, stem cell and greater bioscience technologies are developing into increasingly complex arenas, including involving advances in gene editing. It will be important to continue to observe how stem cell research remains intertwined with ethical, legal, and policy questions on domestic and international fronts.

List of Acronyms and Abbreviations CIHR Canadian Institutes of Health Research CIRM California Institute for Regenerative Medicine hESc Human embryonic stem cell HFEA Human Fertilisation and Embryology Authority (United Kingdom) HHS Department of Health and Human Services, once was DHEW: Department of Health, Education and Welfare iPSc Induced pluripotent stem cell IRB Institutional Review Board ISSCR International Society for Stem Cell Research

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IVF In vitro fertilization NAS National Academy of Science NIH National Institutes of Health SCNT Somatic cell nuclear transfer

Acknowledgments Parts of this chapter are based on two of Josephine Johnston’s previous publications: Paying egg donors: exploring the arguments, Hastings Center Report 2006; 36(1): 28e31 and Stem cell protocols: the NAS guidelines are a useful start, Hastings Center Report 2005; 35(5): 16e17.

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[69] Canadian Institutes of Health Research. Human pluripotent stem cell research: guidelines for CIHR-funded research. Canadian Institutes of Health Research; 2002. Available from: www.cihr-irsc.gc.ca/e/28216.html. [70] Canadian Panel on Research Ethics. Tri-council policy statement: ethical conduct for research involving humans. Chapter 12 Section F. Government of Canada; 2014. Available from:, http://www.pre.ethics.gc.ca/eng/policy-politique/initiatives/tcps2-eptc2/Default/. [71] Canada Assisted Human Reproduction Act 2004. Canada. [72] Johnston J. Is research in Canada limited to “surplus” embryos? Health Law Rev 2006;14(3):3. [73] Research Involving Human Embryos Act 2002. Commonwealth of Australia. [74] Research Involving Human Embryos Act 2006. Commonwealth of Australia. [75] Bioethics and Safety Act No. 9100 2008 (Republic of Korea). English translation available from: http://www.mbbnet.umn.edu/scmap/ KoreanBioethics.pdf. [76] Johnston J. Paying egg donors: exploring the arguments. Hast Cent Rep 2006;36(1). [77] Park SB. South Korea steps up stem-cell work. Nature 2012;10:1038. [78] Prohibition of Genetic Intervention (Human Cloning and Genetic Manipulation of Reproductive Cells) Law, 5759-1999. Israel. [79] Walters L. Human embryonic stem cell research: an intercultural perspective. Kennedy Inst Ethics J 2004;14(1):3e38. [80] Human Cloning and Other Prohibited Practices Act 2004. Chapter 131B. Singapore. [81] George RP, Landry DW, Co-Chairmen C. The stem cell debates: lessons for science and politics, appendix E: overview of international human embryonic stem cell laws. New Atl 2012. Available from: http://www.thenewatlantis.com/publications/appendix-e-overview-ofinternational-human-embryonic-stem-cell-laws. [82] World Medical Association Declaration of Helsinki. Ethical principles for medical research involving human subjects. 2008. Available from: http://www.wma.net/en/30publications/10policies/b3/17c.pdf. [83] Gurmankin AD. Risk information provided to prospective oocyte donors in a preliminary phone call. Am J Bioeth 2001;1(4):3e13. [84] van Leeuwen FE, Klip H, Mooij TM, van de Swaluw AMG, Lambalk CB, Kortman M, et al. Risk of borderline and invasive ovarian tumours after ovarian stimulation for in vitro fertilization in a large Dutch cohort. Hum Reprod 2011;26. [85] Protection of human subjects, basic HHS policy for protection of human research subjects, criteria for IRB approval of research. 45 CFR x46.111(a). [86] Uniform Anatomical Gift Act 1987. Available from: http://www.uniformlaws.org/shared/docs/anatomical_gift/uaga_final_aug09.pdf. [87] Murray TH. Gifts of the body and the needs of strangers. Hast Cent Rep 1987;17(2):30e8. [88] Andrews LB. My body, my property. Hast Cent Rep 1986;16(5):28e38. [89] Department of Health and Human Services. Principles and guidelines for recipients of NIH research grants and contracts on obtaining and disseminating biomedical research resources: final notice. Fed Regist 1999;64(246):72090e6. [90] Krimsky S. Science in the private interest: has the lure of profits corrupted biomedical research? IEEE Technol Soc Mag 2006;25(1):10e1. [91] Heller MA, Eisenberg RS. Can patents deter innovation? The anticommons in biomedical research. Science 1998;280(5364):698e701. [92] Porter G, Denning C, Plomer A, Sinden J, Torremans P. The patentability of human embryonic stem cells in Europe. Nat Biotechnol 2006; 24(6):653e5. [93] Thomson JA. (Wisconsin Alumni Research Foundation). Primate embryonic stem cells. United States US 5843780 A; 1998. [94] Rabin S. The gatekeepers of hES cell products. Nat Biotechnol 2005;23(7):817e9. [95] Diamond NJ. Stem cells and the trajectory of section 101 jurisprudence after Myriad. Albany Law J Sci Technol 2016;45. [96] Smith S. Claiming a cell reset button: induced pluripotent stem cells and preparation methods as patentable subject matter. Boston Coll Law Rev 2015:1577. [97] California Public Health Code of Regulations 2006. 17 Cal. Code of Regs. [98] Hyun I. What’s wrong with human/nonhuman chimera research? PLoS Biol 2016;14(8):e1002535. [99] National Institutes of Health. NIH research involving introduction of human pluripotent cells into non-human vertebrate animal pregastrulation embryos. 2015. NOT-OD-15-158. Available from: http://grants.nih.gov/grants/guide/notice-files/NOT-OD-15-158.html. [100] National Institutes of Health. Request for public comment on the proposed changes to the NIH guidelines for human stem cell research and the proposed scope of an NIH Steering Committee’s consideration of certain human-animal chimera research. Fed Regist 2016;81:51921e3. Available from: https://grants.nih.gov/grants/guide/notice-files/NOT-OD-16-128.html.

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76 Ethical Considerations* Ronald M. Green Dartmouth College, Hanover, NH, United States

INTRODUCTION Since the first development of human embryonic stem cells (hESCs) in 1998 [2,3] there has been a consensus in the scientific community that pluripotent cells hold great promise for developing new treatments for a variety of serious and currently untreatable disease conditions [4,5]. Although clinical successes in this area have been slow in coming, research is ongoing and novel therapies are beginning to show promise in clinical trials [6,7]. Nevertheless, because hESC research involves the manipulation and destruction of human embryos, the field has also been a focus of ethical controversy and opposition. In the course of these debates, many challenging ethical questions have been raised. Scientists, clinicians, or patients using hESCs or therapies must formulate their own answers to these questions. Society, too, must address them to determine the extent to which hESC research may require oversight and regulation. This chapter presents the most pressing of these questions, and critically examines some of the answers that have been proposed to them.

IS IT NECESSARY TO USE HUMAN EMBRYOS? Whether it is permissible to destroy human embryos to create hESC lines is a leading question, to which we turn in question 2 later. However, in late 2007, teams led by Yamanaka in Japan and Thomson in the United States announced success in the use of gene transfer technology to produce induced pluripotent stem cell (iPSC) lines [8,9]. Some have argued that this development obviates the need to develop or use hESC lines [10]. They assume that if we can directly manipulate somatic cells and perhaps even produce patient-specific (autologous) stem cells there is no reason to destroy human embryos or use hESCs derived from such destruction. However, there are both scientific and ethical reasons for questioning this assumption. Scientifically, there is a question of whether iPSCs will prove suitable for use in human transplant and cell regeneration therapies. Early iPSCs exhibited high rates of tumorigenicity in mice, possibly a result of the use of retroviral vectors to carry pluripotency-inducing transcription factors, including the cancer-related factor c-Myc [11,12]. Reports suggest that the genomic reprogramming process in iPSCs may be less complete than that which takes place in the fertilized egg, causing the resulting cells to exhibit abnormal expansion and early senescence [13]. There is also a concern that because iPSCs are produced from somatic cells that have been exposed to aging and toxins, they may exhibit harmful mutations when used in clinical therapies. Because of these concerns and the possible costs of patient-specific therapies, some argue that banks of tissue-matched hESCs offer better prospects for therapies. In a review of these issues, Hug and Hermere´n conclude that “if we consider all the aspects of safety, it is hardly possible to determine which therapy based on which type of cells would be safer according to the present state of knowledge” [14]. Research to answer these questions is ongoing (for example, [15,16]), but doing so will require a better understanding of the reprogramming process in embryonic cells. For this reason alone, the use of human embryos is likely * This chapter is a substantial revision and updating of the chapter “Ethical Considerations” that appeared in Atala et al. [1].

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to continue to be an important feature of stem cell research [17]. At this time, there exist many useful hESC lines created from donated human embryos. There are also currently hundreds of thousands of frozen embryos remaining from infertility procedures that will likely be destroyed and that could be used for research [18]. If stem cell research involving embryos were halted, this vast resource would go to waste. Many people feel that it is unwise to foreclose any of the available paths to developing pluripotent cells for regenerative medicine research. The view of the National Research Council remains in force: “The application of stem cell research to therapies for human disease will require much more knowledge about the biological properties of all types of stem cells” [4]. Ethically, it can also be asked whether the production and use of iPSCs is free of concerns. It is widely assumed that iPSCs avoid ethical controversy because they are produced from somatic cells, which unlike early embryos have no inherent potentiality of developing into a human being. But the work of Nagy and others shows that embryonic stem cells, when inserted into tetraploid embryos, are able to develop the placental material needed for further development [19e24]. This possibility raises complex ethical questions. Is potentiality morally relevant if it is accompanied by such intensive technical interventions? Embryos produced by in vitro fertilization (IVF) or nuclear transfer (cloning) also require intensive technical interventions, and opponents of hESC research normally condemn the use of these as sources for hESCs. If it is acknowledged that the potentiality for full human development, however assisted, confers moral status on an entity, then iPSC research may not be the ethical panacea that many of its proponents believe it to be. As Giuseppe Testa [25] observes: The point is not that one can no longer use the potentiality argument because fibroblasts might now be considered as potential persons in a more concrete way than ever before. That is, there is no longer, if there ever was, any ready-made grid of boundaries that biology can let us see as if they were simply out there and that can serve as neutral justification for political choices.

Finally, it might be asked whether, even within the framework of hESC research, we must destroy embryos. In 2005, the Bush administration’s President’s Council on Bioethics issued a White Paper encouraging research in alternate methods of hESC derivation, including the use of arrested or developmentally nonviable embryos [26]. Several of these proposals, such as the deliberate genetic manipulation of embryos to prevent their normal development, are not free of ethical controversy [27,28]. A method involving single-cell blastomere biopsy has also been developed and implemented that could obviate the need to destroy the embryo to develop an hESC line [29]. While this approach raises questions of risk to the manipulated embryos, it can be justified in the context of preimplantation genetic diagnosis (PGD), where cells are routinely removed as a diagnostic procedure. Reviewing all these considerations, we can say that considered scientifically and ethically there is no easy route around the destruction and use of human embryos in pluripotent stem cell research.

IS IT MORALLY PERMISSIBLE TO DESTROY A HUMAN EMBRYO? hESC lines are usually made by chemically and physically disaggregating an early, blastocyst-stage embryo, and removing its inner cell mass. At this stage the embryo is composed of approximately 200 cells, including an outer layer of differentiated placental material, and the undifferentiated (pluripotent) cells of the inner cell mass. The embryo dies as a result of this procedure. New methods that permit the development of pluripotent cell lines without destroying an embryo have not yet replaced this standard method for developing hESCs, and most existing hESC lines were created this way. Hence the question remains: may we intentionally kill a developing human being to expand scientific knowledge and potentially provide medical benefits? At one end of the spectrum of answers are those who believe that, in moral terms, human life begins at conception when a new, self-developing genome comes into being. For many of those holding this view, the early embryo is not morally different from a child or an adult human being. It cannot be used in research that is not to its benefit, and it cannot be used without its consent [30e32]. Furthermore, proxy consent by parents in such cases is inadmissible, since it is an accepted rule of pediatric research that parents may not volunteer a child for risky studies that are not to its benefit. Many Roman Catholics, evangelical Protestants, and some Orthodox Jews take the position that life (morally) begins at conception, and they oppose hESC research. At the other end of the spectrum are those who believe that the embryo is not yet fully a human being in a moral sense. They hold a “developmental” or “gradualist” view of life’s beginning [33e37]. They do not deny that the early embryo is alive and has the biological potential to become a person, but they believe that other features are needed for the full and equal protection we normally accord children and adults and that these features only develop gradually across the full term of gestation. These features include such things as bodily form and the ability to feel or

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think. Since the early embryo lacks organs, it cannot have these features or abilities. They also note that the very early embryo lacks human individuality, since it can still undergo twinning at this early stage, and two separate embryos with distinct genomes can fuse to become a single individual [38]. Some dismiss this argument, maintaining that the possibility of twinning or fusion does not reduce the genetic uniqueness or moral status of the earliest developing cells [39], but it is hard to see how the claim that “a person begins at conception” can easily withstand these biological facts. Finally, the very high mortality rate of such embryos (most never implant) reduces the force of the argument from potentiality [40,41]. Those who hold this developmental view do not agree on the classes of research that warrant the destruction of embryos, but most support some form of hESC research. Their reasoning is that, although the early embryo merits some respect as a nascent form of human life, the lives and health of children and adults outweigh whatever claim it possesses [42]. Each individual faced with involvement in hESC research must arrive at his or her own answer to the question of the status of the early human embryo and when, if ever, it may be destroyed. Legislators and others must also wrestle with these issues. On March 9, 2009, President Barack Obama issued an executive order authorizing federal funding for research involving the use, but not the derivation, of hESC lines. Whether this position will be sustained by the Trump administration remains to be seen.

MAY ONE BENEFIT FROM OTHERS’ DESTRUCTION OF EMBRYOS? If, as some maintain, the human embryo is a morally protectable entity that cannot be intentionally destroyed, can researchers, clinicians, or patients justify the downstream use of a cell line produced from its destruction? This raises the more basic question of whether we can ever benefit from deeds with which we morally disagree or regard as wrong [43,44]. It is also the question of when a connection with wrongdoing becomes complicit with it [45]. Why is it morally wrong to benefit from others’ misconduct? One answer is that by doing so we may encourage similar deeds in the future. This is most apparent in cases where our conduct directly instigates wrongdoing, such as when we authorize theft or receive stolen goods [45]. However, it seems less objectionable to benefit from others’ wrongdoing when their deeds are independently undertaken and not in any way prompted or encouraged by us. For example, few would object to using the organs from a young victim of a gang killing to save the life of another dying child. The use of organs benefits one person and in no way encourages teen violence. Can similar logic apply to stem cell research using spare embryos remaining from infertility procedures? It helps to remember that most embryos used to produce hESC lines are left over from infertility procedures. Couples using IVF routinely create more embryos than can safely be implanted. There are already hundreds of thousands of these embryos in cryogenic freezers in the United States and around the world [18]. Despite strenuous efforts, including some supported by the US government, few frozen embryos are adopted [46,47]. Most of these supernumerary embryos will be destroyed. In 1996, British law mandated the destruction of 3600 such embryos [48]. Does using hESC lines made from surplus embryos encourage either the creation or destruction of embryos? Although a downstream researcher, clinician, or patient may abhor the deeds that led to the existence of an hESC line, including the creation and destruction of excess human embryos in infertility medicine, nothing that a recipient of an hESC line chooses to do is likely to alter, prevent, or discourage the continuing creation or destruction of human embryos, or make the existing lines go away. Those who use such embryos may also believe that if they refuse to use an hESC line, they would forgo great therapeutic benefit. People in this position will struggle with the question of whether it is worthwhile to uphold a moral ideal or engage in a symbolic act when doing so has no practical effect, and when it threatens to expose others to harm. It is noteworthy that in an August 2001 address to the nation, President George W. Bush adopted a version of the position that allows one to benefit from acts one morally opposes. Stating his belief that it is morally wrong to kill a human embryo for others’ benefit, the President nevertheless permitted the use of existing stem cell lines on the grounds that the deaths of the embryos had already occurred [49]. Similarly, Germany permits the importing and use of hESC lines created before January 1, 2002, the date on which the Bundestag passed a law governing such matters (this date was later moved to May 2007). These initiatives reflect the belief that it does not encourage further destruction of embryos to permit the use of previously generated cell lines. President Bush did not go so far as to permit the use of lines that could in the future be derived from embryos certainly slated for destruction because their progenitors choose never to use them. However, in July 2009, the National Institutes of Health (NIH) implemented an executive order issued in March 2009 by President Obama that authorized funding for the use, but not the derivation, of new hESC lines from embryos remaining from infertility procedures [50]. Opponents of hESC research have criticized this derivation-versus-use distinction as morally problematical, but apart from the moral issues,

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this distinction is legally necessary because the DickeyeWicker Amendment prohibits federal funding for research in which a human embryo is destroyed [51]. In summary, the downstream use of stem cells derived from embryos remains a source of moral controversy and disagreement. Each researcher or clinician who opposes embryo destruction will have to examine his or her own conscience in the light of the foregoing considerations to determine how much they wish to associate themselves with the use of hESC lines created in ways to which they morally object.

MAY WE CREATE AN EMBRYO TO DESTROY IT? Is it ever morally permissible to create an embryo deliberately to produce a stem cell line? This was done in the summer of 2002 at the Jones Institute in Norfolk, Virginia [52]. Those in favor of this research defend it on two grounds. First, they say that in the future, if we seek to develop stem cell lines with special properties, such as closer genetic matches to tissue recipients or properties introduced by gene editing (see question 7 later), it may be necessary to produce stem cell lines to order using donor sperm and eggs. Second, they argue that it is ethically better to use an hESC line created from embryos that have been produced for just this purpose, with the full and informed consent of their donor progenitors, than to use cell lines from embryos originally created for a reproductive purpose. Those who believe the early embryo is our moral equal oppose the deliberate creation of embryos for research or clinical use. They are joined by some who do not share this view of the embryo’s status, but who believe that it is morally repellant to deliberately create a potential human being only to destroy it [53]. They argue that this research opens the way to the “instrumentalization” of all human life and the use of children or adult human beings as commodities. Some ask whether such research does not violate the Kantian principle that we should never use others as “a means only” [54]. On the other side of this debate are those who believe that the lesser moral status of the early embryo permits its creation and destruction for lifesaving research and therapies [55,56]. The proponents of this research direction ask why it is morally permissible to create supernumerary embryos in IVF procedures to help couples have children, but morally wrong to do the same thing to save a child’s life. They are not persuaded by the reply that the status of the embryo is affected by its progenitors’ intent, and that it is therefore permissible to create excess embryos for a “good” (reproductive) purpose, but not for a “bad” (research) purpose. They point out that the embryo is the same entity, and its status should not depend on its progenitors’ intentions. We do not ordinarily believe that a child’s rights are dependent on its parents’ intent or degree of concern [57]. They conclude that it is not parental intent that warrants the creation of excess of embryos in such cases, but the embryo’s lesser moral status and the likelihood of significant human benefit from its use. These same considerations, they believe, justify deliberately creating embryos for stem cell research.

MAY WE CLONE HUMAN EMBRYOS? This question arises in connection with the patient-specific stem cell technology known as “human therapeutic cloning.” It involves the deliberate creation of an embryo using somatic cell nuclear transfer (cloning) technology to produce an immunologically compatible (isogenic) hESC line [58]. Immune rejection could occur if the embryo used to prepare a line of hESCs for transplant does not share the same genome as the recipient. This would be the case whether the cell line was created from a spare embryo or from one made to order. Therapeutic cloning offers a way around this problem. In the case of a diabetic child, for example, the mother could donate an egg whose nucleus would then be removed. A cell would be taken from the child’s body and its nucleus inserted into the egg cytoplasm. With stimulation, the reconstructed cell would divide, just like a fertilized egg. If the resulting embryo were transferred back to a womb, it could go on to birth and become a new individualda clone of the child. But in therapeutic cloning, the blastocyst would be dissected and an hESC line prepared. Growth factors could be administered to induce the cells to become replacement pancreatic cells for the child. Because these cells contain the child’s own DNA, and even the same maternal mitochondrial DNA, they would not be subject to rejection. Research has also shown that the small amount of alien DNA from the mitochondria of a donor egg would probably not provoke an immune response [59]. Although this is a promising technology, it raises a host of novel ethical questions. One is whether the embryonic organism produced in this way should be regarded as a “human embryo” in the accepted sense of that term [60]. Those who believe that “life begins at conception” tend to answer this question affirmatively, even though cloned

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“embryos” are not the result of sexual fertilization [61]. They believe that it is no more permissible to create and destroy a cloned embryo than to do so with one produced by sexual fertilization. They base their view on the biological similarities between cloned and sexually produced embryos and on the argument that both have the potential to become a human being. However, as we have seen, if the embryo’s status rests on its potential, some degree of potentiality attaches today to all bodily cells, which no one would argue should be withheld from use in biomedical research or therapy. Therapeutic cloning implies the availability of human oocytes and raises the special question of ovulation induction. This is an invasive medical procedure, with both known and undetermined risks [62,63]. Not only must egg donors be informed of these risks, but steps also must be taken to preserve the voluntary nature of their consent. This includes avoiding financial incentives that create an “undue influence” on donors’ judgment. It also includes preventing them from being pressured into producing excess eggs or embryos for research in return for discounts on infertility services [64]. Fears about coercion and the exploitation of poorer women or women of color have led some to oppose paid egg donation for research [65,66]. Several states, including California and Massachusetts, have passed laws prohibiting this practice. Nevertheless, experience has shown that women will not donate eggs for either reproductive or research purposes without adequate compensation [67]. In the face of these problems, New York State reversed the legal trend and approved payment for research egg donation. Those who defend payment for oocytes point out that payment to research subjects involved in risky research is a common practice. They also ask why it is permissible to pay reproductive but not research egg donors [68]. The Ethics Committee of the American Society for Reproductive Medicine has supported payment for research egg donors. It conceptualizes this in terms of appropriate compensation for a donor’s time, inconvenience, and discomfort, but not for the eggs themselves [69,70]. Finally, there is a moral question specific to cloning itself. The more scientists are able to perfect therapeutic cloning, the more likely it is that they will sharpen the skills needed to accomplish reproductive cloning, which aims at the birth of a cloned child. There is a broad consensus in the scientific and bioethics communities that, for the foreseeable future, cloning technology poses serious health risks to any child born as a result of it [71]. There are also serious, unresolved questions about the psychological welfare of such a child [72]. Finally, there is the possibility that embryos created for therapeutic cloning research might be diverted to reproductive cloning attempts. All these concerns raise the question: do we really want to develop cloning technology for the production of isogenic stem cells if doing so hastens the advent of reproductive cloning [73]? The advent of iPSCs and the prospect of efficiently producing patient-specific stem cell lines have weakened the arguments for therapeutic cloning with its many associated controversies.

MAY WE USE HUMAN STEM CELLS TO CREATE CHIMERAS? In stem cell research, human-to-animal chimera experiments involve the transfer of pluripotent or multipotent human stem cells into animals at embryonic, fetal, or postnatal stages of development to study stem cell behavior [74]. Some forms of chimera research are common and widely acceptable. For example, hESCs are routinely used to form teratomas in immunodeficient mice to assess stem cell quality and developmental potential. The creation of humanized antibody systems in mice is also central to cancer immunotherapy research. Researchers have expressed interest in creating humanized organs or tissues in larger animals for disease modeling, drug testing, and perhaps eventual organ transplant [75]. Animal models with human cells in the brain can be used to study many human brain diseases, including Parkinson’s, Alzheimer’s, and schizophrenia, and may be useful models for testing new drugs [76]. However, these proposals raise many ethical questions because hESCs or human islet-derived precursor cells (hiPCs) inserted into an animal embryo prior to gastrulation can incorporate themselves into neural or reproductive tissues. Modifying neural tissues risks the creation of an animal with significantly increased potential for human sentience and self-awareness, a concern that is greatly increased in the case of nonhuman primates whose brain architecture might reasonably support human-like cognition and feelings. Humanizing reproductive tissues risks the inadvertent mating of two animals with such tissue in their gonads and a resulting pregnancy or birth of a human child from an animal womb. Some have argued that either prospect is ethically unacceptable because it represents a threat to human dignity [77,78], although what constitutes human dignity or its violation is difficult to assess [74]. Nevertheless, in terms of recognized human subject protections and animal welfare considerations, few would disagree that it is wrong to create animals with significantly humanized brains or to bring about a human pregnancy in an animal uterus.

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Part of the difficulty in coming to terms with these questions are the many variables involved. Research and therapeutic benefits that are often speculative must be weighed against risks, which depend in part on the nature of the animals being used, the times at which pluripotent cells are inserted into the developing organism, the likelihood of these cells integrating themselves into various animal tissues and organ niches, and whether the organism will be allowed to come to term. Compounding this complexity is the “irreducible degree of uncertainty about the cognitive nature of the new chimeric animal, and how it would manifest distress, anxiety, or other factors relevant to one’s assessment of animal welfare” [74]. These complexities counsel case-by-case assessment of research directions, protocols based upon careful monitoring of outcomes, and the best evolving information about risks and benefits [66,79e81]. The International Society for Stem Cell Research (ISSCR) summarizes this position when it calls for “specialized oversight” of such research utilizing “baseline animal data grounded in rigorous scientific knowledge or reasonable inferences” and involving “a diligent application of animal welfare principles” [82]. In August 2015, the NIH imposed a moratorium on its funding of human-to-animal chimera research pending public comment on revised guidelines. These guidelines include the creation of a high-level internal NIH steering committee to provide broad policy oversight for this research. In addition, the guidelines include specific prohibitions on (1) inserting hESCs or hiPSCs into nonhuman primate blastocyst-stage embryos, and (2) breeding of animals where the introduction of hESCs or hiPCs may contribute to a germ line (i.e., make human eggs or sperm) [76]. A 2011 report by the British Academy of Medicine also recommends a prohibition on “substantial functional modification of the NHP [normal pressure hydrocephalus] brain, such as to engender ‘human-like’ behavior,” though it offers somewhat more lenient permission, with oversight, for “substantial modification of an animal’s brain that may make the brain function potentially more ‘human-like,’ particularly in large animals” [83]. Despite disagreements by national bioethics bodies at the margins of research possibilities, there is a consensus that chimera research involving the humanization of animal brains and any research that risks a human birth requires ongoing, case-by-case, specialist scrutiny of the scientific and ethical issues involved.

MAY WE GENETICALLY MODIFY HUMAN EMBRYOS? Gene editing has recently been given new precision and applicability by development of CRISPR-Cas9 technology. It is now possible to target specific gene sequences in somatic cells, stem cells, gametes, or embryos for deletion or modification. This raises the prospect of human genetic engineering for the purpose of disease prevention and treatment and for genetic enhancement. For example, it might be possible to use genetically modified cells or stem cells to provide therapies for HIV AIDS or sickle cell anemia. Pathologies that begin at early uterine development might be prevented by the modification of parental gametes or the early embryo itself. Germline interventions like these have the added advantage of preventing future transmission of the disease-causing mutations. In many cases it is true, the same goal can be accomplished without gene editing by use of PGD and embryo selection, but gene editing permits the introduction of novel gene sequences not available in the parental lineage. It can also facilitate genetic enhancement, the creation of human beings who are “better than well” [27,28]. Despite its capabilities, CRISPR-Cas9 has deficiencies. It can alter and disrupt DNA at locations other than the intended ones (“off-target results”); it can change the DNA in some but not all, resulting in a mosaic of altered and unaltered cells, and can generate immune responses if introduced into the body. Although the CRISPR-Cas9 system is still undergoing development to reach the level of safety where it could be used in clinical applications, there is broad consensus that it would be premature and ethically questionable to attempt germline gene editing at this time, even for serious genetic diseases. In December 2015, leading scientists and bioethicists convened an international summit on human gene editing that concluded: It would be irresponsible to proceed with any clinical use of germline editing unless and until (i) the relevant safety and efficacy issues have been resolved, based on appropriate understanding and balancing of risks, potential benefits, and alternatives, and (ii) there is broad societal consensus about the appropriateness of the proposed application.

The report continues: At present, these criteria have not been met for any proposed clinical use: the safety issues have not yet been adequately explored; the cases of most compelling benefit are limited; and many nations have legislative or regulatory bans on germline modification. However, as scientific knowledge advances and societal views evolve, the clinical use of germline editing should be revisited on a regular basis.[84].

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ARE THERE SPECIAL CONSIDERATIONS GOVERNING THE USE OF STEM CELLS IN CLINICAL RESEARCH AND CLINICAL APPLICATIONS? Like all research or clinical translation involving the use of animal or human subjects, stem cell research is governed by a set of fundamental ethical principles that have been embodied in national and international codes and laws. As enunciated by the ISSCR, these principles include (1) research integrity; (2) primacy of patient welfare; (3) respect for research subjects; (4) transparency; and (5) social justice [82]. Some commentators have warned against “stem cell exceptionalism”dthe singling out of stem cell research for unwarranted special attention [74]. Thus it may be asked whether the use of stem cells in research or clinical treatments raises any novel ethical questions. Given the controversial and moral status of hESCs and the long-lasting but still uncharted powers of stem cells themselves, it seems wise to add special provisions to the guidelines governing research and clinical applications when hESCs and hiPSCs are involved. Such guidelines have been developed by the Chief Medical Officer’s Expert Group in Great Britain [85,86], by private ethics boards at the Geron Corporation [87] and Advanced Cell Technology [88] in the United States, by committees of the National Research Council and Institute of Medicine [66,80,81], by the NIH [50,89], and by the ISSCR [82,90,91]. These various guidelines tend to share several features.

Stem Cell Research Oversight/Embryo Research Oversight Committee Review There is agreement that, in addition to the usual review of research proposals by an Institutional Review Board (IRB), it is appropriate that there be another layer of review provided by a stem cell research oversight (SCRO) committee, or, as it has more recently been termed, an embryo research oversight (EMRO) committee. The name change reflects the advent of hiPCs, the use of which, except for those having “organismal potential,” raises fewer issues than the use of hESCs. In its most recent guidelines, the ISSCR recommends EMRO review for research involving human embryos, embryo-derived cells or that which entails “the production of human gametes in vitro when such gametes are tested by fertilization or used for the creation of embryos,” and research on human totipotent cells that have the potential to sustain embryonic or fetal development [82]. In contrast, use of iPSCs requires human subjects review but does not require specialized EMRO review so long as the research does not generate human embryos or totipotent cells [82]. The ISSCR also exempts from EMRO review research with established hESC lines that are confined to cell culture or involve routine and standard research practice, such as assays of in vitro differentiation or teratoma formation in immune-deficient mice. EMROs can operate at institutional or higher levels, and should include as members scientists and/or physicians not directly engaged in the research under consideration but with relevant expertise, ethicists, “members or advisors familiar with relevant local legal statutes governing the research,” and “community members, unaffiliated with the institution through employment or other remunerative relationships, who are impartial and reasonably familiar with the views and needs of research subjects, patients and patient communities who could be benefited by stem cell research, and community standards” [82].

Donor and Procurement Issues Because hESC and therapeutic cloning research require a supply of human gametes and embryos, any of which raise sensitive issues of moral status and familial relationships, steps must be taken to elicit the informed consent of donors, to protect their privacy, and to minimize any risks to which they might be subject. Because hiPCs are derived from human somatic material, existing IRB regulations concerning review and informed consent cover their derivation, and no EMRO committee review is needed. Nevertheless, the possibility of the use of such cells in chimeric animals and special transplantation issues recommend EMRO review in some cases. To aid researchers, the ISSCR has offered the outlines of a sample consent document for these purposes [82]. In general, explicit and contemporaneous informed consent (consent given at the time of procurement) is required when donors are asked to provide biomaterials for the creation of embryos, hESCs, or immortalized hiPC lines. Contemporaneous consent is not necessary if researchers procure somatic cells from a tissue bank, but in such cases the tissue bank’s consent documents must evidence that donors consented to the use of their biomaterials for gamete creation or for stem cell research. There is wide consensus that donors of cells, sperm, or embryos should not be compensated, although it is regarded as permissible to provide compensation for the time, inconvenience, and discomfort associated with the donation of oocytes for research so long as payment does not vary according to the planned use of the oocytes, the number or quality of oocytes retrieved, the number or outcome of prior donation

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cycles, or the donor’s ethnic or other personal characteristics [70]. Payment levels should not be so high as to constitute an “undue influence” on a woman’s decision-making. Informed consent requires that donors fully understand the nature of the research being undertaken (for example, that it may require the destruction of an embryo), and that they explicitly consent to it. Whenever possible, donors should be given the option of agreeing to some forms of hESC research but not others (e.g., the creation of stem cells but not cloned embryos). It is morally impermissible to elicit somatic cells, sperm, eggs, or embryos for the production of stem cell lines without informing donors that an immortalized, pluripotent cell line might result that could be widely used in research or therapeutic applications. If there are likely to be commercial benefits flowing from the research, donors must also be informed of this, and their rights (if any) in such benefits should be clearly specified. If the research involves therapeutic cloning, both egg and somatic cell donors must be informed that a cloned embryo and a cell line with the egg donor’s mitochondrial genetic material and the somatic cell donor’s nuclear DNA will result. iPSC somatic cell donors should also be informed about the developmental and commercial possibilities of their cellular materials. Special questions also arise about the use of somatic cells, embryos, or sperm whose provenance is uncertain, including embryos created with sperm whose donor is unknown [92]. An emergent problem given the possible widespread use of hESCs and iPSCs is the discovery of incidental genetic findings in the course of downstream research that might be of health importance to donors. ISSCR and other bodies recommend the development of a policy on whether and how incidental findings will be returned to donors and the communication of this policy to donors at the time of the informed consent process. Successful implementation of this policy depends on the traceability of cell lines and strict compliance with material transfer agreements [82]. In conducting research, efforts should be made to preserve donor privacy by removing identifying information from gametes, embryos, and cell lines, and keeping this information apart in a secure location. In view of the controversy surrounding much of this research, donors can be subjected to harassment or embarrassment if their association with the research is revealed without their consent. In an era where it is increasingly possible to identify donors from the DNA of the resulting stem cells, donors should be warned of possible risks to their privacy and that of their family members. Finally, there is a question about the use of hESC lines previously produced (or produced abroad) under less comprehensive guidelines. The NIH faced this question in its effort to develop guidelines following President Obama’s executive order permitting funding for hESC research. Many hESCs available for use had not been developed in precise conformity to the newly developed guidelines. The NIH met this problem by forming a Working Group of the Advisory Committee to the Director to consider exceptions that conform to the spirit of the guidelines. Rather than “grandfathering” preexisting hESCs, the Working Group is empowered to judge hESC lines “derived in a responsible manner” to be eligible for use in NIH-funded research [50].

Research Conduct Guidelines also apply to the overall conduct of research. These include the requirement that no embryo used in hESC or therapeutic cloning research be allowed to develop beyond 14 days. This internationally recognized limit is based on the substantial changes that occur at gastrulation, which marks the beginning of individualization and organogenesis [93]. At this time, ISSCR guidelines also prohibit reproductive cloning research; research in which human embryos (or organized cellular structures that might manifest human organismal potential) are gestated ex utero or in any nonhuman animal uterus, or have undergone modification of their nuclear genome and are implanted into or gestated in a human or animal uterus; and research in which animal chimeras incorporating human cells with the potential to form human gametes are bred to each other [82]. Supervision and accountability of all staff and scientists involved in this research to prevent any diversion of gametes or embryos to reproductive purposes is required.

Clinical Translation Stem cell-based treatments have been the clinical standard of care for some conditions, such as hematopoietic stem cell transplants for leukemia, but research on stem cell-based therapies has greatly expanded in recent years. Investigational New Drug applications have been filed for several therapies, and clinical trials are under way. As might be expected, problems have already developed. In late 2009, the Food and Drug Administration (FDA) asked the Geron Corporation to halt its trial on a groundbreaking stem cell therapy for spinal cord injury when a review revealed the formation of cysts in some (animal) trial subjects [94].

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All tissue transplantation involves risks for the recipient, but for several reasons these risks are magnified in the case of stem cell transplants: • Stem cells are novel therapeutic agents whose development and manufacture require innovative procedures for ensuring purity and homogeneity. Additional risks derive from the fact that cells are likely to be expanded in culture and/or exposed to xenoculture materials, viruses, or other infectious agents before transplantation. The likely use of these cells by large populations increases the need for careful oversight in cells’ development and manufacture. • Cell types may differ in their ability to proliferate and safely implant in the body. Their use poses risks of infusional toxicity, severe immune responses, and tumorigenesis. An additional problem is ectopic or off-target insertion. • Animal models may not accurately reflect toxicity in humans. In response to these challenges, in 2008 a multidisciplinary, international task force of the ISSCR proposed ethical guidelines for clinical translational research. The task force issued 39 recommendations, most of which fall into five broad categories: (1) cell processing and manufacture, (2) preclinical studies, (3) clinical research, (4) stem cell-based medical innovation, and (5) considerations of social justice [91]. With regard to cell processing and manufacture, the task force recommended that these be conducted under “scrupulous, expert, and independent review and oversight.” As a rule, minimally manipulated products (cells maintained in culture under nonproliferating conditions for short periods of time) require less oversight than those subjected to extensive manipulations, such as genetic alterations. The same is true of autologous versus allogenic use and use for homologous versus nonhomologous functions (Recommendations 8 and 9). Good manufacturing practice must be applied to manipulated products and those destined for allogeneic use. Donors of cells should be screened for infectious diseases and the donor should give written informed consent that covers the likely storage, future manipulations, analyses, uses of their cells, commercial potential, and possible risks to the donor’s privacy, including exposure of genetic information (Recommendation 3). Preclinical studies are meant to provide evidence of product safety and proof-of-principle of therapeutic effect. This normally requires sufficient studies in animal models, including larger animals where structural tissue needs to be tested in a load-bearing model (Recommendations 11 and 14). Researchers must develop and implement a clear plan to assess the risks of tumorigenicity for any cell product (Recommendation 18). Cell cultures and animal models should be used to test the interaction of cells with drugs to which recipients will be exposed, including the immunosuppressants planned for recipients (Recommendation 19). Clinical trials of stem cell research must conform to internationally accepted principles governing the protection of human subjects, including regulatory oversight, peer review by an expert panel independent of the investigators and sponsors, fair subject selection, informed consent, and patient monitoring. In addition, there are special issues raised by stem cell research. Because of their undetermined risks stem cell-based therapies, as a rule, must aim at being clinically competitive or superior to existing therapies. Where there are already efficacious therapies for a disease condition, the stem cell-based intervention must be of low risk and offer a potential advantage (such as better functional outcome; single procedure vs. lifelong drug therapy). Greater risks are permissible where there is no efficacious therapy and where the disease condition is severely disabling and life threatening (Recommendation 25). Patients need to be informed that cell-derived products have never been tested before in humans, that researchers do not know whether they will work as hoped, and that, unlike many pharmacological products or even many implantable devices, stem cells may stay in the body and generate adverse effects for the lifetime of the patient. To respect their values, and because some subjects may have moral objections to the use of embryo-derived cells, subjects should be informed about the source of the cells (Recommendation 28). To determine the consequences of cellular implantation, and with consideration of cultural and familial sensitivities, subjects should be asked for autopsy in the event of death (Recommendation 31). Considerations of social justice apply to all research involving human subjects but receive special importance in view of public involvement in this emergent research area. Among the recommendations of the task force are public engagement in the policy making of governmental agencies (Recommendation 37), fair allocation of benefits and risks, and the need to develop alternative models of intellectual property, licensing, product development, and public funding to promote fair and broad access to the new diagnostics and therapies (Recommendation 38). One justice consideration especially pertinent to stem cell research is the establishment of stem cell collections with genetically diverse sources of cell lines (Recommendation 38). The task force concludes its recommendations with an aspirational goal that companies, subject to their financial capability, should offer affordable therapeutic interventions to persons living in resource-poor countries who would not otherwise have access to these therapies.

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Universities and other institutions are asked to incorporate this requirement in their intellectual property licenses (Recommendation 39).

Clinical Application Recommendations for the ethical introduction of stem cell therapies take place against a background marked by the presence of many clinical programs offering “stem cell therapies” outside the boundaries of proper clinical trials and ethical or legal oversight [95]. These unsupervised stem programs exploit a loophole in the FDA regulation of biologics that permits the use of cells that have been subject to “minimal manipulation” [96]. In some cases desperate patients have suffered serious illnesses, including growths caused by proliferating and ectopic transplanted stem cells [97]. In its 2016 report, the ISSCR issues a stern rebuke to such programs. The ISSCR condemns the administration of unproven stem cell-based interventions outside of the context of clinical research or medical innovation compliant with the guidelines in this document and relevant laws, particularly when it is performed as a business activity. Scientists and clinicians should not participate in such activities as a matter of professional ethics. For the vast majority of medical conditions for which putative “stem cell therapies” are currently being marketed, there is insufficient evidence of safety and efficacy to justify routine or commercial use [82].

The ISSCR guidelines do not altogether prohibit innovative clinical therapies. They affirm that “unproven stem cell-based interventions” can be provided to at most a very small number of patients outside the context of a formal clinical trial but in accordance with a set of “very restrictive conditions.” These include the existence of a written plan outlining the scientific rationale and justification of the procedure, why it has a reasonable chance of success, and why it should be attempted compared to existing treatments. This plan should offer a full characterization of the types of cells being transplanted, how they will be administered, and a commitment to clinical follow-up and data collection to assess the procedure’s effectiveness and adverse effects. The plan should be peer reviewed by independent experts, and have the full support and accountability of the health care institution where it is based. Patients must provide voluntary informed consent and understand that the intervention is unproven. There should be an action plan for adverse events, including insurance coverage for patients. Finally, the conduct of innovative treatments requires a commitment by clinician-scientists to use their experience with individual patients to contribute to generalizable knowledge. This includes communication of all outcomes to the scientific community in professional meetings and journals, and timely movement to formal clinical trials [82].

CONCLUSION Fully answering all of the questions this chapter has raised would require an ethical treatise. However, in my view, the moral claims of the very early embryo do not outweigh those of children and adults that can be helped by hESC and therapeutic cloning technologies. While iPSC research is promising, reliance on it should not replace the use of hESC lines produced through proven techniques, and that are currently available. As we move forward to clinical translational research, care should be taken that there is proper oversight for these novel and potentially harmful therapies. Some may disagree with these conclusions. Continuing dialog about these questions and clearer scientific research results will bring us closer to a consensus on these issues.

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[88] Green RM, Olsen DeVries K, Bernstein J, Goodman KW, Kaufmann RW, Kiessling AA, et al. Overseeing research on therapeutic cloning: a private ethics board responds to its critics. Hastings Cent Rep 2002;32:27e33. [89] National Institutes of Health. Guidelines for research using human pluripotent stem cells (effective August 25, 2000. 65 FR 5 1976) (corrected November 21, 2000. 65 FR 69951). Bethesda, MD: National Institutes of Health; 2000. [90] ISSCR (International Society for Stem Cell Research). Guidelines for the conduct of human embryonic stem cell research. 2006. Available at: http://www.isscr.org/home/publications/guide-clintrans. [91] ISSCR (International Society for Stem Cell Research). Guidelines for the clinical translation of stem cells. December 3, 2008. Available at: http://www.isscr.org/home/publications/ClinTransGuide. [92] Sugarman J, Siegel A. How to determine whether existing human embryonic stem cell lines can be used ethically. Cell Stem Cell 2008;3:238e9. [93] O’Rahilly R, Mu¨ller F. Human embryology and teratology. New York: Wiley-Liss; 1992. [94] Hyder N. Geron issues statement on halted stem cell trial. Bionews September 6, 2009;524. Available at: http://www.bionews.org.uk/page_ 47650.asp. [95] Turner L, Knoepfler P. Selling stem cells in the USA: assessing the direct-to-consumer industry. Cell Stem Cell 2016;19:1e4. [96] FDA (Food and Drug Administration). Minimal manipulation of human cells, tissues, and cellular and tissue-based products draft guidance for industry and Food and Drug Administration staff. 2014. Available at: http://www.fda.gov/BiologicsBloodVaccines/ GuidanceComplianceRegulatoryInformation/Guidances/CellularandGeneTherapy/ucm427692.htm. [97] Kolata G. A cautionary tale of ‘stem cell tourism’. NY Times June 22, 2016. Available at: http://www.nytimes.com/2016/06/23/health/acautionary-tale-of-stem-cell-tourism.html?rref¼collection%2Ftimestopic%2FNew%20England%20Journal%20of%20Medicine&action¼click &contentCollection¼timestopics®ion¼stream&module¼stream_unit&version¼search&contentPlacement¼3&pgtype¼collection.

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C H A P T E R

77 Overview of the US Food and Drug Administration Regulatory Process Carolyn Yong1, David S. Kaplan2, Andrea Gray1, Laura Ricles1, Anna Kwilas1, Scott Brubaker1, Judith Arcidiacono1, Lei Xu1, Cynthia Chang2, Rebecca Robinson3, Richard McFarland3 1

Center for Biologics Evaluation and Research, FDA, Silver Spring, MD, United States; 2Center for Devices and Radiological Health, FDA, Silver Spring, MD, United States; 3Advanced Regenerative Manufacturing Institute, Manchester, NH, United States

INTRODUCTION AND CHAPTER OVERVIEW The field of regenerative medicine encompasses a breathtaking array of interdisciplinary scientific approaches that address a broad spectrum of clinical needs. Ongoing advances in scientific knowledge related to cell biology, gene transfer therapy, biomaterials, immunology, engineering principles, and technology applicable to biological systems foster innovation of regenerative medicine products. This places the regenerative medicine community in a position to address a number of challenging and critical health needs including the treatment of disease conditions resulting from pancreas, liver, and kidney failure; structural cardiac valve repair; skin and wound repair; and orthopedic applications, among others. Scientific challenges confronting this field include expanding the knowledge base in each discipline as well as developing an interdisciplinary approach for identifying and resolving key questions. The US Food and Drug Administration’s (FDA’s) regulatory review process mirrors the scientific challenges with regard to the development of review paradigms that cross scientific disciplines. This chapter will provide a brief historical review of the FDA and its organizational structure and discuss topics pertaining to the regulation of regenerative medicine products including possible regulatory pathways for combination products and relevant jurisdictional issues. Sources of information concerning FDA regulatory policies important to regenerative medicine product developers will also be discussed. It is essential for individuals, institutions, and companies responsible for the clinical trials of regenerative medicine products to be aware of FDA regulatory policies and how to obtain the necessary information. These entities are collectively referred to in FDA regulations as Sponsors (the term “Sponsor” for drugs and biologics is defined in 21 Code of Federal Regulations [CFR] 312.3[b] whereas “Sponsor” is similarly defined in 812.3[n] for devices). Suggestions will also be provided regarding how to engage FDA effectively during the development of a novel regenerative medicine product.

BRIEF LEGISLATIVE HISTORY OF UNITED STATES FOOD AND DRUG ADMINISTRATION Medical products regulated by the FDA include human and animal drugs, medical devices, and biological products such as vaccines, cellular and gene therapies, and blood products. Among the therapeutic agents of biological origin regulated by the FDA are cellular therapies, including products derived in whole or part from human tissue, and xenotransplants. In addition to medical products for human use, the FDA regulates food other than most meat and poultry; radiation-emitting products for consumer, medical, and occupational use; cosmetics; medical products Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00077-1

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for veterinary use; and animal feed. The FDA’s role in medical product regulation extends throughout the entirety of the product life cycle. Depending on the product category, this may mean oversight, including review and inspection, of clinical trials, of the premarket product approval/clearance process, of manufacturing controls, controls over labeling, and registration and listing requirements. After a product is marketed, the FDA also continues its oversight in a variety of ways, including inspections, mandatory and voluntary postapproval (e.g., Phase IV) studies, and surveillance of adverse events reported to the FDA. FDA laws and regulations have developed over time, partly in response to serious medical adverse events or by other public health and safety concerns. Early regulation of biological products was prompted in part by the death of 13 children in 1901 after the administration of diphtheria antitoxin prepared from a source contaminated with tetanus. In response, Congress passed the Biologics Control Act in 1902. This act provided for regulation of viruses, serums, toxins, and analogous products; required licensing of manufacturing establishments and manufacturers; and provided the government with inspectional authority. The Act focused on requiring control of manufacturing processes for producing biological products, reflecting the extent to which the starting source material and the manufacturing process defined the final product. In 1906, Congress passed the Federal Food and Drugs Act, proposed in part in reaction to the meat-packing industry conditions described in Upton Sinclair’s book, The Jungle. Although the primary focus of the Act was on food safety, the law also required that drugs be provided in accordance with standards of strength, quality, and purity unless otherwise specified in the label. Premarket review of new drugs was not required until the passage of the 1938 Food, Drug, and Cosmetic Act (FD&C Act), which repealed the earlier 1906 Federal Food and Drugs Act. In 1937, the sulfa drug Elixir Sulfanilamide, previously available only in tablet or powder form to treat streptococcal infections, was marketed as a liquid using diethylene glycol, an analog of antifreeze, as a formulating solvent. This change in formulation, which was made without the requirement for premarket review, resulted in over 100 deaths, many of whom were children, and prompted the passage of the 1938 FD&C Act. The 1938 Act also put medical devices and cosmetics under FDA authority and authorized factory inspections. The Public Health Service Act (PHS Act), passed in 1944, incorporated the 1902 Biologics Control Act and is the current legal basis for the licensing of biological products. Because most biological products also meet the definition of “drugs” under the FD&C Act, they are also subject to regulation under that Act. The requirement for premarket demonstration of efficacy and the authority for FDA oversight of clinical trials were provided by the KefauvereHarris amendments to the FD&C Act in 1962. These amendments were prompted in part by the tragic adverse events resulting from use of thalidomide as a nonaddictive prescription sedative. This drug, which was not approved as a sedative in the United States, resulted in thousands of birth defects in children born outside this country when taken by pregnant women during the first trimester. The Medical Device Amendments to the FD&C Act were passed in 1976, after reports of safety issues with respect to the Dalkon Shield intrauterine device. The Medical Devices Amendments included risk-based requirements for premarket notification or approval of medical devices. Before 1976, FDA authority was limited to taking action against marketed devices found to be unsafe or ineffective.

LAWS, REGULATIONS, AND GUIDANCE The previous section summarized the history of laws that form the underpinning of FDA medical product regulation. This section provides a brief description of how laws are made and implemented, the procedures for promulgating regulations, and a description of how the FDA develops and uses guidance documents. Laws are created as an outcome of legislative activity conducted in the US Senate and House of Representatives resulting in passage of a bill. Once Congress passes a bill, it becomes law if it is signed by the president. If the president vetoes the bill, it becomes law if two-thirds of the Senate and House of Representatives vote in favor of the bill. A federal law also is denoted as a public law and may contain a name, such as the FD&C or PHS Acts. These laws are then incorporated into the US Code (USC), which is updated every 6 years, with supplements published regularly to incorporate changes to statutes between updates. Drug, biologic, and device laws can be found in the USC at: • Drugs and Devices: Title 21 Chapter 9, and • Biologics: Title 42 Chapter 6A. When laws are passed, government agencies such as the FDA often implement them by promulgating regulations. Sometimes an agency may elect to promulgate regulations on its own, whereas other laws may explicitly require an agency to issue regulation. The process for making regulations must be performed in accordance to the Administrative Procedures Act (Title 5, USC, Chapter 5). This Act generally requires agencies such as the FDA to provide public notice and opportunity for comment as part of the rule-making process.

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FDA regulations are contained in the CFR. Regulations for drugs, biologics, devices, tissues, and combination products, along with related regulations, may be found in various parts of Title 21 of the CFR. The following is a list of key regulatory provisions: • • • • • • • • • •

Drugs: 21 CFR Parts 200e299, 300e369; Biologics: 21 CFR Parts 600e680; Devices: 21 CFR Parts 800e898; Human Cells, Tissues, and Cellular and Tissue-based Products (HCT/P): 21 CFR Parts 1270/1271; Combination Products: 21 CFR Parts 3 and 4; Recalls: 21 CFR Part 7; Informed Consent/Institutional Review Boards (IRBs): 21 CFR Parts 50/56; Financial Disclosure by Clinical Investigators: 21 CFR Part 54; Good Laboratory Practice for Nonclinical Laboratory Studies: 21 CFR Part 58; and Good Guidance Practices: 21 CFR Part 10.115;

Guidance documents are nonbinding publications that describe the FDA’s interpretation of policy pertaining to a regulatory issue or set of issues related to: • the design, production, labeling, promotion, manufacturing, and testing of regulated products; • the processing, content, and evaluation or approval of submissions; and • inspection and enforcement policies. Guidance documents, which are developed in accordance with Good Guidance Practices found at 21 CFR 10.115, are intended to clarify the FDA’s current thinking related to regulatory issues and procedures. Unlike regulations and laws, guidance documents are not enforceable. Therefore, Sponsors may elect to choose alternate approaches that still comply with existing laws and regulatory requirements. In most cases, guidance documents are issued in draft form for public comment before implementation. In cases in which prior public participation is not feasible or appropriate, the FDA may issue a guidance document for immediate implementation without first seeking public comment. The publication of guidance documents reflects the FDA’s efforts to convey up-to-date information on general and cross-cutting topics to those involved in the developing field of regenerative medicine. When considering developing a guidance document, the FDA may freely discuss related issues with the public. In fact, the FDA may hold a public meeting, advisory committee meeting, or workshop to obtain input on scientific issues. Finally, after receiving public input, the FDA will evaluate submitted comments and finalize the document. Guidance documents are a useful way for the FDA to communicate current thinking to the public. Within the arena of regenerative medicine, it is valuable to be aware of both product-specific and cross-cutting guidance documents. Some of the more pertinent guidance documents to this field, such as those related to preclinical testing, manufacturing practices, and clinical trial design, are discussed in this chapter. In addition to guidance documents, the FDA may refer to guidelines published by the International Council for Harmonization of Technical Requirements for Pharmaceuticals for Human Use (ICH). ICH is an international effort to harmonize regulatory requirements. ICH guidelines, similar to FDA guidance documents, are nonbinding.

FOOD AND DRUG ADMINISTRATION ORGANIZATION AND JURISDICTIONAL ISSUES Scientific development of regenerative medicine products involves extensive testing and planning before marketing authorization. It can be helpful for individuals and organizations involved in product development to engage in early dialog with the appropriate FDA review unit to receive and consider the FDA’s comments on the design of the preclinical and clinical development plan. This section describes the FDA’s organizational structure and provides basic information regarding jurisdictional decisions made to determine the appropriate regulatory pathway for a broad range of products. The FDA is an agency within the Department of Health and Human Services (DHHS) consisting of seven centers and the Office of the Commissioner. Three of the centers are responsible for regulating medical products for humans. The Center for Biologics Evaluation and Research (CBER) has jurisdiction over a variety of biological products, including blood and blood products, vaccines and allergenic products, and cellular, tissue, and gene therapies, as well as some related devices. The Center for Devices and Radiological Health (CDRH) has jurisdiction over diagnostic and therapeutic medical devices, administration of the Mammography Quality Standards Act program, and ensuring the safety of radiation-emitting products. The Center for Drug Evaluation and Research (CDER) has jurisdiction over a variety of drug products, including small molecule drugs and well-characterized biotechnology-derived therapeutic biological products that include monoclonal antibodies and cytokines.

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For many medical-use products, it is clear which center within the FDA should have primary jurisdiction for the premarket review. For other products, including some technologically novel products under development, determining which center has jurisdiction for review may be unclear. Important starting points for determining product jurisdiction are the formal regulatory definitions of biological products, drugs, devices, and combination products. The formal definitions are as follows: • Biological Product (42 USC 262[i](1)): A virus, therapeutic serum, toxin, antitoxin, vaccine, blood, blood component or derivative, allergenic product, protein (except any chemically synthesized polypeptide) or analogous product, or arsphenamine or derivative of arsphenamine (or any other trivalent organic arsenic compound), applicable to the prevention, treatment, or cure of a disease or condition of human beings. • Drug (21 USC 321[g] [1]): (A) Articles recognized in the official US Pharmacopeia, official Homeopathic Pharmacopeia of the United States, or official National Formulary, or any supplement to any of them; and (B) articles intended for use in the diagnosis, cure, mitigation, treatment, or prevention of disease in man or other animals; and (C) articles (other than food) intended to affect the structure or any function of the body of man or other animals; and (D) articles intended for use as a component of any articles specified in clause (A), (B), or (C). • Device (21 USC 321[h]): An instrument, apparatus, implement, machine, contrivance, implant, in vitro reagent, or other similar or related article, including any component, part, or accessory, which is: (1) recognized in the official National Formulary, or the US Pharmacopeia, or any supplement to them, (2) intended for use in the diagnosis of disease or other conditions, or in the cure, mitigation, treatment, or prevention of disease, in man or other animals, or (3) intended to affect the structure or any function of the body of man or other animals, and which does not achieve its primary intended purposes through chemical action within or on the body of man or other animals and which is not depend on being metabolized for the achievement of its primary intended purposes. • Combination Product (21 CFR 3.2[e]): (1) A product composed of two or more regulated components, that is, drugedevice, biologicedevice, drugebiologic, or drugedeviceebiologic, that are physically, chemically, or otherwise combined or mixed and produced as a single entity; (2) two or more separate products packaged together in a single package or as a unit and composed of drug and device products, device and biological products, or biological and drug products; (3) a drug, device, or biological product packaged separately that according to its investigational plan or proposed labeling is intended for use only with an approved individually specified drug, device, or biological product where both are required to achieve the intended use, indication, or effect and where upon approval of the proposed product the labeling of the approved product would need to be changed, for example, to reflect a change in intended use, dosage form, strength, route of administration, or significant change in dose; or (4) any investigational drug, device, or biological product packaged separately that according to its proposed labeling is for use only with another individually specified investigational drug, device, or biological product where both are required to achieve the intended use, indication, or effect. The FDA’s Office of Combination Products (OCPs), located in the Office of the Commissioner, has broad administrative overview responsibilities covering the regulatory life cycle of drugedevice, drugebiologic, devicee biologic, and drugebiologicedevice combination products. When jurisdiction is uncertain, Sponsors may contact the OCP for assignment of primary regulatory review responsibility for combination and other medical products. Jurisdictional determinations are made after a formal submission process called a Request for Designation. The appropriate FDA center jurisdiction is determined by considering the primary mode of action of the product.

APPROVAL MECHANISMS AND CLINICAL STUDIES There are several premarket approval pathways for medical products, depending on whether the product is a drug, biological product, or device. Approval pathways, explained in more detail subsequently, include the New Drug Application (NDA) for drugs and the Biologics License Application (BLA) for biologics. The Premarket Approval Application (PMA), Humanitarian Device Exemption (HDE), 510(k) clearance mechanism, and De Novo classification process are various regulatory pathways used for medical devices. Clarification regarding the type of application needed for a particular regenerative medicine product may be helpful to the Sponsor early in development to enable the Sponsor to discuss the data needed for a marketing application during the planning stage. NDAs and BLAs are applications for licensure under the safety and effectiveness requirements of the FD&C Act. Biologics are further subject to the approval standards set forth in the PHS Act that requires demonstration that the product is safe, pure, and potent. Further information concerning the licensure of drugs and biological products is

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provided in “Guidance for Industry: Providing Clinical Evidence of Effectiveness for Human Drugs and Biological Products” [1]. For a medical device, the appropriate regulatory pathway and classification depend heavily on the associated risks; Class I includes devices with the lowest risk and Class III includes devices with the highest risk. A PMA is an application for approval for most Class III medical devices; the Sponsor must show reasonable assurance of safety and effectiveness [2]. The premarket notification, or 510(k), clearance process applies to products that are “substantially equivalent” to a Class I or II (or in a few cases, a Class III) medical device already on the market [3]. The De Novo process provides a pathway to Class I or Class II classification for medical devices for which there is no legally marketed device to enable a demonstration of “substantial equivalence” but for which regulatory controls provide a reasonable assurance of safety and effectiveness [4]. Under medical device regulation, a product can also gain approval through an HDE, which is a marketing approval for certain devices for rare conditions and is exempt from the effectiveness requirements of a PMA but requires demonstration of safety and probable benefit [5]. To qualify for this type of application, a Sponsor would first need to receive a designation from the FDA Office of Orphan Products Development that the medical device is a Humanitarian Use Device intended for treatment or diagnosis of a disease or condition that affects or is manifested in not more than 8000 individuals per year in the United States. Many but not all combination products are approved or cleared under a single marketing application. For example, depending on the specific facts, including the primary mode of action of the product, a combination biological device could be licensed under the biologics authorities or approved under the medical device authorities. After approval of a marketing application, there are also postmarketing requirements such as reporting [6].1,2 In addition, modifications to the product or labeling may require prior approval. The FDA has published regulations and guidance documents that address submission and approval processes for modifications to marketed products [3,6].3,4,5 Compliance with manufacturing requirements is also an ongoing Sponsor obligation. The FDA has issued a guidance document entitled “Guidance for Industry and FDA Staff: Current Good Manufacturing Practice Requirements for Combination Products,” which provides direction regarding applicable manufacturing requirements for combination products [7]. Owing to the relatively new nature of regenerative medicine and its developmental status, postapproval topics will not be further discussed in this chapter. When clinical investigation is needed to evaluate the safety and efficacy of an investigational product before marketing approval, an Investigational New Drug (IND) application is required for drugs and biologics and an Investigational Device Exemption (IDE) is generally required for devices [8,9]. For both types of applications, the Sponsor needs to submit a description of the product and manufacturing process, preclinical studies, a clinical protocol, information on any other prior investigations such as human clinical studies, and a rationale for the study design. An IRB review and informed consent are also required. The FDA has 30 days to review the application to determine whether the study may proceed. The contents are specifically laid out in FDA regulations for each type of application. Requirements for the content of an IND can be found at 21 CFR 312.23, and for an IDE at 21 CFR 812.20. For some products, there may be applicable guidance with respect to the manufacturing information and the preclinical data needed to support the study. For example, “Guidance for FDA Reviewers and Sponsors: Content and Review of Chemistry, Manufacturing, and Control (CMC) Information for Human Somatic Cell Therapy Investigational New Drug Applications (INDs)” provides information on characterization and manufacturing of a cellular product to be submitted in an IND [10]. Information on a general preclinical study design for regenerative medicine products using cellular or gene therapies can be found in the “Guidance for Industry: Preclinical Assessment of Investigational Cellular and Gene Therapy Products” [11]. For medical devices, “Guidance for Industry and FDA Staff, Investigational Device Exemptions (IDEs) for Early Feasibility Medical Device Clinical Studies, Including Certain First in Human (FIH) Studies” [12] provides information on early feasibility study IDEs that allow for early clinical evaluation of significant risk devices. Early feasibility study IDEs may be appropriate early in device development when clinical experience is necessary because nonclinical testing methods are not available or adequate to provide the information needed to advance the developmental process. Applicable regulations and guidance should

1

Reporting for Biological Products: 21 CFR 314.80 and 21 CFR 314.81, 21 CFR 600.14, 21 CFR 600.80, 21 CFR 600.81, 21 CFR 601.28, 21 CFR 601. 70, and 21 CFR 601.93. 2

Postmarketing Reports for Applications for FDA Approval to Market a New Drug (NDA): 21 CFR 314.80 and 314.81.

3

PMA supplements: 21 CFR 814.39. (2007).

4

Supplements and Changes to an Approved NDA: 21 CFR 314.70, 314.71, 314.72. (2007).

5

BLA 21 CFR 814.39, 21 CFR 314.70, 21 CFR 314.71, 21 CFR 314.72.

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be further consulted for information on adverse event reporting, labeling, study conduct and monitoring, and other topics related to requirements for conducting an IND [8] and IDE [13e15]. For information on general clinical study design and conduct issues, the FDA has many guidance documents that may be helpful [16,17]. For some indications there may be guidance documents that apply across technologies, such as the “Guidance for Industry: Chronic Cutaneous Ulcer and Burn WoundsdDeveloping Products for Treatment” [18] and “Guidance for Industry: Cellular Therapy for Cardiac Disease” [19]. In addition, guidance documents not directly applicable for a specific product, indication, or technology may be worth consulting, because the documents may provide some insights into general clinical issues, such as assessment parameters that may be valuable.

MEETINGS WITH INDUSTRY, PROFESSIONAL GROUPS, AND SPONSORS Although the terminology and procedures may vary, all three FDA centers performing medical product review encourage meetings with Sponsors to address questions before a regulatory submission and at specific developmental milestones. When requesting a formal or informal meeting with the FDA, it is helpful to provide background information as well as specific discussion questions. Further information about formal meetings, such as what to include in a meeting request and what type of information to include in the briefing package submitted before the meeting, is provided in “Guidance for Industry: Formal Meetings Between the FDA and Sponsors or Applicants” [20] and “Guidance for Industry and FDA Staff, Requests for Feedback on Medical Device Submissions: The PreSubmission Program and Meetings with Food and Drug Administration Staff” [21]. Early-stage device meetings are addressed in “Early Collaboration Meetings Under FDA Modernization Act (FDAMA): Final Guidance for Industry and CDRH Staff” [22]. The FDA also interacts with organizations representing a group of interested parties (e.g., the International Society for Cellular Therapy, the American Association for Blood Banks, the Biotechnology Industry Organization, and Pharmaceutical Research and Manufacturers of America), which provides an opportunity to discuss topics of interest to the FDA and the organization. These interactions can be valuable for the FDA and stakeholders because they are a way to understand general issues of concern better, as opposed to product-specific discussions with individual firms. In addition to such interactions and meetings with individual sponsors, the FDA has various advisory committees that review available data and information and make recommendations related to a variety of issues, many of which are pertinent to the field of regenerative medicine. Advisory committees will be discussed further in the Advisory Committee Meetings section.

REGULATIONS AND GUIDANCE OF SPECIAL INTEREST FOR REGENERATIVE MEDICINE The topics discussed thus far have been of general applicability to medical product regulation: marketing pathways, clinical trial regulation, meetings, guidance development, and related topics. This section will review a few topics of particular interest to the scientific community engaged in developing regenerative medicine products: FDA regulations of human tissue products, product characterization for cellular products, FDA policy and guidance on xenotransplantation, and gene therapy.

Regulation of Human Cells and Tissues Intended for Transplantation An understanding of the regulations applicable to cells and tissues is important for developers of regenerative medicine products because human cells or tissues comprise the whole of many products or are a key component of them. In 1997, noting the fragmented approach to the regulation of human cellular and tissue-based products, the FDA issued the “Proposed Approach to the Regulation of Cellular and Tissue-Based Products” [23]. This document proposed a tiered risk-based approach to regulating these products. According to the proposed approach, products posing a lesser degree of risk would be subject to the rules designed to minimize communicable disease risks and additional regulatory requirements would be imposed on products posing an additional risk. The proposed approach to regulating human tissues was implemented in three parts, collectively referred to as the “tissue rules”: Establishment Registration and Listing, Donor Eligibility, and Current Good Tissue Practice (CGTP). The complete

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set of rules went into effect on May 25, 2005 and is codified in 21 CFR Part 1271. The tissue rules derive from the statutory authority of section 361 of the PHS Act, which addresses the control of spread of communicable diseases. Because the tissue rules apply to all human cellular and tissue-based products, it is important for Sponsors of regenerative medicine products to be aware of these rules, as well as the specific additional requirements for biologics or devices that may apply, depending on the particular regulatory pathway for their products. With some exceptions that are noted in the tissue rules, human cells or tissue intended for implantation, transplantation, infusion, or transfer into a human recipient are regulated as an HCT/P. Examples of HCT/Ps are musculoskeletal tissue, skin, ocular tissue, heart valves, dura mater, reproductive tissue, gestational tissue such as umbilical cord tissue and amniotic membrane, and hematopoietic stem/progenitor cells. Specifically excluded are vascularized organs, minimally manipulated bone marrow, blood products, xenografts, secreted or extracted products such as human milk and collagen, ancillary products, and in vitro diagnostic products. Tissue rules require HCT/P establishments to do the following: • register and list their HCT/Ps with FDA (21 CFR 1271 Subparts A and B); • evaluate donors through screening and testing to reduce risk for transmission of infectious diseases through transplantation (21 CFR 1271 Subpart C); and • follow CGTP to prevent the spread of communicable disease (21 CFR 1271 Subpart D). Additional requirements for reporting, labeling, inspections, importation, and enforcement are described in 21 CFR 1271 Subparts E and F; these provisions apply only to HCT/Ps regulated solely under Section 361 of the PHS Act, and therefore would not apply to most regenerative medicine products. The Establishment Registration rule defines the circumstances under which a product would be subject to the HCT/P rules only (21 CFR 1271.10), and when there would be additional regulatory oversight such that a BLA, PMA, or other marketing application would be required (21 CFR 1271.20). Products that meet all of the following conditions are regulated by the FDA solely under the HCT/P rules: (1) the HCT/P is not more than minimally manipulated, (2) the HCT/P is intended for homologous use, (3) the HCT/P is not combined with a drug or device (with certain exceptions), and (4) the HCT/P does not have a systemic effect and is not dependent upon the metabolic activity of living cells for its primary function (except for autologous use or allogeneic use in a first or second degree blood relative, or reproductive use). If any of these four conditions are not met, a marketing application is required. The Tissue Reference Group handles various inquiries from stakeholders concerning application of the HCT/P rules including generating recommendations for consideration for CBER, CDRH, and OCP regarding whether specific HCT/Ps meet the criteria specified in 21 CFR 1271.10 for regulation solely under Section 361 of the PHS Act. Additional information and documents regarding these rules, as well as electronic forms for registration and listing, can be found on the FDA’s website for Tissue and Tissue Products [24]. A joint FDAeCenters for Disease Control and Prevention (CDC) workshop held in 2007 on the Processing of Orthopedic, Cardiovascular, and Skin Allografts is relevant to the regenerative medicine field. The workshop discussed pertinent information regarding current clinical practices, experiences, expectations, and assumptions of safety when using tissue allografts; the challenges of processing allografts with a focus on disinfection and sterilization methods; and the usefulness, reliability, and validation of tissue culturing methods. Some manufacturing topics discussed at the workshop are addressed in “Guidance for Industry: Current Good Tissue Practice (CGTP) and Additional Requirements for Manufacturers of HCT/Ps” [25]. The FDA issued “Guidance for Industry: Eligibility Determination for Donors of Human Cells, Tissues, and Cellular and Tissue-Based Products (HCT/Ps)” to assist establishments in making donor eligibility determinations with compliance to the Donor Eligibility rule (21 CFR 1271 Subpart C) [26]. Since 2007, multiple guidance documents have been published to provide updated or new recommendations for HCT/P donor screening and donor testing as well as guidance involving regulatory considerations for HCT/Ps from adipose tissue, to meet minimal manipulation criterion, interpret and apply homologous use, and interpret the scope of the same surgical procedure exception under 21 CFR 1271.15(b) [27].

Human Cellular Therapies Although therapeutic products composed of or containing cells vary greatly in the specific details of their application, cell or tissue source, manufacturing process, and characteristics, there are regulatory considerations that apply to all cellular preparations being developed as investigational regenerative medicine products intended for early-phase clinical studies. Control of the source material, demonstrated control of the manufacturing process,

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and characterizations of the cellular product that results from the manufacturing process are three such safety and quality considerations and will be discussed briefly next. The cell source will vary for different products and may be autologous or allogeneic, undifferentiated stem or progenitor cells, or terminally differentiated cells. Ensuring the safety of source cellular materials used during manufacture of an investigational regenerative medicine product begins by determining the eligibility of the donors selected to provide the source material through screening and testing. This screening and testing are part of the tissue rules described earlier in this chapter. Autologous products are not required to comply with the donor screening and testing requirements in the tissue rules but they carry labeling requirements to communicate risks for infectious substances. However, if autologous tissue either is positive for specific pathogens or has not been screened or tested, it is recommended that manufacturers document whether tissue culture methods could propagate or spread viruses or other adventitious agents to persons other than the recipient [10]. Donor eligibility determination is required for all allogeneic donors of cells and tissues. There are other aspects of the cell source may raise substantial concerns. In addition to screening and testing donors for communicable disease agents, according to the “ICH Guidance on Quality of Biotechnological/Biological Products: Derivation and Characterization of Cell Substrates Used for Production of Biotechnological/Biological Products,” the FDA has suggested that Sponsors consider the importance of evaluating donor medical history information and the relevance of conducting specified molecular genetic testing as part of an overall comprehensive assessment program to establish the suitability of a specific cellular preparation for use in manufacturing a regenerative medicine somatic cellular product [28]. The rationale for and feasibility of collecting additional information about molecular genetic testing were discussed in a public meeting of the FDA Biological Response Modifiers Advisory Committee (now known as the Cellular, Tissue, and Gene Therapies Advisory Committee [CTGTAC]) convened on Jul. 13e14, 2000, on the topic of “Human Stem Cells as Cellular Replacement Therapies for Neurological Disorders” [29]. A description of the physiological source of the cellular material, including the tissue of origin and phenotype such as hematopoietic, neuronal, fetal, or embryonic, conveys important information about the cells and their critical attributes. Control of the manufacturing process provides assurance about the consistent, reproducible production of the cellular component. Often, manufacturing will involve a multistep process that must be performed using aseptic techniques to prevent introduction of microbial contamination [30]. In addition to these precautionary techniques, the final product resulting from the manufacturing process should be demonstrated to be free of viable contaminating organisms. For manufacturing processes that involve in vitro cell culture, the cell product should also be tested for mycoplasma contamination, which may be introduced by manufacturing reagents or the culture facility environment [31]. A final rule became effective on Jun. 4, 2012, which amended the sterility regulations in part to provide manufacturers of biological products with greater flexibility, as appropriate, and encourage use of scientific and technological advances in sterility test methods as they become available.6 Many types of reagents may be used to manufacture the cellular component of a product, including those that promote cellular replication or induce differentiation, and those used to select targeted cell populations, specifically serum, culture medium, peptides, cytokines, and monoclonal antibodies. It is possible that the manufacturing process uses reagents that are FDA approved or cleared, clinical grade, or research grade. Depending on the grade of a particular reagent, additional documentation may be needed to verify the source, safety, and performance of the reagent. For example, some materials, such as serum, may be human- or animal-derived products, which have an increased risk for containing adventitious agents and therefore require further documentation of the safety testing performed on each lot of material. It is essential that reagents be properly qualified [10,31e33]. Demonstration of manufacturing control is evidenced by strict adherence to standard operating procedures and quality control assessment of manufacturing intermediates, as well as testing of the final cellular preparation. Because of inherent biological complexity, it is unlikely that a unique biomarker or other single analytical test will be sufficient to permit full characterization of a cellular product. Accordingly, as recommended in “Guidance for Industry: Guidance for Human Somatic Cell Therapy and Gene Therapy,” the FDA asks Sponsors to provide documentation that their testing paradigm developed for the final cell product encompasses a multiparametric approach that may involve biological, biochemical/biophysical, and/or functional characterization [10,33,34]. Therefore, in addition to microbiological testing, tests developed to demonstrate the following should be conceived to determine the degree to which the characteristics of the manufactured cell preparation conform to desired and specified criteria [10,33,34]: 6

Amendments to Sterility Test Requirements for Biological Products (77 FR 26162).

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• identity of the cell product (physical and chemical characteristics identifying the product as being what is designated on the label and distinguishing the product from other products manufactured in the same facility); • purity (freedom from contaminants including endotoxin, residual reagents, and unintended cell populations); and • potency/biological activity (the specific ability of the cells, as indicated by appropriate laboratory tests, to effect a given result relevant to the specific indication for use). This process can be challenging for a number of reasons. For example, the mechanism of action associated with a cell product may be incompletely understood, which may constrain the ability to develop a specific potency assay. Direct assessment of potency for a cellular preparation may not be possible owing to a lack of appropriate in vitro or in vivo assay systems. On Feb. 9e10, 2006, FDA CTGTAC discussed this challenging topic and obtained input on alternative approaches for performing potency assessments of cellular therapy products [29]. A guidance document was issued in Jan. 2011, to provide manufacturers of cellular and gene therapy products with recommendations for developing tests to measure potency [35]. Because process and product development are iterative and will continue throughout the life cycle of the cellular product, continued multiparametric characterization of the cellular product throughout the manufacturing process may aid in identifying critical process steps, establishing relevant specifications and acceptance criteria, and demonstrating comparability after manufacturing changes. Ensuring the safety of cell products that in and of themselves constitute a regenerative medicine product or that constitute a component of a product requires demonstrated control over each facet of the manufacturing process. This assurance begins with acquisition of the source material and is carried forward through manufacturing and characterization of the final cellular preparation using specified analytical tests based in large measure on the intrinsic biological properties of the cell product.

Xenotransplantation The success of allogeneic organ transplantation has increased the demand for human cells, tissues, and organs. Scientific advances in the areas of immunology and molecular biology, coupled with the growing worldwide shortage of transplantable organs, have led to increased interest in xenotransplantation. In addition to the potential use of xenotransplantation to address the shortage of human organs for transplantation, there are increasing efforts to use other xenotransplantation products in treating disease. An example of this is the investigational use of encapsulated porcine pancreatic islet cells to treat type 1 diabetes. Along with the promise of xenotransplantation are a number of challenges, including the potential risk for transmission of infectious agents from source animals to patients and the spread of any zoonotic disease to the general public [36]. In addition, the potential exists for recombination or reassortment of source animal infectious agents, such as viruses, with nonpathogenic or endogenous human infectious agents, to form new pathogenic entities. These considerations demonstrate the need to proceed with caution in this area. The US Public Health Service (PHS) Agencies, including the FDA, the National Institutes of Health (NIH), the CDC, and the Health Resources and Services Administration, have worked together to address the risk of infectious disease transmission, publishing the “PHS Guideline on Infectious Disease Issues in Xenotransplantation” [37]. This Guideline discusses xenotransplantation protocols, animal source, clinical issues, and public health issues. After publication of the PHS Guideline, the FDA published a guidance document entitled “Guidance for Industry: Source Animal, Product, Preclinical and Clinical Issues Concerning the Use of Xenotransplantation Products in Humans” to build on concepts in the PHS Guideline and provide more specific advice regarding xenotransplantation product development and production, and xenotransplantation clinical trials [38]. Xenotransplantation is defined in the PHS Guideline and FDA Guidance as any procedure that involves the transplantation, implantation, or infusion into a human recipient of live cells, tissues, or organs from a nonhuman animal source or human body fluids, cells, tissues, or organs that have had ex vivo contact with live nonhuman animal cells, tissues, or organs. Examples of investigational xenotransplantation products are: • transplantation of xenogeneic hearts, kidneys, or pancreatic tissue to treat organ failure, or implantation of neural cells to ameliorate neurological degenerative diseases; • administration of human cells previously cultured ex vivo with live nonhuman animal, antigen-presenting, or feeder cells; and • extracorporeal perfusion of a patient’s blood or blood component through an intact animal organ or isolated cells contained in a device to treat liver failure.

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The use of genetically modified source animals has the potential to overcome physiological and immunologic barriers to successful xenotransplantation as demonstrated in pig to nonhuman primate animal models [39]. When using genetically engineered (GE) animals for xenotransplantation, consultation with the Center for Veterinary Medicine (CVM) before initiating a clinical trial is recommended. GE animals are regulated by CVM under the New Animal Drug provisions of the FD&C Act and its enabling regulations (21 CFR 511 and 514). The requirements and recommendations for producers and developers of GE animals and their products are further clarified in “Guidance for Industry 187: Regulation of Genetically Engineered Animals Containing Heritable Recombinant DNA Constructs.” Review teams for xenotransplantation products consist of experts from CBER, CVM, and, when applicable, CDRH. Medical products that do not contain living cells are not considered to be xenotransplantation products by this definition. Therefore, products that include some common animal-derived components such as collagen, small intestinal submucosa, and heart valves do not automatically fall under this category. When a product does meet the definition, xenotransplantation guidelines are applied as appropriate to the specific product. The FDA encourages any potential sponsor of a xenotransplantation product to familiarize themselves with available documents that can be found on the FDA’s website [40].

Gene Therapy Human gene therapy seeks to modify or manipulate the expression of a gene or to alter the biological properties of living cells for therapeutic use. Products that mediate their effects by transcription or translation of transferred genetic material, or by specifically altering host (human) genetic sequences, are considered gene therapies. Gene therapy products are diverse and include genetically modified viruses (viral vectors), genetically modified microorganisms (e.g., bacteria, fungi), genome-edited/editing products, and ex vivo genetically modified human cells. The FDA regulates human gene therapy products as biological products. The field of gene therapy holds great promise for treating a wide array of illnesses, from genetically inherited diseases such as cystic fibrosis or hemophilia to heart disease, acquired immune deficiency syndrome, graft versus host disease, and cancer. The use of gene therapy in the areas of wound healing, tissue repair, and tissue engineering is also being investigated. Many of the same regulatory concerns already discussed for human cellular therapies apply to gene therapies, including appropriate qualification of source materials, proper control of the manufacturing process, and sufficient characterization of the final product. There are a number of safety issues associated with gene therapy, however, which are unique to this area. Safety issues specific to gene therapy trials may include generation of replication-competent virus (when using replication-incompetent vectors), vector and/or transgene-associated immune responses, toxicity associated with transgene overexpression, and inadvertent germline transmission of vector. These risks are exemplified by (1) the death of a study subject in 1999 owing to toxicity after administration of high-dose adenovirus vector [41], and (2) genomic integration of retroviral vectors, which has been shown to result in genotoxicity. In the latter case, five children developed leukemia and one died as the direct result of altered gene expression owing to retroviral vector integration [42,43] Detailed recommendations from the FDA regarding what type of CMC information to submit in an early-phase study of gene therapy products are available in the “Guidance for FDA Reviewers and Sponsors: Content and Review of Chemistry, Manufacturing, and Control (CMC) Information for Human Gene Therapy Investigational New Drug Applications (INDs)” [33]. This guidance covers product manufacturing and characterization information (from component qualifications to final formulation procedures), product characterization and release testing (including microbiological testing, identity, purity, potency, as well as other testing), and product stability, giving specifics in these areas that are pertinent to gene therapy. Special considerations should also be made when conducting preclinical assessment of gene therapy products. For instance, preclinical testing for gene therapy products includes tests designed to determine the biodistribution (i.e., localization and persistence) of the vector in various tissues, as well as assess transgene expression. Additional information on the evaluation of biodistribution in preclinical studies can be found in Section IV.B of the “Guidance for Industry: Preclinical Assessment of Investigational Cellular and Gene Therapy Products” [44]. Clinical trials evaluating gene therapy products may also differ in design from clinical trials for other types of pharmaceutical products because of the distinctive features of these products as well as the target patient populations, outcome measures, and risks involved. Therefore, the FDA has also published advice regarding clinical trial design in the “Guidance for Industry: Considerations for the Design of Early-Phase Clinical Trials of Cellular and Gene Therapy Products” [45].

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Gene therapies may differ from conventional drugs in that vector and transgene expression may persist for the lifetime of the subject. In these cases, there is a risk for delayed adverse effects. Indeed, the previously mentioned leukemias in a clinical study of gene therapy for the treatment of X-linked severe combined immunodeficiency did not occur until approximately 3e15 years after exposure to the retroviral vector. These events highlight the need for long-term assessment of study subjects receiving gene therapies. The FDA has discussed these issues, noting that the assessment of risk is based on the persistence of vector sequences, the potential for integration into the host genome, and transgene-specific effects. The FDA has published “Guidance for Industry: Gene Therapy Clinical TrialsdObserving Subjects for Delayed Adverse Events,” which addresses the duration and types of observations to be performed based on the patient population and the risks presented by the gene therapy product [44]. Although regulatory authority for gene therapy clinical investigations rests with FDA, the NIH serves an important complementary role. In addition to funding a number of gene therapy research studies, the NIH provides an important forum for open public deliberation on the scientific, ethical, and legal issues raised by novel FIH gene therapy clinical research applications. This is accomplished through the Recombinant DNA Advisory Committee (RAC), an expert advisory committee to the NIH director [46]. Clinical studies discussed in this forum include studies funded by the NIH, as well as industry-funded studies conducted at clinical sites receiving NIH funding for recombinant DNA research. Also, since the 2013 report by the National Academy of Science and Medicine (previously known as the US Institute of Medicine), NIH-supported clinical investigators have needed to register gene transfer protocols with the NIH Office of Science Policy [47]. In consultation with appropriate regulatory and/or oversight authorities including, IRBs and Institutional Biosafety Committees, the RAC may also identify protocols for public review based on criteria outlined in the 2013 report.

CelleScaffold Combination Products Cellescaffold combination products often face unique product development challenges because of their inherent complexity. These products often combine metabolically active cells and extracellular matrix or other scaffold components into complex three-dimensional structures, which makes the manufacture, characterization, and study of these products a challenge. Such a complex product that is derived from chemically or physically combining multiple entities cannot be defined solely by the characteristics of the individual components alone. Other factors such as product assembly and the resulting cellescaffold interactions have critical roles in determining final product characteristics. Furthermore, depending on the intended therapeutic function, these products may be designed to remodel or degrade in vitro during processing in bioreactors and/or in vivo after transplantation during clinical use, thereby precluding complete functionality testing at the time of product release. Packaging, shipping, and shelf life for these dynamic cellescaffold products are also nontrivial considerations. As with other products, product safety and efficacy of cellescaffold products need to be supported by a combination of appropriate in vitro and in vivo preclinical testing. The FDA draws on its extensive experience in regulating mammalian cell products and other tissue-derived products in evaluating product safety and efficacy. Many of the important tests needed for the individual components, such as sterility, mycoplasma, pyrogenicity/ endotoxin, scaffold characteristics, adventitious agent testing, cellular viability, identity, and purity, are applicable for the combined product as well. Demonstration of product potency and/or performance is also necessary. The complexity of these products, however, presents challenges for product characterization. Animal models are often used to characterize safety and performance of cellescaffold products. However, these models have limitations that are confounded by anatomical as well as physiological differences between the animals and humans. Therefore, characterization of these products may require the development of new scientific techniques and assays [48]. In vitro analyses of cell scaffold products may have a critical role in product characterization and evaluation for safety, efficacy, and potency. Nondestructive product testing and rapid characterization tests may also aid in product development. The goals, challenges, and methods of in vitro analyses of cellescaffold products were discussed at an FDAeNational Institute of Standards and Technology (NIST) cosponsored workshop [49]. Challenges and approaches of correlating product characteristics with clinical performance of absorbable medical devices, similar to materials that may be used as cell scaffolds, were discussed at an ASTM InternationaleFDA cosponsored workshop [50]. A critical consideration for developers of cellescaffold combination products is determining which tests need to be conducted on individual components before assembly and which are most relevant after product assembly. For

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example, if undifferentiated cells are directed to differentiate during culture after combination with the scaffold, the individual cellular component will have a phenotype different from that of the cellular component in the final product. In addition, biocompatibility testing to determine the potential for an unacceptable adverse biological response to the scaffold component is generally most appropriately conducted on the scaffold before combination with the cellular component to prevent convolution of the results. The biocompatibility tests that need to be performed will depend on the type and duration of contact the scaffold will have with the body. Regarding biocompatibility, the FDA issued a guidance document on Jun. 16, 2016 to provide further clarification and updated information regarding the use of International Standard Organization (ISO) 10993-1, “Biological evaluation of medical devicesd Part 1: Evaluation and testing within a risk management process” to support applications to the FDA [51]. For many innovative products, such as cellescaffold combinations used for regenerative medicine, the final product and instructions for use can be expected to undergo iterative modifications over time. Consequently, refinement of the product by the Sponsor and review of product modifications by the FDA will be an ongoing process. It is critical for the Sponsor to have a good understanding of the product and key scientific and/or clinical issues that can affect the safety and efficacy of the product, including the establishment of appropriate manufacturing controls to ensure product quality and consistency. When changes are made in the composition or manufacture of the cell and/or scaffold component of the combination product, it is essential that the Sponsor fully evaluate the impact of such changes to the final product’s quality and function.

PRECLINICAL DEVELOPMENT PLAN The primary goals of the preclinical development program for an investigational device or pharmaceutical product is to establish a scientific rationale for the use of the product in a specific clinical investigation and to demonstrate an acceptable product safety profile. Thus, the conduct of pharmacology and toxicology animal studies is important to identify and characterize possible local and systemic adverse effects. These studies also serve to determine potential safe dose levels in humans and to guide selection of a dose escalation scheme (if applicable) and a clinical monitoring paradigm. For cellular therapy, gene therapy, and cellescaffold combination products there are frequently additional product-specific safety questions that might need to be addressed before initiation of a clinical trial. For example, for cellescaffold constructs, there may be safety concerns related to the cells (e.g., inappropriate cell proliferation or differentiation, potential for cell migration from the scaffold) as well as the scaffold (e.g., its ability to perform a specific function at the anatomic site of implantation, degradation properties, structural integrity). These safety concerns may be independent of, or interdependent on, each component of the construct. As a result of their increasing complexity, many of these products do not have an established paradigm for preclinical evaluation. Sponsors are therefore encouraged to approach the FDA early in their preclinical development program to discuss the scope and design of the preclinical studies that will support the use of their investigational product in a clinical population. The FDA has held public discussions at Advisory Committee meetings and published Guidance documents that focus on appropriate preclinical testing of cellular therapy, gene therapy, and cellescaffold combination products. These discussions and publications provide valuable recommendations and considerations for Sponsors as they prepare their preclinical development plan. In addition to various FDA Guidance documents that are referenced throughout this chapter, a list of relevant Advisory Committee meetings can be found in the Advisory Committee Meetings section.

CLINICAL DEVELOPMENT PLAN The goal of the clinical development program is to establish a product’s relative safety and efficacy under specific directions for use in particular diseases or conditions. In the field of regenerative medicine, variability in the product poses unique challenges in clinical trial design and conduct. An additional challenge is the need, with many of these products, to observe their integration into the host over a prolonged period. To facilitate the development of cellular therapy and gene therapy products, the FDA published the “Guidance for Industry: Considerations for the Design of Early-Phase Clinical Trials of Cellular and Gene Therapy Products” [45]. This guidance provides recommendations regarding selected aspects of the design of early-phase clinical trials of cellular therapy and gene therapy products.

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However, because of the wide variety of regenerative medicine products and their potential applications, a case-bycase assessment is warranted for the design of each clinical trial. The FDA encourages sponsors to communicate with the FDA throughout product development. Such communications can include formal meetings (e.g., presubmission meeting; end of Phase 2 meeting) and requests (e.g., request for Special Protocol Assessment for products regulated as a biologic, or an Agreement Meeting for products regulated as a device) and informal interactions [20,22,52]. In addition, the FDA encourages IND sponsors to plan ahead by developing and submitting a Target Product Profile (TPP) early in product development. The TPP is an evolving draft of the product labeling and should be revised as clinical development proceeds. Submission of the TPP can facilitate communications between the FDA and the sponsor and enable the FDA to provide more comprehensive advice regarding the overall development program [53].

Food and Drug Administration’s Standards Development Program Since its inception, the development and use of standards have been critical to the mission of the FDA. The use of standards in FDA medical product regulation began with the 1906 Federal Food and Drugs Act. As defined in accordance with the standards of strength, quality, and purity in the US Pharmacopeia and the National Formulary, drugs could not be sold in any other condition unless the specific variations from the applicable standards were plainly stated on the label (Federal Food and Drugs Act, 1906). In current times, federal government agencies, including the FDA, are encouraged when practical to use voluntary consensus standards, whether domestic or international, when performing regulatory activities in lieu of government-unique standards that are developed by the government for its own uses. Standard-setting activities include the development of performance characteristics, testing methodology, manufacturing practices, product standards, scientific protocols, compliance criteria, ingredient specifications, labeling, or other technical or policy criteria. As with guidance document development, in which Good Guidance Practices describe the FDA’s procedures for developing and using guidance documents, specific regulations describe FDA participation in outside standardsetting activities. Regulations governing this participation can be found in 21 CFR 10.95. In addition, the FDA’s Staff Manual Guide 9100.1 establishes agency-wide policies and procedures related to standards management activities to ensure a unified approach to standards within the FDA. Constructive FDA participation in organizations responsible for developing standards applicable to the products regulated by the agency is considered essential. The FDAMA of 1997 provides for the recognition of national and international standards in medical device reviews for IDEs, HDEs, PMAs, and 510(k)s [54]. A “recognized consensus standard” is a consensus standard that the FDA has evaluated and recognized, in full or in part, for use in satisfying a regulatory requirement and that the FDA has published in a Federal Register notice. A “consensus standard” is a standard developed by a private sector standards body using an open and transparent consensus process. Conformance with recognized consensus standards is strictly voluntary for a medical device manufacturer, who may choose either to conform to applicable recognized standards or to address relevant issues in another manner. Standards may also be used in support of nondevice applications when appropriate and not in conflict with an FDA regulation or Guidance. Lists of recognized standards, Guidance Documents, and Standard Operating Policies and Procedures can be found on the FDA’s website. The scientific and manufacturing novelty of many regenerative medicine products presents unique challenges for meeting regulatory requirements with respect to product testing, developing performance characteristics, testing methodologies, scientific protocols, product standards, ingredient specifications, and compliance criteria. Increased development and use of standardized analytical methods and metrics have the potential to facilitate product characterization and reduce time to market by leveraging industry and government efforts. CDRH and CBER actively engage with Standards Development Organizations (SDOs) such as ASTM International and the ISO. CDRH and CBER staff attend standards meetings and workshops and participate in developing the standards and formally voting on the publication of standards. ASTM Committee F04 on Medical and Surgical Materials and Devices Division IV is actively engaged in developing standards for tissue engineered medical products (TEMPs). F04 Division IV consists of six subcommittees: (1) Classification and Terminology, (2) Biomaterials and Biomolecules, (3) Cells and Tissue Engineered Constructs, (4) Assessment, (5) Adventitious Agent Safety, and (6) Cell Signaling. The ASTM TEMPs group has developed more than 25 published standards, including standard guides, standard practices, and test methods, and has approximately 30 draft standards under preparation. The first of these standards were developed for substrates, biomaterials, natural materials such as collagen, alginate, and chitosan, terminology, cells and cell processing, bone morphogenetic protein, assessment of adventitious agents, and test methods for

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characterizing biomaterials. These standards are reviewed on a regular basis by the appropriate ASTM subcommittee (SC) to ensure that the standards reflect current scientific knowledge. The FDA liaisons elected to represent the Agency on each of the six subcommittees review the standards to ensure that standards developed do not conflict with current FDA regulatory practices. Some examples of approved standards with which the TEMPs group was involved include standards for characterizing and testing biomaterial scaffolds, quantifying cell viability within biomaterial scaffolds, and repairing articular cartilage in vivo. Another SDO with which FDA is involved is ISO, a nongovernmental international organization that develops consensus standards in collaboration with both the private and public sector. Standards for regenerative medicine/tissue engineering products are developed in Technical Committee (TC) 150 SC 7, Tissue Engineered Medical Devices, and in TC 194, Biological Evaluation of Medical Devices. Within TC 150, SC 7 are two active Working Groups (WGs): WG 1 Management of Risk, and WG 3 Tissue-engineered medical products for skeletal tissues. TC 194 SC 01 is responsible for Tissue Product Safety, and within SC 01 are four WGs: WG 01 Risk Assessment, Terminology, and Global Aspects, WG 02 Sourcing Controls, Collection, and Handling, WG 03 Elimination and/or Activation of Viruses and Transmissible Spongiform Encephalitis (TSE) Agents, and WG 04 TSE Elimination. ISO standards of particular interest to regenerative medicine products are those in the 10993 series (10993-1e20). Additional ISO standards-setting activities for regenerative medicine/tissue engineering products take place in TC 276, Biotechnology, WG 01 Terminology, WG 02 biobanks and bioresources, WG 03 analytical methods, WG 04 bioprocessing, and WG 05 data processing and integration. Standards are available through ISO’s website: http:// www.iso.org. There are many benefits to standards development, adoption, and use. Participation in standards development activities benefits the regenerative medicine community in that these activities can facilitate the development and maintenance of guidance for industry, address issues not addressed in FDA Guidance documents, facilitate product design, and lead to international harmonization. From a business perspective, participation in standards development activities can provide a competitive advantage by influencing technical contents, reduce the cost of production by improving factory flexibility and supply chain management, foster innovation and support research and development, and increase trade by shortening the time between concept and global availability. Standards can be used by developers of regenerative medicine products to meet regulatory expectations. In turn, the benefit to the regenerative medicine community helps the public by providing products that are more thoroughly characterized with a greater understanding of the conditions of use of the products. Thus, the FDA has a critical role in providing support for standards development activities.

ADVISORY COMMITTEE MEETINGS As mentioned in the Meetings With Industry, Professional Groups, and Sponsors section, because of the diversity of innovative technology evaluated by FDA review staff, the FDA makes use of scientific Advisory Committees or Panels (for medical devices) to complement its internal review process. These advisors provide independent, professional expertise and technical assistance to contribute to scientific regulatory decision-making processes. Outside experts can be asked to review data or make recommendations about study designs across a product or clinical area; outside advisors can also be helpful at earlier stages of product development. Expertise on the advisory committee often includes scientific, statistical, and clinical experts, as well as consumer representation, patient advocates, and industry participation. Most meetings are public, and there is an opportunity for public participation in the form of public comment. There are 34 Advisory Committees as well as a number of SCs and one DHHS Committee administered by CBER. The areas of responsibility for the Panels and committees are divided along medical specialty areas. The Advisory Committee for cellular, tissue, and gene therapy products, as mentioned earlier, is the CTGTAC. This committee has discussed a number of topics that are of potential interest to product developers in the regenerative medicine area, including: • • • • •

Hematopoietic stem cells for hematopoietic reconstitution (Feb. 2003), Allogeneic islet cell therapy for diabetes (Oct. 2003), Somatic cell cardiac therapies (Mar. 2004), Cellular products for joint surface repair (Mar. 2005), Potency measures for cell, tissue, and gene therapies (Feb. 2006),

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• Cellular therapies derived from human embryonic stem cellsdconsiderations for preclinical safety testing and patient monitoring (Apr. 2008), • Animal models for porcine xenotransplantation products intended to treat type 1 diabetes or acute liver failure (May 2009), • Clinical issues related to FDA draft guidance “Preparation of IDEs and INDs for Products Intended to Repair or Replace Knee Cartilage” (May, 2009), • Testing of replication competent retrovirus/lentivirus in retroviral and lentiviral vector-based gene therapy products (Nov. 2010), • Cellular and gene therapy products for the treatment of retinal disorders (Jun. 2011), and • Oocyte modification in assisted reproduction for the prevention of transmission of mitochondrial disease or treatment of infertility (Feb. 2014). The presentations for each topic, as well as transcripts of each discussion, are available on the FDA’s website referenced at the end of this chapter [29]. The Medical Devices Advisory Committee consists of 18 Panels that cover the medical specialty areas. Panel meetings are held regularly to discuss specific products. A few examples of past Panel meetings are: • • • • • • • •

General and Plastic Surgery Panel regarding dermal fillers (Nov. 2008), Orthopaedic and Rehabilitation Devices Panel regarding a bone putty (Mar. 2009), Ophthalmic Devices Panel regarding a visual prosthetic device (Mar. 2009), Neurological Devices Panel regarding a spinal surgery sealant (May 2009), Ear, Nose, and Throat Devices Panel regarding an implantable hearing system (Dec. 2009), Orthopedic and Rehabilitation Devices Panel regarding metal-on-metal hip arthroplasty systems (Jun. 2012), Circulatory Systems Devices Panel regarding an atrial appendage device (Oct. 2014), and General and Plastic Surgery Devices Panel regarding wound care products combined with antimicrobials or other drugs (Sep. 2016).

A complete list of upcoming Medical Devices Advisory Committee Panel meetings and searchable archive of past meetings with agendas and materials can be found on the FDA’s website [55].

FOOD AND DRUG ADMINISTRATION RESEARCH AND CRITICAL PATH SCIENCE The FDA recognizes the complexity of the scientific issues related to regenerative medicine products. FDA research laboratories have an important role in ensuring that the agency stays abreast of the rapid changes and developments affecting the entire field as well as addressing specific regulatory science questions. In 1992, CBER researchers began systematic efforts to uncover mechanisms that control the behavior of cells subjected to various environmental stimuli, particularly those encountered during normal wound healing, regeneration, and prenatal development. Those studies led to the discovery of several growth factors and feedback mechanisms that control these pathways [56e63]. Other research addressed interactions between hematopoietic and mesenchymal cell lineages both in vitro and in vivo [64,65]. The success of those efforts led to the recruitment of additional investigators in these areas. In 2004, the FDA introduced the Critical Path Initiative to identify and support research priorities that are expected to advance innovation in medical products. The Critical Path Opportunities List “presents specific opportunities that, if implemented, can help speed the development and approval of medical products” [66]; it is available on the FDA’s website. A number of the research topics on the Critical Path Opportunities List have applications to regenerative medicine, such as developing characterization tools for cell therapy and tissue engineering, biomarkers for cardiovascular disease, and advanced imaging technologies. FDA laboratories are actively engaged in these and other research questions that will facilitate the advancement of the field of regenerative medicine. A priority of the Critical Path Initiative is updating and modernizing techniques, to ensure that the agency and the research community have the tools necessary to bring safe and effective products to market. Efforts include promoting collaborations spanning multiple centers and regulatory jurisdictions across the FDA as well as among the FDA and other relevant organizations (e.g., other agencies, academic organizations, regulated industry). Developing research collaborations and the requisite infrastructure to support those and other efforts will facilitate review of combination products, which are often seen in regenerative medicine. An example of collaborative work across FDA laboratories is a multiinvestigator project at CBER using a battery of state-of-the-art analytic techniques chosen to provide complementary data on cell state and seeking to develop

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new biomarkers for cell therapies. The project aims to characterize mesenchymal stromal cells from a number of perspectives, including genetic stability, proteomic and phosphoproteomic analysis, microRNA analysis, messenger RNA profiling by microarray and quantitative polymerase chain reaction analysis, chromatin immunoprecipitation, and examination of the potential for cells to mature and contribute to the formation of organs and tissues. Furthermore, the study will look for molecular differences between cells from early versus late passage numbers. Importantly, the same cells that go through this panel of tests will be implanted in a mouse model of hind-limb ischemia, allowing for the correlation of product characterization data with the in vivo outcome with regard to localization, differentiation, and functionality. This FDA research project may yield information that will be useful for product characterization, in-process testing, lot release criteria, developing comparability and stability protocols, and predicting cell fate and function after receiving a cell therapy. A major issue associated with the clinical use of cellular therapies is predicting what happens to the cells after administration. Another FDA Critical Path research study at CBER will advance cell therapy by developing methods for in vivo tracking and imaging of neural stem cells (NSCs) after transplantation. NSCs from adult, fetal, and embryonic sources have been proposed as treatments for degenerative conditions such as Parkinson disease and for the repair of tissues damaged by stroke and spinal cord injury. Magnetic resonance imaging and single photon emission computed tomography are being used to determine cellular location and the persistence of engraftment qualitatively and quantitatively. The goal of this project is to develop methods for evaluating biomarkers that may be predictive of NSC function. The FDA’s research projects often involve collaborations with other federal and academic partners to employ new technologies to address regulatory science questions. For example, a collaboration between FDA and the NIST is using automated microscopy to characterize the differentiation of mesenchymal stem cells (MSCs). The goal is to improve the safety of MSC products by developing robust assays that can be used for in-process and lot release testing. FDA Critical Path research also addresses some of the challenges faced in both product development and product evaluation. For example, after the observation of unexpected toxicity of adenoviral vector gene therapy in a clinical trial, CBER research provided insight into how adenovirus vectors cause toxicity and developed an animal model for gene therapy in the context of preexisting liver disease [67]. CBER researchers and regulators also worked with a consortium from industry and academia to develop reference material for adenoviral vector particles [68]. FDA research is ongoing to understand the nature of toxicity of systemically delivered adenovirus and mechanisms for vector clearance to improve the safety of gene therapy trials [69]. Some FDA laboratories are engaged in research projects related to tissue. CDRH scientists are studying the effects of mechanical and electrical stimulation on cardiac cell cultures and how the parameters modulate cellular physiology. A CBEReCDRH collaboration is examining the relationship between encapsulation of chondrocytes in a scaffold material, with and without mechanical stress, on the status of several signaling pathways. Another intercenter collaborative research project between CDRH and CDER conducts a comprehensive translational assessment of human induced pluripotent stem cell-derived cardiomyocytes to evaluate drug-induced arrhythmias [70] and to predict individual patient response to drugs by using patient-specific stem cells collected in an FDA-sponsored clinical trial [71]. A Critical Path collaborative research project with partnerships outside the US government involves CDRH, CBER, the University of Maryland, the University of Akron, and Tufts University, and evaluates the accuracy and reproducibility of three-dimensional (3D) printed scaffolds containing cells. This project assesses the material and biological properties of the printed scaffolds in the presence or absence of cells on printing accuracy. The FDA’s interest in 3D printing, also known as additive manufacturing, is in part because it is driving the innovation of medical products, including in the area of biological and tissue engineered products for regenerative medicine purposes. 3D printing involves the layer-by-layer deposition of material to produce a 3D part from a digital design file. With the increase in use of 3D printing and the uncertainty of how the technology can affect the safety and effectiveness of the products, interest in 3D printing has increased at the FDA and led to the formation of the Additive Manufacturing Working Group. The WG held a public workshop on Oct. 8e9, 2014, to obtain input from stakeholders, entitled “Additive Manufacturing of Medical Devices: An Interactive Discussion on the Technical Considerations of 3D Printing” [72]. The workshop aimed to bring together experts in the field to discuss with the FDA the current state of the art of 3D printing, and included discussions on 3D printing of medical devices, biologics, and pharmaceuticals. Some of the considerations discussed during the public workshop are presented in FDA web content and documents on 3D printed devices [73] and a technical review [74] These documents primarily focus on medical devices, with important considerations for evaluating 3D printed devices including the following: the effect of build orientation and location on final device performance, including mechanical and physical

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properties; process validation of the 3D printing systems to ensure consistency between print jobs; sterilization and removal of residual materials; and characterization of material properties before and after printing. In the case of printing biological products, also known as bioprinting, applications may include printing tissue scaffolds, constructs containing living cells [75], and even small tissue [76] and organ structures [77]. 3D printing may offer an approach to make a medical product with a complex architecture, such as porous structures or internal lattice structures. In addition, cells and biomaterials could potentially be printed simultaneously with more precise spatial control to produce constructs with desired properties. Although the FDA has experience regulating 3D printed medical products [74] no FDA-approved or cleared biological products incorporate 3D printing. These types of products will typically be regulated along the same pathways as non-3D printed products. However, because of the complexity of 3D printed biologics, there may be some additional technical considerations for sponsors to take into account and evaluate compared with none3D printed products. Some of these considerations may include, but are not limited to, the printing parameters and consistency, material selection, finishing steps, biocompatibility, mechanical and physicochemical properties, and the biological function of the finished product. Known applications of bioprinting technology used for research purposes include skin [76], cartilage [78], bone [79], nerve [80], and blood vessels [81]. With accelerating advancement in this area, the FDA continues to communicate with the public regarding considerations for 3D printed medical products. Through ongoing research collaborations, and discussions with stakeholders, the FDA is committed to fostering safe and effective innovation in the area of 3D printed medical products. With support through the Critical Path Initiative, FDA research laboratories provide an important source of inhouse expertise in regenerative medicine and other cutting-edge technologies and research areas. Although by no means exhaustive, the examples provided here demonstrate the diverse range of topics under investigation at the FDA. Especially in consideration of the rapid change and development of the regenerative medicine field, Critical Path research efforts ensure that the FDA stays abreast of current innovations. The FDA’s research programs provide an important source of the latest science to inform the regulatory process and bring safe regenerative medicine products to market.

OTHER COORDINATION EFFORTS Because of the highly interdisciplinary nature of regenerative medicine, the FDA recognizes the need to build collaborative efforts to review products successfully in this area. The FDA is therefore a partner in the MultiAgency Tissue Engineering Science (MATES) Interagency Working Group. This partnership, which spans more than a dozen federal agencies, is designed to provide a forum to facilitate communication and coordination across the government regarding activities in tissue engineering and regenerative medicine. The full strategic plan can be found at the MATES website [82]. In another example of interagency interaction, collaboration, and coordination, the FDA has established Memoranda of Understanding (MOU) agreements with two different NIH institutes that involve FDA scientific review staff and NIH extramural research program officers. The MOUs cosigned by the National Institute of Neurological Disorders and Stroke and the National Heart, Lung and Blood Institute incorporate safeguards to protect from disclosure shared, nonpublic information such as trade secrets and confidential commercial information, identities of study participants and other personal information, privileged and/or predecisional agency information, research proposals, progress reports, and/or unpublished data or information protected for national security reasons. Under the MOU agreements, participants are able to hold unfettered discussions and exchange information that enables the respective agencies to maintain currency with respect to ongoing scientific activities that could affect regenerative medicine from both the laboratory and clinical research perspectives. Interagency MOU interactions identify gaps in knowledge related to the state of available scientific information and familiarity with FDA regulatory expectations. This, in turn, contributes to identification of promising basic research with the potential for clinical translation. There are also efforts to promote collaboration within the agency. For example, the FDA Commissioner’s Fellowship Program (CFP) is facilitating a collaboration related to the regulation of regenerative medicine across the FDA’s product jurisdictions. The CFP was established in 2008 to attract new talent to the agency while providing an opportunity for those to learn about FDA regulatory science for future careers outside the agency. Within its annual cohort of fellows, a regenerative medicine fellowship program has been established. Because many regenerative medicine products involve combining biologics and device technology, regenerative medicine fellows work across both CBER and CDRH to facilitate cross-agency collaboration and conduct research projects related to the regulation of regenerative medicine products.

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Interaction between the FDA and the scientific and regulated communities is an important area of collaboration. Workshops are one such example, including the previously mentioned FDA-NIST cosponsored session regarding cellescaffold products. Workshops provide valuable opportunities for the agency to receive input from outside scientific experts and other stakeholders. An additional common example of this type of activity is the liaison meeting, in which the FDA directly engages in dialog with professional societies (e.g., International Society for Stem Cell Research, American Association of Blood Banks, American Association of Tissue Banks, International Society for Cellular Therapy) regarding scientific or regulatory issues related to a certain research area. The FDA is a founding member of the National Academy of Medicine (NAM) Forum on Regenerative Medicine. This forum, which is similar in structure and function to NAM’s longstanding Drug Forum, provides a neutral convening mechanism for interested parties from academia, industry, government, patient/provider organizations, regulators, foundations, and others to discuss difficult issues facing the application of, and opportunities for, regenerative medicine. The Kidney Health Initiative, an existing publiceprivate partnership that includes the FDA, has begun a roadmapping project to develop renal replacement therapies. CBER and CDRH have a leading role in this effort. All of these activities ensure that the FDA receives continuous input on the latest scientific discoveries to inform the regulation of safe products.

CONCLUSION The field of regenerative medicine is exciting, with scientific advances leading to the promise of future therapies for current unmet medical needs for patients. The FDA regulatory approach to the evaluation of medical products includes an ongoing assessment of how the science of those products informs regulatory policy. To meet the needs of the challenging array of products that are on the horizon, the FDA intends to continue the current dialog with the scientific community and product Sponsors as part of its mission to develop science-based regulatory review policies that are robust and predictable.

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C H A P T E R

78 Regenerative Medicine ManufacturingdChallenges and Opportunities Paul Cohen1, Joshua G. Hunsberger2, Anthony Atala2 1

North Carolina State University, Raleigh, NC, United States; 2Wake Forest Institute for Regenerative Medicine, Wake Forest University, Winston-Salem, NC, United States

WHY REGENERATIVE MEDICINE MANUFACTURING? Why regenerative medicine manufacturing? The answer is simple. For regenerative medicine-based technologies and products to become the next standard of care, advances are needed in manufacturing processes. This chapter will highlight the current challenges facing regenerative medicine and list what the primary challenges are that are hindering widespread adoption. We will then consider some of the challenges in lack of standards and lack of definition of quality, as well as some of the technical challenges. After highlighting these challenges, we will delve into the opportunities that exist in regenerative medicine manufacturing. These opportunities exist in areas of regulatory and standard setting and we will cover some of work currently being done in these areas. There are also opportunities for new technologies and we will cover some of the advances being made for scale-up, reducing costs, quality control systems, automation, modular plug and play systems, and bioprinting. Next we will consider regenerative medicine manufacturing systems of the future. While these systems do not currently exist in full form today we believe these systems will be part of every major clinical regenerative medicine program in the next 20 years. We will conclude this chapter by reviewing the global landscape focusing on technical societies and domestic and international efforts.

CURRENT CHALLENGES AND OPPORTUNITIES IN REGENERATIVE MEDICINE MANUFACTURING Primary Challenges for Widespread Adoption The primary challenges for widespread adoption of regenerative medicine-based technologies have been reviewed extensively [1e7]. Table 78.1 highlights some of these challenges. While the science of regenerative medicine continues to progress, for widespread adoption we must solve several roadblocks related to the manufacturing aspect of these treatments. Currently, high cost, quality, and logistics issues threaten the promise of bringing regenerative medicine therapies to those in need. The lack of efficient, controllable manufacturing processes makes cost and quality difficult to achieve. Moreover, Food and Drug Administration (FDA) approval procedures make continuous improvement far more difficult and costly. More efficient, cost-effective manufacturing will require advances in automation and control to remove human work and reduce variability where possible, because differences in procedures induce increased variability. Improved biosensors will also be needed to control quality and consistency, which will also drive down cost. Cell expansion is an example of an area, common to all therapies, in need of cost-effective scale-up. Progress is being made in the development of scalable suspension cultures that can generate billions of human cells using

Principles of Regenerative Medicine, Third Edition https://doi.org/10.1016/B978-0-12-809880-6.00078-3

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TABLE 78.1

Current Challenges in Regenerative Medicine Manufacturing

Need for scale-up (e.g., expand to billions of cells) High costs of manufacturing regenerative medicine product Lack of sufficient quality control systems for in-line sensing Lack of automation Lack of closed and modular systems Lack of standards for regenerative medicine

single-use bioreactors [8,9]. This necessitates advances in bioreactor design, as well as sensors to meet the large quantities of cells for a given product. Additionally, it should reduce cost and make these therapies more widely available [10].

Lack of Standards Regenerative medicine covers a broad range of therapies as it manipulates genes, cells, and tissues to repair or replace diseased, damaged, or missing organs. This may include bone, skin, as well as other cells and tissues. The transition from laboratory to clinic has been slow due to the inherent complexity of the treatments and a lack of standards and standardization that complicates manufacturing thereby increasing cost. Many traditional measurements of efficacy, potency, purity, and quality that work with traditional pharmaceutical manufacturing may not be sufficient for regenerative medicine treatments. The National Institute of Standards and Technology (NIST) has developed a regenerative medicine and advanced therapies laboratory program that advances cell therapy, gene therapy, and tissue engineering by developing measurement infrastructure, including enabling tools, reference materials, methods and protocols, and bioinformatics and modeling tools [11]. NIST also has a regenerative medicine biomanufacturing program that is seeking to develop measurement solutions, serve as a neutral ground for the discussion of underpinning measurements and other manufacturing needs, and also lead and contribute to the development of standards [12]. NIST has also entered into a memorandum of understanding with the Standards Coordinating Body for Gene, Cell, and Regenerative Medicines and Cell-Based Drug Discovery to advance the field by developing consensus standards. ASTM International and the International Organization for Standardization (ISO) have a partnership in additive manufacturing where they have jointly released the Additive Manufacturing Standards Development Structure [13]. This provides a framework for meeting the needs for new standards in areas including regenerative medicine. ASTM also has many technical committees that serve as working groups to develop new standards in areas such as additive manufacturing (Committee F42) and medical and surgical materials and devices (Committee F04). ISO is another standard setting organization that brings together technical experts from around the world to develop international standards. For instance, there are ISO standards currently available for tissue-engineered medical products (ISO 19090:2018) and quality management systems (ISO 90001:2015 and ISO 13485:2016). For cell therapies, standards for imaging tools and protocols for cell characterization (identification, quantity, and function) are needed, as well as reference materials for fluorescence imaging and flow cytometry. For tissue engineering standards for scaffold design, structure and properties are needed as are standards for the computational models required. FDA has released guidance for additive manufactured devices that are implanted [14]. While these are not yet approved and some may not apply to scaffolds, they illustrate the complexity of standards development for future manufacturing systems and the need to incorporate them into process design. These are summarized in Table 78.2. In addition to standards, the manufacture of regenerative medicine products needs standardization of procedures and protocols to minimize variability, thereby increasing quality and decreasing costs.

Logistics Currently, cells are preserved and stored at cryogenic temperatures. This impacts handling procedures, cost, efficacy, safety, equipment requirements, retrieval, and effects on the regenerative medicine product. In addition, the ability to distribute product to point-of-use locations is constrained by the number of cryogenic containers.

CURRENT CHALLENGES AND OPPORTUNITIES IN REGENERATIVE MEDICINE MANUFACTURING

TABLE 78.2

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Summary of Recent Food and Drug Administration (FDA) Guidelines for Additive Manufactured Devices

Area

Components

Overall device design

Comparison of desired feature sizes and tolerances to process

Patient-matched device design

Effects of imaging, software interacting with design models, complex design file conversion, cybersecurity and personally identifiable information

Software workflow

File format conversions, digital device design to physical device, build volume placement, addition of support material, slicing, build paths, machine parameters and environmental conditions, automated software validation

Material controls

Starting material, material reuse

Postprocessing

Full documentation

Process validation and acceptance activities

Process validation, revalidation, acceptance procedures, test coupons

Quality data

Analysis of sources of quality data to identify existing and potential causes of nonconforming product

Device testing

Description, mechanical testing, material characterization, residue removal and sterilization

Labeling

Recommended for Additive Manufactured (AM) devices that are patient matched, may include patient identifier, use, final design iteration or version used to produce the device

Sensors and Quality Control Systems There are significant challenges with the development of sensors and quality control systems for regenerative medicine manufacturing, although they differ between cell expansion, tissue, and organs. Product loss or slow growth for cell-based therapy batches may slow testing or attainment of full expansion. For cells, tissue and organ sensors or sensor suites are needed to properly control processes and assure quality.

Bioprinting While there a many bioprinters on the market, none meet the proposed FDA guidelines and can be used for human patients. As summarized in Table 78.2, there are many issues covered in the FDA proposed guidelines and no currently marketed bioprinter meets these. It is important to note that issues involving hardware design, software, material control, and validation will make this challenging.

Scale-Up and Automation Benchtop research does not easily scale up for production for a number of reasons. Since there is little standardization of processes causing poor reproducibility the translation to production is difficult. Also, it is time consuming and expensive to change protocols once approved by the FDA. Standardized protocols, whether for manual or automated manufacturing, will help reduce variability and enable scale-up. However, the human element adds uncertainty and cost. Therefore automation for autologous products is crucial for widespread adoption. Currently, there is no marketed, fully integrated, modular manufacturing system. Collectively, we have highlighted some of the opportunities in regenerative medicine manufacturing and listed them in Table 78.3. We mention that there can be technologies developed that can assist with scaling-up processes such as using single-use technologies [15]. There is also an opportunity for developing solutions for reducing costs, which could include multiple solutions discussed further in this perspective article [10]. Other opportunities exist with developing in-line sensing systems, automating processes through integration of robotics and artificial intelligence, and developing plug and play infrastructures to accommodate different manufacturing processes. A final set of opportunities exists in the development of standards and ensuring seamless integration of these new regenerative medicine-based products with regulatory pathways. The FDA has published two new draft guidance documents that are intended to assist with bringing innovative, safe, and effective products to patients as efficiently as possible. The first guidance document is called “Expedited Programs for Regenerative Medicine Therapies for Serious Conditions.” The second document is called “Evaluation of Devices Used with Regenerative Medicine Advanced

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TABLE 78.3

Current Opportunities in Regenerative Medicine Manufacturing

Develop technologies for scale-up (e.g., single-use technologies) Develop solutions for reducing cost (e.g., defined media) Develop solutions for quality control systems (e.g., microfabricated biosensor technologies for in-line sensing) Develop technologies that can automate processes (e.g., integration of robotics and artificial intelligence into manufacturing processes) Develop solutions for modular systems (e.g., plug and play infrastructure to accommodate different modules) Current ASTM and ISO efforts for developing standards (e.g., ISO 19090:2018 for tissue-engineered medical products) FDA guidance documents for advancing regenerative medicine-based products (e.g., Technical Considerations for Additive Manufactured Medical Devices)

Therapies.” The first guidance document describes the Fast Track designation and breakthrough Therapy designation, which are available for regenerative medicine therapies. It also describes requirements for the new Regenerative Medicine Advanced Therapy (RMAT) designation program, which was created by the 21st Century Cures Act. The second draft guidance document on devices provides the FDA’s current thoughts on evaluating devices used in the recovery, isolation, and delivery of RMATs.

ENVISIONED REGENERATIVE MEDICINE MANUFACTURING SYSTEMS OF THE FUTURE Envisioned regenerative medicine manufacturing systems of the future will enable these technologies, therapies, and products to be commercialized and made widely available. Table 78.4 illustrates some of these envisioned system areas, their system attributes, and companies of interest that have some of the capabilities needed to build these systems. Fig. 78.1 captures some of the envisioned manufacturing systems of the future that could bioengineer new tissues and organs by integrating a clinical-grade 3D bioprinting approach into the manufacturing process. Next we will cover some of the system attributes we believe will be critical for these advanced regenerative medicine manufacturing systems. Standardization will be critical for future regenerative manufacturing systems due to the inherent variability and need to make procedures more consistent within and between manufacturing sites. Both manual and automated manufacture will rely on adherence to standards and standardization. NIST is working to establish standards for cell measurement, live cell imaging, quantitative flow cytometry, and bioprinting. One concern is that measurements of efficacy, potency, purity, and quality that are used for pharmaceuticals may be insufficient for regenerative medicine. Moreover, standardized procedures for cell expansion, size of biopsies, and other critical operations are needed for manual and automated manufacture. Fully integrated, modular, and automated manufacturing systems are critical and this will help to drive standardization since automation will depend, to a certain extent, on this. Such a system would utilize closed pods or cassettes for all processing to remove the possibility of contamination. This “plug and play” design philosophy would also lend flexibility to these manufacturing systems allowing them to produce a wide array of regenerative medicine products for individuals. An integrated modular robotic system with closed pods could have the ability to exploit standard processes, including cell harvesting, cell processing, nutrient addition, tissue digestion, incubation, imaging and characterization, tissue banking, quality control, and preservation prior to use or shipping. The collection of data, scheduling, and control of such a system will be complex and require new sensors and smart controls. The use of machine learning to sense quality characteristics and remove human visual perceptions (and biases) will be needed. Control for processes that can be highly variable may take weeks and must be done in a way that gets needed cells, tissue, and organs to critically ill patients reliably. Smart manufacturing refers to

ENVISIONED REGENERATIVE MEDICINE MANUFACTURING SYSTEMS OF THE FUTURE

TABLE 78.4

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Envisioned Manufacturing Systems of the Future

Future System Areas

System Attributes

Fully integrated/modular/ closed/sterile/automated manufacturing systems

• • • • •

Xeno-free defined media systems

• Synthetic serum for human immune cells and mesenchymal stem cells first. Following its success it can be extended to other tissue sources in the human body • Universal “basal” media • Synthetic, defined serum substitute “panel”

Supply chain and logistic systems

• Platform technologies for shipping human stem cells and mesenchymal stem cells should be initially tested • Further development of formulations and methods for extending liquid storage stability • Further development and optimization of formulations and methods for freeze drying

Biosensing systems

• • • •

Nondestructive quality control systems

• • • • • •

Automated and closed patientspecific systems

• Integrated, standardized disposables for product types • Processes that are patient specific and would involve modeling the patient’s anatomy • Easily adaptable semiuniversal automated system • Semiuniversal disposables • Automated injector with detailed process control • Novel single-use sensors/sensing approaches for noninvasive monitor/control of process parameters and detection of microbial contamination • Disposable single-use bioreactors that will support all steps from seeding through harvest • Instrumentation that can provide CO2 and temperature control without traditional incubators

Cell and tissue expansion systems

• • • •

Closed, integrated purification, formulation and vial fill Seamless media to bioreactor/cell culture vessel transition Cell concentration standard method to maintain viability/potency Customizable modules for expansion and cell retrieval Automated cell handling in a fully controlled aseptic environment. Scale-up for mass production • Passaging and layering multiple types of cells in a closed system • Automatic monitoring of glucose utilization/lactose production to adjust nutrient supply for continuous feed • Off-the-shelf closed systems that can be easily upscaled

3D printer for engineered tissues Microfabrication of biosensors Better oxygen and glucose monitoring sensors Improved material for adherent cells to maximize cell seeding and harvesting

One area is lactate and ammonia management In-line measurement of cell density Real-time cell “state”/phenotype monitoring Data capture and mining/correlation Microscope with software for recognizing and quantifying cellular structures Disposable sensors or built-in sensors for growth factor levels in addition to pH, osmolarity, O2, and CO2. Glucose, ammonia, and Kþ would be great as well • There will not be a path to a universal solutiondtoo divergent. Pick either most universal quality control specification/parameter (most therapy agnostic) or highest impact upcoming therapydpropose and prototype a solution

Determine if same system could be scaled to fit both small- and large-scale needs In-line cell “state”/phenotype monitoring Scalable bioreactor technology Novel single-use format for high-density cell culture (adaptable adherent or suspension, or tissue constructs) • Design for scale-out strategy for autologous therapies

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FIGURE 78.1 Envisioned future manufacturing systems: illustrated here is an envisioned future manufacturing system where 3D bioprinting will be incorporated into an end-to-end modular manufacturing system that can produce at scale bioengineered tissues and organs.

“fully-integrated, collaborative manufacturing systems that respond in real time to meet changing demands and conditions” [16]. It has the potential to use data and models to control such complex systems. Another aspect of an integrated automated manufacturing system is the ability to share information. Blockchains, a computer science construct that uses a distributed recording and storage of transaction records in a manner that safeguards patient confidential information, can be used to link information through the process to understand cause and effect and better control production. Xeno-free defined media systems will be able to support a large panel of clinically relevant cell types that will be needed for cell and gene therapies, as well as tissue engineering applications. These systems will greatly accelerate product development times, as well as reduce costs and variability by removing serums and recombinant proteins and replacing them with defined biochemical substitutes. The RegenMed Development Organization is currently using a consortium-based model to develop a platform technology that seeks to build a universal media system similar to the one we envision here to support clinical cell manufacturing. Efforts like these will greatly accelerate the field. Biosensing systems will be integrated into regenerative medicine manufacturing processes to monitor the viability, phenotypic characteristics, biomechanical characteristics, and physiologic responses of cells, scaffolds, organoids, organoid systems (e.g., body-on-a-chip personalized medicine systems), tissues, and bioengineered organs. These biosensing systems will be microfabricated and allow improved monitoring of oxygen, glucose, metabolites, pH, temperature, and many additional attributes that can be customized based on the cellular and tissue-specific systems being engineered and manufactured. These biosensing systems can also be integrated into 3D bioprinting systems to ensure the integrity of starting materials (e.g., cells, biomaterials, bioinks, etc.) and confirm the viability, phenotypic characteristics, biomechanical characteristics, and physiological responses from 3D bioprinted organoids, tissues, and organs. Nondestructive quality control systems will provide in-line measurement of specific quality control attributes needed in regenerative medicine manufacturing processes of the future. One area of need will be improved monitoring of lactate and ammonia management, as well as other biochemical markers that provide important information to maintain the viability of cellular and tissue systems. These in-line measurements will capture data in real time and have artificial intelligence systems to make correlations and predictions on safety and quality attributes of the current and final clinical product that will also save time and money by only advancing products with a high probability of possessing final safety and quality attributes needed for the manufactured clinical product. Cell and tissue expansion systems will be adaptable to both small- and large-scale needs. The miniaturized systems will enable autologous patient-specific therapies at the bedside, while the large-scale systems will provide a pathway for allogeneic therapies to be manufactured at industrialized scale for widespread use. Within these two systems, in-line cell “state” or phenotype monitoring will be critical to ensure safety and quality attributes are maintained throughout the manufacturing process. There will also be opportunities for novel single-use formats for highdensity cell cultures.

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Supply chain and logistic systems must be used carefully when considering the expansion of regenerative medicine therapies. Current technologies are expensive as companies must buy and maintain an appropriate number of cryogenic storage containers for shipment. Otherwise, timely shipment and product quality cannot be guaranteed. Current thinking and research seeks novel cell preservation technologies for regenerative medicine. This may include the development of advanced tissue and organ preservation technologies and processes for cell banking, such as specimen harvest, cell retrieval, selection and expansion, tissue typing, storage, and distribution. Promising paths include the development of new and more stable dimethyl sulfoxide-free stabilizer formulation compositions, lyophilization formulations, and characterization of the effects on the product (cellular function) both pre- and postthaw identification of reagents that serve to preserve cellular function upon freeze drying. Additionally, novel drying technologies that optimize formulations for reconstitution are needed. Specialized packaging for cells, tissues, and organs is needed, as well as training to properly employ their technologies, as well as appropriate infrastructure at point-of-care facilities. Solutions may utilize short-term cryostorage coupled with new techniques for distribution. Such hybrid storage would enable initial cryogenic storage followed by noncryogenic distribution for high-quality products. Customized designs for closed preservation systems coupled with sensors may also be employed to monitor quality and enable longer preservation time.

GLOBAL LANDSCAPE FOR REGENERATIVE MEDICINE MANUFACTURING The global landscape for regenerative medicine manufacturing is exciting and filled with tremendous potential. Next we will highlight some of the technical societies that are forming in this area, as well as the numerous manufacturing-focused initiatives that are being launched around the world. Table 78.5 lists the initiatives, focus areas, and websites where you can obtain additional information.

Technical Societies There is a new technical society recently formed called the Regenerative Medicine Manufacturing Society (RMMS). RMMS is a new organization that has a vision of enabling the adoption of manufacturing platform technologies into standards, regulatory pathways, and commercial products by assembling a diverse network of stakeholders. These stakeholders fall into these major categories: (1) industry, (2) academia (basic and clinical research), (3) regulatory agencies, (4) nonprofit organizations, and (5) investment and funding agencies. RMMS has a strategic goal to bring these stakeholders together to achieve a number of regenerative medicine manufacturing aims, including the following: (1) enabling the development of scale-up tools and reagents; (2) enabling the development of standards; and (3) seamless integration of platform technologies with regulatory pathways. Additional information on this new technical society can be found at http://regenmedmanufacturing.org/. The first RMMS conference TABLE 78.5

Global Landscape in Regenerative Medicine

Initiative

Focus Areas

Website

Medical Technology Enterprise Consortium Regenerative medicine technologies and manufacturing

https://mtec-sc.org/

National Institute for Innovation in Manufacturing Biopharmaceuticals

Biopharmaceutical manufacturing

http://www.niimbl.us/

Biofab USA

Biofabrication

https://www.armiusa.org/

Centre for Commercialization of Regenerative Medicine

Cell and gene therapies Regenerative medicine technologies

https://ccrm.ca/

Catapult

Cell and gene therapy

https://ct.catapult.org.uk/

Japandregulatory

New regulations to accelerate regenerative medicine therapeutics

Regenerative Medicine Manufacturing Society

Regenerative medicine manufacturing

http://rmmanufacturingsociety.org/about/

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was held in Miami, Florida, in January 2018 in concert with the World Stem Cell Summit and brought together leaders from industry, academia, and government.

Domestic Efforts Efforts currently under way in the United States to advance regenerative medicine manufacturing are growing every day. 21st Century Cure’s Act, also known as the Cures Act, was signed into law on December 13, 2016, with the intention of assisting the acceleration of medical product development and bringing new advances to patients in a faster, more streamlined manner. The Cures Act seeks to incorporate the perspectives of patients into the development of drugs, biological products, and devices in the US FDA’s regulatory process. The Cures Act has the intention of speeding up the development and review of novel medical products (including medical countermeasures) and gives new authority to assist the FDA in building and retaining scientific, technical, and professional experts by establishing new product development programs such as (1) RMAT, a new expedited option for eligible biologics, and (2) the Breakthrough Devices Program to accelerate the review of innovative medical devices. The Cures Act also charges the FDA with improving the regulation of combination products by creating one or more intercenter institutes to coordinate activities in major disease areas between the drug, biologics, and device centers. In total, the Cures Act provides $500 million over 9 years to assist the FDA in implementing this law to accelerate the discovery, development, and delivery of 21st century cures. Currently, the FDA has submitted a work plan to Congress that lays out a budget for advancing seven strategic areas of activity: (1) patient-focused drug development, (2) advancing new drug therapies, (3) modern trial design and evidence development, (4) patient access to therapies and information, (5) antimicrobial innovation and stewardship, (6) medical device innovations, and (7) improving scientific expertise and outreach at the FDA. The National Network for Manufacturing Innovation (e.g., Manufacturing USA) has funded two institutes dedicated to advance regenerative medicine manufacturing through public/private partnerships. These institutes join 12 others, each focusing on a technology or industry group by bringing together industry, academia, and government to increase competitiveness through applied research, outreach, and workforce development. The National Institute for Innovation in Manufacturing Biopharmaceuticals (NIIMBL) has a mission “to accelerate biopharmaceutical manufacturing innovation, support the development of standards that enable more efficient and rapid manufacturing capabilities, and educate and train a world-leading biopharmaceutical manufacturing workforce, fundamentally advancing U.S. competitiveness in this industry” [17]. NIIMBL will initially seek to understand biomarkers for potency, develop sensors for these biomarkers, and use this knowledge and developments for cost-effective scale-up for mesenchymal stem cells, T-cells, and induced pluripotent stem cells. The Advanced Regenerative Medicine Institute “will make practical the large-scale manufacturing of engineered tissues and tissue-related technologies, to benefit existing industries and grow new ones” [18]. The institute seeks to develop innovations across five thrust areas: (1) cell selection, culture, and scale-up; (2) biomaterial selection and scale-up; (3) tissue process automation and monitoring; (4) tissue maturing technologies; and (5) tissue preservation and transport. The Medical Technology Enterprise Consortium (MTEC) is a nonprofit biomedical technology consortium that collaborates with multiple government agencies under a 10-year renewable Other Transaction Agreement with the US Army Medical Research and Material Command. MTEC fosters integrated research partnerships to provide solutions to military, veterans, and the civilian population. Consortium thrust areas include: (1) prevention, diagnosis, and treatment of infectious diseases, (2) care of combat casualties, (3) clinical and rehabilitative medicine, (4) military operational medicine, (5) medical simulation and information sciences, and (6) advanced medical technologies. Significant emphasis is placed on aspects of regenerative medicine that will facilitate manufacturing scale-up.

International Efforts There are also many regenerative medicine international efforts and three specific efforts that we will focus on here include: (1) the Centre for Commercialization of Regenerative Medicine (CCRM), (2) Catapult, and (3) Japan. CCRM is a nonprofit, public/private consortium whose mission is to generate sustainable health and economic benefits through global collaboration in cell and gene therapy, as well as regenerative medicine. CCRM has recognized there is an unmet need in advancing cell, gene, and regenerative medicine-based treatments to the clinic. This

REFERENCES

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Canadian-based effort uses stem cells, bioengineering, and biomaterials combined with their consortium and infrastructure to coordinate product development and commercialization. The Cell and Gene Therapy Catapult is a network of centers in the United Kingdom designed for innovation in specific strategic areas to promote future economic growth. Each Catapult center specializes in a different area of technology, including areas such as cell and gene therapy, high-value manufacturing, transport systems, medicines discovery, and others. This Catapult network provides access to technical capabilities, equipment, and additional resources to develop new products and services on a commercial scale. Japan is advancing regenerative medicine by enabling early commercialization and early reimbursement. This effort is in response to half the population of Japan, which is over age 50, and regenerative medicine-based therapies have been limited in the past because of the difficulty in progressing through Japan’s Pharmaceuticals and Medical Devices Agency. Prime Minister Shinz o Abe has launched an economic revitalization plan that includes U110 billion ($1 billion) in funding for stem cell research. In addition, two new regenerative medicine laws were enacted. Law No. 84/2013 provides an amendment for the Pharmaceutical Affairs Act, renamed the Pharmaceutical and Medical Device Act, and provides conditional marketing authorization. This new regulation provides revenue for regenerative medicine products as soon as this conditional approval is granted, which could be achieved after a small patient number phase 1 safety trial. Law 85/2013 is part of the Safety of Regenerative Medicine Act, and is for clinical and physician-led research. This Act focuses on the cells and allows them to be processed outside hospitals at accredited cell-processing centers to enable manufacturing processes that are more robust with the intention of promoting a safer product through a tier-based risk-dependent analysis.

SUMMARY AND CONCLUSIONS This chapter on regenerative medicine manufacturing presented a comprehensive overview of current challenges and opportunities that currently exist in this space. We have also provided a perspective on envisioned regenerative medicine manufacturing systems of the future with focus on specific attributes that we believe will be critical for widespread adoption of regenerative medicine-based technologies and products. Moreover, we briefly described the global landscape for regenerative medicine manufacturing and highlighted some of the technical societies and efforts currently under way.

References [1] Bayon Y, Vertes AA, Ronfard V, Egloff M, Snykers S, Salinas GF, et al. Translating cell-based regenerative medicines from research to successful products: challenges and solutions. Tissue Eng 2014;20(4):246e56. [2] Hourd P, Ginty P, Chandra A, Williams DJ. Manufacturing models permitting roll out/scale out of clinically led autologous cell therapies: regulatory and scientific challenges for comparability. Cytotherapy 2014;16(8):1033e47. [3] Hunsberger J, Harrysson O, Shirwaiker R, Starly B, Wysk R, Cohen P, et al. Manufacturing road map for tissue engineering and regenerative medicine technologies. Stem Cells Transl Med 2015;4(2):130e5. [4] Martin I, Simmons PJ, Williams DF. Manufacturing challenges in regenerative medicine. Sci Transl Med 2014;6(232):232fs16. [5] Ratcliffe E, Thomas RJ, Williams DJ. Current understanding and challenges in bioprocessing of stem cell-based therapies for regenerative medicine. Br Med Bull 2011;100:137e55. [6] Williams DJ. Overcoming manufacturing and scale-up challenges. Regen Med 2011;6(6 Suppl.):67e9. [7] Williams DJ, Sebastine IM. Tissue engineering and regenerative medicine: manufacturing challenges. IEE Proc e Nanobiotechnol 2005;152(6): 207e10. [8] Kehoe DE, Jing D, Lock LT, Tzanakakis ES. Scalable stirred-suspension bioreactor culture of human pluripotent stem cells. Tissue Eng 2010; 16(2):405e21. [9] Kwok CK, Ueda Y, Kadari A, Gunther K, Ergun S, Heron A, et al. Scalable stirred suspension culture for the generation of billions of human induced pluripotent stem cells using single-use bioreactors. J Tissue Eng Regen Med 2017;12. [10] Hunsberger J, Goel S, Allickson J, Atala A. Five critical areas that combat high costs and prolonged development times for regenerative medicine manufacturing. Curr Stem Cell Rep 2017;3. [11] NIST. Regenerative medicine and advanced therapies laboratory programs. 2017. [12] NIST. Regenerative medicine biomanufacturing. 2017. Available from: https://www.nist.gov/programs-projects/regenerative-medicinebiomanufacturing. [13] ISO AIa. Additive manufacturing standards development structure. [14] (FDA) UFaDA. Technical considerations for additive manufactured medical devices. 2017. Available from: https://www.fda.gov/ downloads/MedicalDevices/DeviceRegulationandGuidance/GuidanceDocuments/UCM499809.pdf. [15] Boedeker B, Goldstein A, Mahajan E. Fully disposable manufacturing concepts for clinical and commercial manufacturing and ballroom concepts. Adv Biochem Eng Biotechnol 2017. PMID: 29101419.

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[16] i-scoop. Smart industry and smart manufacturing- industrial transformation. 2018. Available from: https://www.i-scoop.eu/manufacturingindustry/. [17] NIIMBL. The National Institute for Innovation in Manufacturing Biopharmaceuticals. Available from: http://www.niimbl.us/. [18] ARMI. Advanced Regenerative Medicine Institute; 2018. Available from: https://www.armiusa.org/.

Index

‘Note: Page numbers followed by “f” indicate figures, “t” indicate tables and “b” indicate boxes.’

A

A1AT deficiency. See a1-Antitrypsin deficiency (A1AT deficiency) AA. See Acetic acid (AA); Activin A (AA) AAO. See Anodic aluminium oxide (AAO) AAV. See Adeno-associated virus (AAV) AbATE trial. See Autoimmunity-Blocking Antibody for Tolerance trial (AbATE trial) ABCA4 gene, 354 Abcg2+ SP cells, 250e251, 263e264 AC. See Articular cartilage (AC) Ac-NRARADADARARADADA-CNH self-assembling peptide hydrogel, 1090 Accessory Limb Model, 44 Accuracy, 95 Acellular grafts, 1232 Acellular renal scaffold, 1171e1172 Acellular scaffolds, 505, 810, 1064, 1132 Acellular tissue matrices, 1269e1270 b-(1e4)-2-Acetamido-2-deoxy-Dglucopyranose, 641 Acetic acid (AA), 545, 546f, 600 Acetobacter xylinum (A. xylinum), 641e642 N-Acetyl-m-aminophenol, 775e776 N-Acetyl aspartyl-glutamate synthetase (NAAGS), 378e380 N-Acetyl-glucosamine, 641 Acetyl-para-aminophenol (APAP), 773, 775e776 Acetylated low-density lipoprotein (AcLDL), 311e312 Acetylated tubulin-1 (AT-1), 1157 ACI. See Autologous chondrocyte implantation (ACI) Acid hydrolases, 39 Acid solubilization, 643 Acid-catalyzed hydrolysis, 669 Acidebase interaction, 596 ACL. See Anterior cruciate ligament (ACL) AcLDL. See Acetylated low-density lipoprotein (AcLDL) ACOG. See American Congress of Obstetricians and Gynecologists (ACOG) ACP. See Amorphous calcium phosphate (ACP) Acquired immunity, 684e691 Acrylonitrile butadiene styrene, 823 ACT1 peptide, 81 Actin b, 209

Actinemyosin complexes, 395e396 Active targeting, 721 Activin A (AA), 339e340 Actomyosin contractility, 449 Acute inflammation, 679 Acute kidney injury (AKI), 210, 1152, 1165 BRECS treating, 1157e1158 clinical experience with renal assist device to treating, 1154 renal assist device therapy causing by sepsis, 1153e1154 Acute liver failure, hepatocyte transplantation in, 233e234 Acute lymphoblastic leukemia (ALL), 1253 Acute myocardial infarction (AMI), 209e210, 223e224, 253, 258, 261, 318 Acute tubular necrosis (ATN), 143, 1150 AD. See Alzheimer disease (AD) Ad-BMP-7. See Adenoviruses encoding green fluorescent protein, BMP-7 (Ad-BMP-7) AD-cSVF. See Adipose-derived cellular stromal vascular fraction (AD-cSVF) ADA-SCID. See Adenosine deaminaseSCID (ADA-SCID) Adaptive immune system, 715, 938, 1088 cell types and function, 685t AdBMP-2. See Adenoviral bone morphogenetic protein-2 (AdBMP-2) Additional signaling pathways, 7e8 Additive manufacturing techniques (AM techniques), 887, 1360e1361 Adeno-associated virus (AAV), 285, 749e750, 1010e1011 Adenosine deaminase-SCID (ADA-SCID), 198 Adenosine triphosphatase (ATPase), 38, 232 Adenosine triphosphate (ATP), 276, 394e395 Adenoviral bone morphogenetic protein-2 (AdBMP-2), 1186 Adenovirus, 749, 1010e1011 Adenovirus-p21, 81 Adenoviruses encoding green fluorescent protein, BMP-7 (Ad-BMP-7), 915e917 Adherent junctions (AJs), 398

1377

Adhesion adhesion peptide-modified scaffolds, 731 celleECM interactions during healing of cutaneous wounds, 25e26 during regenerative fetal wound healing, 29 signal transduction events during, 19e22 ligands, 442 molecules, 619, 679 peptides, 567 Adipocytes, 286 Adipogenesis, 459 Adipose tissue, 207, 847, 1258 Adipose-derived cells clinical delivery of, 297e300 therapeutic safety of, 301e302 carcinogenesis and tumorigenesis, 301e302 Adipose-derived cellular stromal vascular fraction (AD-cSVF), 384 Adipose-derived MSCs (AMSCs), 707, 795, 955 Adipose-derived stem/stromal cells (ASCs), 83, 104, 207, 295e297, 300e301, 384, 543, 837, 854e855, 1168e1169, 1202, 1264, 1299 immunophenotype, 296 uses and conditioned medium for hair growth, 1299 adipose-derived stem/stromal cellsconditioned media (ADSC-CM), 1299 Adipose-specific MSC, 295 Adjuvancy, 688 Adjuvant therapies, 898e899 antibiotics, 898e899 patient-specific technology, 899 b-Adrenergic agonist, 780 Adriamycin-induced nephropathy, 1167 ADSC-CM. See adipose-derived stem/ stromal cells-conditioned media (ADSC-CM) ADSCs. See Adipose-derived stem/ stromal cells (ASCs) Adult c-Kit+ cardiac cells, 248e250 Adult cardiac progenitor cell types, 250e251 Adult cells, 1309 Adult epicardial progenitor cells, 251e252 Adult epicardium, 251e252

1378 Adult healing process, 31 Adult heart, cardiac stem/progenitor cells in, 248e252 Adult human ventricular cells, 1076 Adult pancreatic islets to stem cells, 335e336 b cells from adult stem/progenitor cells, 341e343 from pluripotent stem cells, 336e341 for replacement therapy, 336 MSCs to modulate immunity, 344e345 Adult populations, 1041 Adult skin adult wound healing and scar formation, 66e68 anatomy of, 66 fibroproliferative scarring, 68e72 underhealing, 72 Adult stem cells (ASCs), 181e182, 241, 254e257, 750, 929, 1168e1169, 1264, 1266e1268. See also Cardiac stem cells (CSC) b cells from, 341e343 populations, 1264 in vitro, 924e925 Adult wound healing, 66e68 classic stages of wound repair, 69f inflammatory cell recruitment to site of tissue damage, 68f Advanced Regenerative Medicine Institute, 1374 Advanced therapy medicinal product (ATMP), 1116, 1125e1126 Advanced tissue engineering-based approaches, 535 Advanced Tissue-Engineered Human Ectypal Network Analyzer (ATHENA), 780 Adventitial cells, 295 Advisory committee meetings in FDA, 1358e1359 AEC. See Amniotic epithelial cells (AEC); Apical epidermal cap (AEC) AEMA. See 2-Aminoethyl methacrylate (AEMA) AF. See Amniotic fluid (AF); Annulus fibrosis (AF) Affinity-based release, 510 AFM. See Atomic force microscopy (AFM) AFMSCs. See Amniotic fluid mesenchymal stem cells (AFMSCs) AFS cells. See Amniotic fluid stem cells (AFS cells) AG. See Anterior gradient (AG) AG gene, 44 AGA. See Androgenetic alopecia (AGA) Agarose, 641, 821e822, 1226 hydrogel, 838 Age-related macular degeneration (AMD), 351, 354, 1208e1210 Aged MuSCs, 281, 283 Agglomerative nesting, 101e102 Aggrecan, 618

INDEX

Aggregates, 1106e1110 Aging, 219 aging effects on MSCs, 1266 cellular heterogeneity in, 107e108 muscle stem celleintrinsic defects in, 283e284 musculoskeletal disorders, 212 AGM region. See Aorta-gonadmesonephros region (AGM region) AgNPs. See Silver nanoparticles (AgNPs) Agrin, 618 AHSCT. See Autologous hematopoietic stem cell transplantation (AHSCT) Airefluid interface models, 774 Aireliquid interface (ALI), 1063 AJs. See Adherent junctions (AJs) Ajuba LIM proteins, 5 AKI. See Acute kidney injury (AKI) ALD. See Alendronate (ALD) Aldehyde dehydrogenase (ALDH), 154e155, 1167e1168 ALDHbr, 154e155 potency assay, 155f Alendronate (ALD), 494 ALD-conjugated AuNPs, 494 Alginate, 454e455, 640e641, 702, 815e816, 821e822, 837, 940, 1048, 1090, 1104, 1270 Alginate-based Cryogel system, 729 Alginic acid, 640 ALI. See Aireliquid interface (ALI) Aligned anisotropic scaffolds, 1230 nanofilaments for contact guidancemediated growth, 1230f Alimentary tract, 1131 anal canal, 1141e1142 colon, 1141 esophagus, 1131e1134 small intestine, 1135e1141 stomach, 1134e1135 in vitro models, 1142e1143 Aliphatic polyesters, 705 ALK. See Anaplastic lymphoma kinase (ALK) Alkaline phosphatase (ALP), 542, 699e700, 854 Alkanethiolates, 443 ALL. See Acute lymphoblastic leukemia (ALL) AlloDerm, 614te615t, 1136 Allogeneic aortic graft segments, 1136 cells, 664, 1264 clinical-grade hMAPCs, 186 collagen, 621 extracellular matrices, 1241 hepatocytes, 239 matrices, 1032 osteodifferentiated ASCs, 707 peripheral blood stem cell transplantation, 197 solid organ transplants, 790e791 stem cell transplantation, 753 transplantation, 198

Allogeneic hematopoietic stem cell transplantation, 157e158, 198 Allografts, 591, 955, 1184 hepatocyte, 234 liver, 239 stem cells, 1290e1291 Alloimmunization, 157 Allosensitization, 688 ALP. See Alkaline phosphatase (ALP) ALS. See Amyotrophic lateral sclerosis (ALS) Alveolar bone, 888 regeneration, 916e917 Alveolar pneumocytes, 1059e1060 Alveolar type 1 pneumocytes (AT1), 1061 Alveolar type 2 pneumocytes (AT2), 1061 Alveolarization process, 1059e1060 Alveolospheres, 1065e1066 Alzheimer disease (AD), 496, 754, 910 Am. See Antheraea mylitta (Am) AM techniques. See Additive manufacturing techniques (AM techniques) Ambulatory renal replacement therapies, future advancements for, 1159e1160 Ambystoma larvae, 40e41 AMD. See Age-related macular degeneration (AMD) AMD3100, 310 American Congress of Obstetricians and Gynecologists (ACOG), 150e151 American Society of Blood and Marrow Transplantation (ASBMT), 149e150 AMI. See Acute myocardial infarction (AMI) Amino acid, 573 amino acidederived polymers, 573e574 sequences, 667 b-(1e4)-2-Amino-2-deoxy-Dglucopyranose, 641 g-Aminobutyric acid (GABA), 1203 N-Aminoethyl aminocaproyl dihydrocinnamoyl (KADD), 339e340 2-Aminoethyl methacrylate (AEMA), 702 Amnion, stem cells from, 133 AEC, 134e135 AF, 134 amniotic fluid stem cells, 138e144 amniotic MSCs, 135e138 placenta, 133e134 Amniotic epithelial cells (AEC), 134e135 Amniotic fluid (AF), 133 function, origin, and composition, 134 Amniotic fluid mesenchymal stem cells (AFMSCs), 133 Amniotic fluid stem cells (AFS cells), 133, 138e144, 140f, 1168. See also Cardiac stem cells (CSC); Embryonic stem cells (ESC); Human embryonic stem cells (hESC); Induced pluripotent stem cells (iPSCs); Mesenchymal stem cells (MSCs) characterization, 138e140 isolation and culture, 138 preclinical studies, 140e144

INDEX

heart, 141e142 hematopoietic system, 142 intestine, 143e144 kidney, 142e143 lung, 143 musculoskeletal system, 141 nervous system, 141 Amniotic mesenchymal stem cells, 135e138 characterization, 136 immunophenotype, 137t isolation and culture, 136 preclinical studies, 136e138 Amorphous calcium phosphate (ACP), 594 Amphibian developmental model, 192 Amphiphilic block copolymers, 572 AMSCs. See Adipose-derived MSCs (AMSCs) Amyotrophic lateral sclerosis (ALS), 172e173 Anagen, 1297 Anal canal, 1141e1142 Analgesics, 575 Anaphylactic reaction, 686 Anaplastic lymphoma kinase (ALK), 52 Anastomosis, 1136 Androgen replacement therapy, 1254, 1254f Androgenetic alopecia (AGA), 1298, 1303, 1305 Androgens induce TGF-B1, 1303 Angioblasts, 317 AngioChip, 1088 Angiogenesis, 27, 298, 313, 317, 680, 701, 751, 1138e1139 Angiogenic growth factors, 312e313 Angiogenin, 1302 Angiopoietin-1, 1088 Animal model, 172e173, 765, 770, 1232 Animalehuman chimeras, 1325e1326 Anisotropic scaffolds, 511e512 for nerve regeneration, 1229e1231 aligned anisotropic scaffolds, 1230 cell-seeded, longitudinally aligned nerve guidance conduits and channels, 1231 ECM molecules, 1231 neurotrophic factors, 1230 Annulus fibrosis (AF), 962 Anodic aluminium oxide (AAO), 1288 Anterior cruciate ligament (ACL), 1179 healing, 1189e1192 biological augmentation, 1190 combined biological and mechanical augmentation, 1191e1192 mechanical augmentation, 1190e1191 Anterior cruciate ligament of knee, 1184e1185 Anterior gradient (AG), 43e44 Anteroposterior (AP), 44 Antheraea mylitta (Am), 1125 Anti-NogoA, 1207 anti-VEGF. See Antivascular endothelial growth factor (anti-VEGF)

Antibiotics, 898e899 Antibody antibodyeantigen, 657 binding proteins, 531 orientation, 529f Anticoagulants, 1151e1152 Antiestrogen resistance 1, 398 Antigen presenting cells (APCs), 684, 715 APC-activating agents, 722 nanoparticle targeting of, 722e727 Antigen(s), 528e529, 684e685, 715 antigen-4, 308 antigen-loaded DCs, 727 antigen-specific cytotoxic Tcells, 715 Antiglycophorin A (GlyA), 315e316 Antiinflammatory cytokines, 41 Antiinflammatory drugs, 575, 1210e1211 Antimicrobial peptides, 898e899 Antisense gene therapy, 1187 Antithymocytic globulin, 999e1000 a1-Antitrypsin deficiency (A1AT deficiency), 235, 750 Antitumor effects of MAPCs, 186 Antivascular endothelial growth factor (anti-VEGF), 354, 1210e1211 Aorta-gonad-mesonephros region (AGM region), 192 AP. See Anteroposterior (AP) APAP. See Acetaminophen. See also Acetylpara-aminophenol (APAP) Apatite cements, 594e595 APCs. See Antigen presenting cells (APCs) Apical epidermal cap (AEC), 37, 41e42 AECenerve interaction, 42e44, 43f Aplastic canine serum, 929 Apligraft, 1287e1288 Apoptosis, 322, 980 celleECM interactions during healing of cutaneous wounds, 28 during regenerative fetal wound healing, 30 signal transduction events during, 24 neuronal, 371 Apoptotic pathways, 68 Aprotinin, 1047e1048 Aquatic vertebrates, 876e877 Araneus diadematus (A. diadematus), 645 Arg-Gly-Asp (RGD). See Arginine-glycineaspartate (RGD) ArgeGlyeAspeSer (RGDS). See Arginineglycine-aspartic acid-serine (RGDS) Arginase, 750 Arginine-glycine-asparate (RGD), 16, 442, 489, 616e617, 632, 656, 729, 811, 941, 1226e1227 peptide, 1226e1227 RGD sequence. See Arginine-glycineaspartate sequence Arginine-glycine-aspartic acid-serine (RGDS), 667, 703e704, 1116 peptides, 670 Argininosuccinic lyase (ASL), 237 Aromatic diisocyanates, 567 Arrhythmogenesis, 255

1379 Arteriovenous (AV), 1029 Arthritic hips and knees, 953 Arthritis, 937 osteoarthritis, 391 rheumatoid, 317 Articular cartilage (AC), 956e958 regeneration, 412 surface, 413 surgery of, 412 Artificial esophageal construct, 1132 Artificial scaffolds, 1063e1064 Artificial urethra, 565 ASBMT. See American Society of Blood and Marrow Transplantation (ASBMT) Ascorbic acid, 762 ASCs. See Adipose-derived stem/stromal cells (ASCs); Adult stem cells (ASCs) ASD. See Autism spectrum disorder (ASD) Asherman syndrome, 1242 Asia’s Stem Cell Center, 1321 ASL. See Argininosuccinic lyase (ASL) Assisted Human Reproduction Act, 1321 Astemizole, 780 Asthma, 220, 223 ovalbumin model, 223 ASTM International, 1368 Astrocytes, 1200e1201 Astrogliosis, 1200 Asymmetric self-renewal, 281e282 AT-1. See Acetylated tubulin-1 (AT-1) AT1. See Alveolar type 1 pneumocytes (AT1) ATHENA. See Advanced TissueEngineered Human Ectypal Network Analyzer (ATHENA) Atherosclerosis, 391 ATMP. See Advanced therapy medicinal product (ATMP) ATN. See Acute tubular necrosis (ATN) Atoh1expression, 872, 874 Atomic force microscopy (AFM), 527 Atorvastatin, 775e776 ATP. See Adenosine triphosphate (ATP) ATPase. See Adenosine triphosphatase (ATPase) AuNPs. See Gold nanoparticles (AuNPs) Autacoids, 888 Autism spectrum disorder (ASD), 161e162 Autografting, 591 Autoimmune diseases, hematopoietic stem cell transplantation for, 199 Autoimmune responses, 717 Autoimmunity, 689 Autoimmunity-Blocking Antibody for Tolerance trial (AbATE trial), 999e1000 Autologous arteriovenous (AV), 1029 Autologous cells, 664, 1251, 1264 cell-based MSC therapy, 222 sheets, 473 grafts, 795

1380 Autologous (Continued ) growth factors in hair follicle regeneration, 1298e1299 immune response, 175 lipotransfer, 300 peripheral blood stem cell transplantation, 196 somatic cells, 1264 stem cell therapies, 1242 tissue grafts, 1224 transplantation, 1252e1253 of cryopreserved ovarian tissue, 1243 vein grafts, 1224e1225 Autologous chondrocyte implantation (ACI), 937e938, 956e957 Autologous hematopoietic stem cell transplantation (AHSCT), 996 Automation, 1369e1370 Autosomal recessive mutation, 354 AV. See Arteriovenous (AV); Autologous arteriovenous (AV) Avidity, 3 Avotermin, 75 Axin, 52e53 Axon regrowth, guiding, 1207e1208 Az-chitosaneQHREDGS, 1090 Azathioprine, 987e988, 999e1000

B

B lymphocytes, 685 B-cell antibody production, 715 antigen receptors, 198 hyperreactivity, 688e689 in vitro effects of MAPCs on, 184 B-strand peptides, 646 Bacterial immune process, 741 BAL. See Bioartificial liver (BAL); Bronchoalveolar lavage (BAL) Balb/C control mice, 690 “Bands of engraftment” pattern, 976, 978f Barth syndrome, 1081 Barx-1 gene, 1134 Basal cells, 1283e1284 Basal epidermal cells, 38 Basal lamina, 662e663 invasion of, 3e4 Basement membrane, 613e615 Basic fibroblast growth factor (bFGF), 114, 311e313, 619, 667, 668f, 703, 914, 1087e1088, 1135, 1185e1186, 1228, 1241, 1298e1299 Basic helix-loop-helix family (bHLH family), 275 Basic multicellular units (BMU), 423 Basilar papilla (BP), 869 BAT. See Brown adipose tissue (BAT) BBB. See Bloodebrain barrier (BBB) BDNF. See Brain-derived neurotrophic factor (BDNF) bFGF. See Basic fibroblast growth factor (bFGF) BFU-E. See Burst-forming unitseerythroid (BFU-E)

INDEX

BGP. See Bone g-carboxyglutamic acid containing Gla protein (BGP) bHLH family. See Basic helix-loop-helix family (bHLH family) Biliary tree, 341e343 Biliary tree stem cells (BTSCs), 341 Bilirubin, 236 Bimodal distribution, 1180 Binding energy (EB), 525 Bio-Oss scaffold, 897e898 Bioactive agents, 911 controlled release of, 510e511 affinity-based release, 510 on-demand release, 510e511 porous structures effect, 510 Bioactive bioceramics, 550 Bioactive ceramics, 705 Bioactive glass, 553 Bioactive hydrogels, 763e764 Bioactive molecules, 859, 892e893 BMPs, 893 PDGF, 893 Bioactive signals, 1269 Bioactive substances, 575 Bioactive substrates, 442 Bioactive surfaces development, 442e446 cellebioactive surface interactions, 443e446 Bioactivity, 697 Bioartificial intestinal segments (BIS), 1136 Bioartificial liver (BAL), 231 and transplantation research, 1107e1108 Bioartificial renal epithelial cell system (BRECS), 1156e1157, 1158f treating acute kidney injury, 1157e1158 Bioassay, 407 Bioceramics, 550 bioactive, 705 CaP, 591e593 inert, 550 natural-based bioceramics, 550 resorbable, 550 Biochemical signaling, 763e764 Biocompatibility, 1088e1090 and bioresponse to biomaterials, 675 fibrosis and fibrous encapsulation, 683e684 immunotoxicity, 684e691 inflammation and wound healing, 676e683 of implanted material, 678 safety and biocompatibility requirements for biomaterial scaffolds, 513e517 foreign body response, 516e517 hemocompatibility, 515 infection and sterilization, 514e515 toxicity, 515 Biodegradability, 697 Biodegradable biomaterials, 1268e1271 advantages and limitations of, 1269t natural collagen matrix, 1270e1271 synthetic scaffolds, 1268e1270 Biodegradable block copolymers, 572

Biodegradable cross-linked polymer networks, 575e580 cross-linked polyesters, 576e580 Biodegradable oligomeric macromers, 579 Biodegradable PLGA microspheres, 690 Biodegradable polymers, 559, 1063e1064, 1169e1170 Biodegradable PUs, 572e573 Biodegradable scaffolds, 907 Biodegradable synthetic polymer, 559e560, 1268e1269. See also Nondegradable synthetic polymers for regenerative medicine, 567e580 amino acidederived polymers, poly(amino acids), and peptides, 573e574 biodegradable cross-linked polymer networks, 575e580 block copolymers of polyesters or polyamides, 572 polyanhydrides, 574e575 polyesters, 567e572 polyphosphazenes, 575 polyurethanes, 572e573 scaffolds, 1139 for structural integrity, 837 Biodegradable synthetic-based scaffolds, 1038 Bioelectricity, 450 Bioengineered kidney-like tissues, 1171 Bioengineered lungs, 793e794, 794f Bioengineered rodent organs, 621e622 Bioengineered sphincter, 1142 Bioengineered teeth, 912 Bioengineering applications, 815 bone, 795 functional lungs, 788e789 human hair follicle, 1304e1306 of liver tissue bioartificial liver and transplantation research, 1107e1108 cancer research, 1106e1107 hepatic tissue engineering, 1103e1106 limitations of current in vitro liver models to test drugs, 1108e1109 liver spheroids, organoids, and aggregates, 1106e1110 toxicology and drug development, 1108 Biofabrication, 772e773 Biogenesis of platelets, 929 Bioglass, 706e707 Bioinert ceramics, 705 Bioinformatic analysis, 171 Bioink(s), 805, 808e826, 835, 955 biodegradable synthetic polymers for structural integrity, 837 bioink materials compatible with printing techniques, 809t categories, 810f chain, 818e819 hydrogel-based bioinks for cell printing, 835e837

INDEX

matrix or matrix-mimicking bioinks, 810e821 printability, 835 scaffold-free cell printing, 838 Biologic scaffold materials, 613 clinical and commercial applications, 622 regulatory considerations for ECMs scaffolds, 622 ECMs, 613e619 ECMecell interactions, 615f intact and solubilized ECMs as scaffold material, 619e622 Biological assays, 540 augmentation, 1190 factors, 942e943 materials, 422 modification of surfaces, 656e659 noise, 100 processes, 633, 636e637, 923 Biological License Agreement (BLA), 150 Biological Product (FDA definition), 1348 Biological proteinepeptide-based nanobiomaterials, 492t, 493 Biological scaffolds, 811, 1188, 1268 to support regeneration, 1063e1065 Biological tissues, 661e662 Biological-based scaffolds, 1030e1033 natural decellularized matrices, 1030e1032 nature-derived polymers, 1032 TESA, 1032e1033 Biologics, 81 Biologics Control Act (1902), 1346 Biologics License Application (BLA), 1348e1349 Biology of ligaments and tendons, 1180 Bioluminescence, 799 Bioluminescence imaging (BLI), 278e279, 278f Biomark chip from Fluidigm Corporation, 97 Biomarker analysis, 383 Biomaterial scaffolds, design principles in function and application-oriented design, 505e513 controlled release of bioactive agents, 510e511 degradation profile, 507e510 mechanical support, 505e506 scaffold morphology, 511e513 traceability and imaging, 513 manufacturability, 517 safety and biocompatibility requirements, 513e517 Biomaterial(s), 31, 559, 642, 651, 675, 690, 695e696, 717e718, 772, 1268e1269, 1281 approach, 717e718, 731 biological modification of surfaces, 656e659 biomedical and biotechnological applications of immobilized biomolecules, 657t

for bioprinting, 835e838 immobilizing biomolecules onto and within biomaterials, 658f interfaces in regenerative medicine, 651 overcoating technologies, 655e656 physicochemical surface modifications, 653e655 surface chemical patterning, 659 surface modification strategies, 651e653 systems as cancer vaccines, injectable, 729e731, 730f for TEHVs decellularized bioscaffolds, 1046e1047 future direction in TEHVs, 1052e1053, 1054t hydrolytically degradable polymers, 1050e1051 natural materials for TEHVs, 1047e1049 PVA, 1049e1050 synthetic biomaterials, 1049 TEHvs. fabrication techniques, 1051e1052 templates, 628e630 Biomechanics, 1045, 1180e1183 contribution to joint function, 1182e1183 uniaxial tensile testing, 1181e1182 Biomedical application, 806 thermoresponsive polymer for, 469e470 Biomedical Pus, 567 Biomedical research, 1318e1320 Biomer, 567 Biomimetic(s) biomaterials, 405e406, 409 deposition method, 551e552 hydrogels, 638 lung-on-a-chips, 792 mechanical stimuli, 457 strategies, 691 surfaces, 560 Biomineralization, 761e762 bone tissue engineering growth and differentiation factors in, 859e861 principles of, 854 scaffolds for, 856e859 stem cells in, 854e856 development and fracture of bone, 853 immunomodulation in bone regeneration, 861e862 Biomolecular factors, 667 Biomolecular orientation control, streptavidin for, 530 Biomolecule(s), 656, 667, 670 conformational stabilization for, 527 delivery, 1201, 1206e1207, 1210e1211 “Bioorthopedic” company, 224 Biophysical cues, 280e281 Bioprinted ear construct, 842, 845f Bioprinted organized muscle construct, 845, 845f Bioprinted tracheal construct, 845f, 847

1381 Bioprinting, 544, 832e835, 846, 1078e1080, 1079f, 1286, 1369. See also Threedimensional bioprinting (3D bioprinting) alginate, 815e816 applications, 811 in vitro biological systems, 840t approach, 1051e1052 biomaterials for, 835e838 extrusion-based printing, 834 hybrid and other mechanisms, 834e835 jetting-based printing, 832e833 laser-assisted printing, 834 strategy, 831e832 technologies, 772e773 Biopsy, 1251, 1305 Bioreactive molecules, 938 Bioreactor, 455, 457, 897e898, 943e946, 944f, 1046, 1139, 1274 bioreactor-assisted recellularization method, 621e622 clinical translation, 945e946 and conditioning, 1083e1086 differentiation, 1084e1086 oxygen supply, 1083e1084 development, 1286 in regenerative medicine bone bioreactors, 795e799 challenges and future directions, 801 design considerations for creating bioreactors, 787e788 lung bioreactors, 788e794 systems, 429 translation of cartilage tissue engineering, 945 Bioscaffold, 613, 621 in cartilage repair, 939e942 natural scaffolds, 940e941 synthetic scaffolds, 941e942 decellularized, 1046e1047 Biosensing systems, 1372 Biotechnological approaches, 523e524 Biowire, 1081 Biphasic CaP compounds, 593 Bipolar cells, 352e353 BIS. See Bioartificial intestinal segments (BIS) Bisulfite, 100 BLA. See Biological License Agreement (BLA); Biologics License Application (BLA) Black box process, 169e170 Bladder, 1264 atrophy model, 1271e1272 biodegradable biomaterials, 1268e1271 cell sources, 1264e1268 clinical trials, 1273e1275 clinical studies, 1274e1275 clinical translation, 1273e1274 dystrophy model, 1273 muscle, 1273e1274 preclinical models experimental animal models for bladder reconstruction, 1272t fibrotic bladder model, 1271e1273

1382 Bladder (Continued ) tissue engineering techniques for bladder regeneration, 1273t tissue regeneration models, 1271 Bladder submucosa (BSM), 1270 Blastema cell migration and accumulation, 41e42 cell proliferation, 37 development, 37 formation, 37e42 cell cycling during, 40e41 differential tissue contributions to blastema, 40 hemostasis and reepithelialization, 38e39 histolysis and dedifferentiation, 39e40 macrophages, 41 growth, 42e45 AECenerve interaction, 42e44 cells interaction from opposite sides of limb circumference, 44e45 Blastocyst, 114e115 Bleomycin, 78, 778e779 BLI. See Bioluminescence imaging (BLI) b-Blocker, 780 Blood, 753e754 bloodematerial interactions, 676e677 clot, 1183e1184 substitutes, 923e933 transfusion, 924 vessel, 313, 317, 424 bioreactors, 431e432 blood vesseleassociated mesoangioblasts, 285e286 Blood and Marrow Transplant Clinical Trials Network, 156 Blood urea nitrogen (BUN), 143, 1152e1153 Blood-retinal barrier (BRB), 1200 Blood-spinal cord barrier (BSCB), 1201 Bloodebrain barrier (BBB), 370e371, 1200 permeability, 373e375 BM. See Bone marrow (BM) BM-MNCs. See Bone marrow mononuclear cell (BM-MNCs) BM-MSCs. See Bone marrow-derived mesenchymal stem cells (BMMSCs) BM-SCs. See Bone marrowederived stem cells (BM-SCs) BMAC technique. See Bone marrow aspirate concentrate technique (BMAC technique) BMDCs. See Bone marrowederived cells (BMDCs) BMF. See Bone marrow failure (BMF) BMP. See Bone morphogenetic protein (BMP) BMP receptor IA (BMPR-IA), 409 BMP receptor IB (BMPR-IB), 409 BMSCs. See Bone marrow stem/stromal cells (BMSCs) BMU. See Basic multicellular units (BMU) Body-on-a-chip, 769e770, 775e783

INDEX

advance of in vitro organoid development, 770e771 multiorgan systems and future applications, 775e783 organ-on-a-chip technologies and applications, 772e775 personalized medicine systems, 1372 perspectives, 783 Bolus delivery of immunomodulatory agents, 721e722 Bombyx mori, 1125 Bombyx mori. See Silkworm (Bombyx mori) Bombyx mori (B. mori), 645, 699, 1271 Bone, 417, 841e842, 955e956 adaptation, 423e424 bioreactors, 429e430, 795e799 bioengineering bone, 795 monitoring environment and tissue development, 799 nonperfused bioreactors, 795 perfusion bioreactors, 795e797 for studying bone development and disease, 798e799 for solving vascularization problem, 797e798, 798f bone development, 853 development, 427, 853 bioreactors for studying bone development and disease, 798e799 fracture healing, 853 fundamentals of bone development and defects, 696 grafting, 406, 854 induction, 406 injuries, 535 matrix, 764 minerals, 493 segments, 791e792 substitute, 591, 593 tissue, 493e494, 761, 795 Bone marrow (BM), 181e182, 205e206, 248e250, 253e254, 295, 307, 923e925 aspirate, 894 BM-derived muscle stem cells, 296 cells, 240 mobilization of, 308e310 hematopoiesis, 193 niche, 193 safety of BM cell infusion, 195e196 stem cells, 257 stromal cells of, 315 transplantation, 195e196 Bone marrow aspirate concentrate technique (BMAC technique), 892, 894e897 Bone marrow failure (BMF), 156 Bone marrow mononuclear cell (BMMNCs), 257, 345, 375e376, 381, 1035e1036 adult trial, 382, 384 biomarker analysis, 383 imaging data, 382e383 pediatric trial, 381e382 longitudinal outcome measures, 381

rationale for using BM mononuclear cells, 381 reduction in therapeutic intensity, 381 Bone marrow stem/stromal cells (BMSCs), 378, 699e700, 1264, 1266 Bone marrow-derived mesenchymal stem cells (BM-MSCs), 83, 135e136, 206, 700e701, 795, 909e910, 955, 1168e1169, 1241 Bone marrowederived cells (BMDCs), 1187 Bone marrowederived hematopoietic stem cells, 351, 361 Bone marrowederived stem cells (BMSCs), 362, 1204 Bone morphogenetic protein (BMP), 6, 49e52, 248e250, 339e340, 405e409, 406f, 411fe412f, 523, 696, 853, 859e860, 892e893, 911, 957, 1300 BMP-2, 205, 409, 427, 572, 763e764, 940 BMP-4, 701 BMP-7, 763e764, 897e898, 1172e1173 Bone regeneration, 761e762, 841e842, 858, 894. See also Hair cell regeneration; Peripheral nerve regeneration bone tissue engineering growth and differentiation factors in, 859e861 principles of, 854 scaffolds for, 856e859 stem cells in, 854e856 development and fracture of bone, 853 immunomodulation, 861e862 nonperfused bioreactors for, 795 perfusion bioreactors for, 795e797 surface modification and functionalization of scaffolds for, 859 Bone sialoprotein (BS), 699e700, 855t, 909e910 Bone tissue engineering (BTE), 761, 795. See also Dental tissue engineering; Functional tissue engineering (FTE); Hepatic tissue engineering cells in, 762 growth and differentiation factors, 859e861 bone morphogenetic proteins, 859e860 nucleotide delivery and gene therapy, 860e861 PTH delivery, 860 nanofibrous scaffolds, 857 natural origin materials calcium phosphates, 551e553 chitosan, 538e540 collagen, 540e543 gellan gum, 543e545 natural-based bioceramics, 550 natural-based polymers, 538 PHAs, 545e546 silicate ceramics, 553e554 silk fibroin, 547e548 starch, 548e550 preclinical models

INDEX

selection considerations based on animal species, 765e766 in vitro preclinical models, 764 in vivo preclinical models, 764e765 principles, 854 scaffolding approaches in, 697e699 hybrid materials, 698 hydrogels, 698e699 immunomodulatory materials, 698 osteoinductive materials, 698 scaffolds for, 856e859 composite materials for bone tissue engineering scaffolds, 858 injectable scaffolds, 858e859 nanofibrous scaffolds for bone tissue engineering, 857 porous and highly interconnected scaffolds, 856 scaffolding design criteria, 856 surface modification and functionalization, 859 three-dimensional printed scaffolds, 859 stem cells in, 854e856 ADSCs, 854e855 ESCs, 855 iPSCs, 855e856 MSCs, 854 Bone g-carboxyglutamic acid containing Gla protein (BGP), 406 Bone-forming cells, 764, 894 Boneeligament-bone complex, 1181 BoneePTebone (BPTB), 1179, 1184 Bony labyrinth, 867 Boston KPro, 1116 Bound delivery systems, 763e764 Bovine collagen, 616, 643 Bovine serum albumin (BSA), 442 Bowman capsule, 1167 BP. See Basilar papilla (BP) BPTB. See BoneePTebone (BPTB) Brain death, 127e128 injury phases primary vs. secondary brain injury, 370e376 Brain-derived neurotrophic factor (BDNF), 911, 1201, 1227e1228 BRB. See Blood-retinal barrier (BRB) BrdU. See Bromodeoxyuridine (BrdU) Breakthrough Devices Program, 1374 Breakthrough Therapy designation, 1369e1370 Breast cancer, 6e7, 108, 398, 775 BRECS. See Bioartificial renal epithelial cell system (BRECS) BRG1 chromatin remodeling factor, 57 Brittle ceramics, 601e602 Bromodeoxyuridine (BrdU), 376, 1167, 1300 Bronchoalveolar lavage (BAL), 1154e1155 Brown adipose tissue (BAT), 297 Brown algae, 640 Brushite cements, 594e595, 601e602 BS. See Bone sialoprotein (BS)

BSA. See Bovine serum albumin (BSA) BSCB. See Blood-spinal cord barrier (BSCB) BSM. See Bladder submucosa (BSM) BSP. See Bone sialoprotein (BS) BTE. See Bone tissue engineering (BTE) BTSCs. See Biliary tree stem cells (BTSCs) Bulge cells, 1297 BUN. See Blood urea nitrogen (BUN) Burst-forming unitseerythroid (BFU-E), 924 Bxb1 phage integrases, 748

C

“C-G” peptide linker, 1123e1124 c-Kit expression, 248e250 c-Kit+ cardiac progenitor/stem cells, 248e250, 263 c-Myc genes, 40, 50, 54, 117, 181, 1168 C-reactive protein (CRP), 1155e1156 Ca10[PO4]6[OH]2. See Hydroxyapatite nanoparticles (HAP nanoparticles) Ca18Mg2[HPO4]2[PO4]12. See Whitlockite CABG. See Coronary artery bypass grafting (CABG) Caco-2 cells, 1139 CAD. See Computer-aided design (CAD); Coronary artery disease (CAD) Cadaveric islet transplantation, 994 Cadaveric nerves, 1232 Cadherins, 2, 772 cadherinecatenin complexes, 398 extracellular domains, 398 switching, 2 Cadmium selenide (CdSe), 490e491 Cadmium sulfide, 490e491 Caenorhabditis elegans (C. elegans), 393, 426 CaHPO4. See Dicalcium phosphate anhydrous (DCPA) Calcium phosphate (CaP), 550e553, 551f, 591, 699e700, 955 bioceramics, 591e593 in bone tissue engineering applications, 552e553 CaP-based ceramics, 891e892 compounds and major properties, 592t CPCs, 593e594 classes of CPCs, 594e595 dental applications, 607 oral, maxillofacial, and craniofacial applications, 606 orthopedic applications, 607 physiochemical properties cohesion, 601 injectability, 599e600 setting/hardening mechanism, 595e599 processing methods, 552 strategies to improving mechanical properties, 601e606 dual setting system, 604e605 fiber reinforcement, 605e606 porosity, 602e604 2D solubility phase diagram for CaP compounds, 592f

1383 Calcium phosphate cements (CPCs), 593e594, 915 basic properties, 593e594 classes, 594e595 apatite cements, 594e595 brushite cements, 595 Calcium phosphate ceramics (CPCs), 540, 706 particles, 656 Calcium polyphosphate scaffolds, 512 Calcium polyphosphate-based bioceramic scaffolds (CPP-based bioceramic scaffolds), 709 Calcium-deficient HA (CDHA), 592e594 Californian Institute for Regenerative Medicine (CIRM), 1316 CAM processes. See Computer-aided manufacturing process (CAM processes) cAMP. See Cyclic adenosine monophosphate (cAMP) Canadian Institutes of Health Research (CIHR), 1320e1321 Cancellous bone, 696 Cancer, 400, 776, 1263 cells, 715 diseases, 405e406 metastasis, 776 nanomedicine in, 718e719 research, 1106e1107 stem cells, 108 Cancer immunotherapy, 715e717, 947 advantages and disadvantages, 717e718 macroscale biomaterial scaffolds, 727e733 to enhancing autologous T cell therapy, 731e733 injectable biomaterial systems as cancer vaccines, 729e731 nanoparticle biomaterials, 718e727 nanomedicine applications, 727 nanoparticle targeting applications, 721 Cancer vaccine, 716 implantable biomaterial scaffolds as, 727e729, 728f injectable biomaterial systems as, 729e731, 730f Cancer-on-a-chip, 774e775 CaP. See Calcium phosphate (CaP) CAR T-cells. See Chimeric antigen receptor T-cells (CAR T-cells) Carbachol, 1272e1273 Carbodiimide chemistry, 442 Carbon based nanobiomaterials, 492t, 493 Carbon nanotubes (CNTs), 450e451, 489, 647e648 Carbonyl diimidazole chemistry, 529 Carbopol, 823e824 N-Carboxy-anhydrides (NCAs), 574 Carboxylated phosphorylcholine, 573 5-Carboxylcytosine, 57 Carboxylic acid, 568 polyesters of, 571 Carboxymethyl cellulose, 601

1384 1,3-bis(p-Carboxyphenoxy)propane (CPP), 574 Carcinogenesis, 301e302 Cardiac anlagen, 247 Cardiac engraftment of cells, 253e254 Cardiac muscle, 494 Cardiac myocytes, 65, 248 Cardiac neural crest cells (CNC cells), 247e250 CNCederived progenitors, 250e251 Cardiac organoids, 779e780, 1080e1082, 1082f Cardiac output (CO), 1157e1158 Cardiac patches implantation, 1091e1093 scaffolds, and bioreactors cardiac patches engineering, 1074e1077 Cardiac precursor pathways, 247 Cardiac progenitor cell (CPC), 209e210 CPC-derived exosomes, 209e210 Cardiac regeneration, 209e210 Cardiac Repair Cell (CRC), 265e266 Cardiac stem cells (CSC), 261e262. See also Amniotic fluid stem cells (AFS cells); Embryonic stem cells (ESC); Human embryonic stem cells (hESC); Induced pluripotent stem cells (iPSCs); Mesenchymal stem cells (MSCs) in adult heart, 248e252 c-Kit+ cardiac progenitor/stem cells, 248e250 cardiac neural crestederived progenitors, 250e251 epicardial progenitor cells, 251e252 Isl1+ cardioblasts, 250 cell-based therapeutics for heart disease, 252e254, 253f clinical trials, 255e264 c-Kit+ CSCs, 263 cardiopoietic stem cells, 255e262 other CSCs, 263e264 combined stem cell therapeutics, 265e266 development of heart from cardiac stem/progenitor cells, 247e248 mechanisms of action, 254e255 cell-based therapeutic strategies for cardiac repair, 254f methods for expansion of adult, 265 Cardiac stem cells in patients with ischemic cardiomyopathy trial (SCIPIO trial), 263 Cardiac stem/progenitor cells, heart development from, 247e248 Cardiac tissue, 450, 846 bioprinting, 1078e1080, 1079f bioreactors and conditioning, 1083e1086 cardiac organoids and organ-on-a-chip engineering, 1080e1082 cardiac patches engineering using cells, scaffolds, and bioreactors, 1074e1077

INDEX

critical issues associating with tissue engineering heart, 1074t engineering culture systems, 1084f engineering ventricle, 1083 goals and issues, 1073 native heart extracellular matrix, 1075f tissue and organ function, 1086e1090 host response and biocompatibility, 1088e1090 mechanical elasticity and strength development, 1086 microfabrication of vasculature, 1089f thrombogenicity and endothelialization, 1087 tissue architecture and electrical conduction, 1086e1087 vascularization, 1087e1088 in vivo studies, 1090e1094 Cardiac-committed cells, 255 Cardiogenic mesoderm, 247e248 Cardiomyocyte, 175, 1075e1076, 1080 mitosis pathways, 247 renewal, 248 in adult humans, 249f Cardiomyogenesis, 250e251 Cardiopoietic cells, 252 Cardiopoietic stem cells adult stem cells, 256e257 BM stem cells, 257 BMMNCs, 257 EPC, 257 MSCs, 258e261, 259f myoblasts, 261, 263f pluripotent stem cells, 255e256 Cardiospheres (CSs), 265 CS forming cells, 265 Cardiovascular disease (CVD), 209e210, 255, 314, 318e319, 1029. See also Heart Cardiovascular tissue, 431, 495 Carolinas Cord Blood Bank (CCBB), 151 Cartilage, 405e406, 405f, 424, 842, 937, 956 bioreactors, 429 morphogenetic proteins, 412 regeneration, 474 surface modification, 938e939 TE for cartilage repair, 938e946 tissue, 496 Cartilage tissue engineering. See also Bone tissue engineering (BTE); Dental tissue engineering for cartilage repair, 938e946 trends, 946e947 Cartilage-derived morphogenetic proteins (CDMPs), 405e406 Cas protein, 742, 744 Cas9 orthologue, 749e750 Cas9 system. See CRISPR-associated protein 9 system (Cas9 system) Catecholamine phenotype, 173 b-Catenin (b-cat), 6e7, 52e53, 874 Catgut sutures, 642e643 Catheter, 1150 Cationic lipids, 749 CB. See Cord blood (CB)

CBAVD. See Congenital bilateral absence of the vas deferens (CBAVD) CBD-VEGF. See VEGF fused with collagenbinding domain (CBD-VEGF) CBER. See Center for Biologics Evaluation and Research (CBER) CBP. See Cyclic adenosine monophosphate response element binding protein (CBP) CBT. See Cord blood transplantation (CBT) CBUs. See Cord blood units (CBUs) CC. See Corpus callosum (CC) CeC-chemokine receptor 2 (CCR2), 373 CC10. See Club cell secretory protein, 10 kD (CC10) CCBB. See Carolinas Cord Blood Bank (CCBB) CCI. See Controlled cortical impact (CCI) CCL20. See Chemokine (CeC motif) ligand 20 (CCL20) CCR2. See CeC-chemokine receptor 2 (CCR2) CCRM. See Centre for Commercialization of Regenerative Medicine (CCRM) CD105 marker. See Endoglin CD105-based immunoisolation method, 205 CD106 marker. See Vascular adhesion molecule-1 CD11a/CD18 marker, 681 CD11b marker, 277 CD11b/CD18 marker, 681 CD11c/CD18 marker, 681 CD133 stem cell marker, 1167 CD146 marker, 206 CD24 marker, 1135 CD25 T cell markers, 682 CD26 marker, 232 CD31 marker, 206, 277 CD34 marker, 194e195, 311 CD34+ surface marker, 154 CD36-thrombospondin binding, 18 CD40 ligand expression, 688 CD41. See Cell surface markers GPIIb/IIIa CD44 receptor, 18 CD45 markers, 206, 277 CD69 T cell markers, 682 CD8+ DCs, 727e729 CD80 molecule, 1268 CD86 molecule, 1268 CD9e/e mice, 210e211 CDC. See Centers for Disease Control and Prevention (CDC) CDCs. See CS-derived cells (CDCs) CDER. See Center for Drug Evaluation and Research (CDER) CDHA. See Calcium-deficient HA (CDHA) CDK. See Cyclin-dependent kinase (CDK) CDMPs. See Cartilage-derived morphogenetic proteins (CDMPs) cDNA. See Complementary DNA (cDNA) CDRH. See Center for Devices and Radiological Health (CDRH) CdSe. See Cadmium selenide (CdSe)

INDEX

CEA. See Cultured epithelial autograft (CEA) CECs. See Circulating endothelial cells (CECs); Corneal endothelial cells (CECs) Cefazolin sodium, 898 Cell adhesion, 443e444, 456, 940 and detachment, thermoresponsive surface, 470 molecules, 763e764 Cell administration, 976e980 approaches for intramuscular transplantation, 976 density of cell injections, 976 efficiency of cell injections, 979e980 potential risks of cell injection procedure, 977e979 Cell and Gene Therapy Catapult network, 1375 Cell injection density, 976 efficiency, 979e980 procedure, 977e979 strategies, 1074e1075 Cell polarity, changes in, 3 Cell printing hydrogel-based bioinks for, 835e837 scaffold-free, 838 Cell sheet engineering clinical applications for, 472e474 cartilage regeneration, 474 cornea reconstruction, 472 esophagus reconstruction after ESD treatment, 473 myocardium regeneration, 473e474 PDL regeneration, 474 combination with scaffold-based engineering, 477e478 intelligence of thermoresponsive polymers for, 469e471 microfabricated intelligent surface, 478e481 produces scaffold-free, 3D tissue constructs, 474e477 Cell sheet layering technique, 474e475, 475f 3D coculture system based on, 475 3D orientation arrangement using, 480 Cell sheet technology, 495 Cell signaling, 391 growth factors and Cell signaling molecules, 80e81 Cell source, 297, 663e664, 1043e1044, 1073 , 1074t, 1166e1169 bladder and ureter cells, 1264 cells in BTE, 762 mechanism of cell therapy, 1266e1268 stem cell sources, 1264e1266 Cell surface annexin II, 18 antigen CD31, 312 markers GPIIb/IIIa, 929 proteoglycans, 442 receptors, 18 Cell survival in recipient, 982

Cell therapy, 175e176, 627, 971, 1165, 1187, 1202e1203, 1368 applications, 318e319 for blood substitutes HSCs, 931e933 megakaryocytes and platelets, 929e931 perspectives, 933 red blood cells, 924e929 cell expansion, 1266 of liver disease choice of sites for hepatocyte transplantation, 231 clinical hepatocyte transplantation, 232e237, 233t current treatments for, 230t hepatocyte transplantation, 237e241 integration of hepatocytes after transplantation, 231e232 multipotentiality, 1266e1267 paracrine effects and immunomodulatory properties, 1267e1268 Cell transplantation, 230, 232e233, 237, 335e336, 971, 1205e1206, 1208, 1211e1212. See also Islet cell transplantation immunology, 239 in skeletal muscle, 972e975 Cell-adhesive ligands, 638 Cell-assisted lipotransfer, 300 Cell-based neuroprotection, 361e362 bone marrowederived stem cells, 362 neural and retinal progenitor cells, 362 umbilical tissue-derived stem cells, 362 Cell-based therapeutics, 247 for heart disease, 252e254, 253f Cell-based therapy, 205, 469, 473e474, 953e954, 956e957, 1166e1172 cell sources, 1166e1169 development, 104e105 engineering of cell-based renal constructs, 1169e1172 kidney regeneration, 1166t for lumbar degenerative disc disease, 962e963 Cell-based tissue engineering, 1269 Cell-binding peptides, 667 Cell-free approach, 1172e1173 Cell-free biomaterials, 1122e1124 collagen-based implants, 1123 decellularized extracellular matrix as implants, 1122e1123 peptide analogs of extracellular matrix, 1123e1124 Cell-free method, 1116 Cell-free organ, 453e454 Cell-free seeded scaffolds, 1271 Cell-graft survival ensuring cell survival in recipient, 982 initial survival, 980e981 long-term survival, 981e982 Cell-homing strategies, 701 Cell-laden bioinks, 819e821, 820f

1385 current translation of three-dimensional bioprinting, 825e826 sacrificial bioinks, 821e823, 822f supporting bioinks and supporting baths, 823e825 Cell-laden GelMa, 836e837 Cell-laden scaffold, 810 Cell-mediated delayed hypersensitivity reaction, 686 Cell-penetrating peptides, 749 Cell-replacement therapy, 355e361 human embryonic stem cellederived retinal pigment epithelium, 355e356 induced pluripotent stem cellederived retinal pigment epithelium, 356e359 photoreceptor transplantation, 360e361 scaffolds for retinal pigment epithelium transplantation, 359e360 surgical techniques for retinal pigment epithelium transplantation, 360 Cell-replacement therapy. See also Extracorporeal renal replacement Cell-seeded, longitudinally aligned nerve guidance conduits and channels, 1231 Cell-seeded scaffolds, 1255, 1271 Cell(s), 231, 1368 behaviors, 788 culture, 523e524, 1033e1034 cycling during blastema formation, 40e41 death postthaw, 1156 division, 22 durability, 1073 encapsulation techniques, 1243e1244 expansion, 1266, 1367e1368 function, 1073 fusion, 240 granules, 29 of hematopoietic origin, 1284 infiltration, 697 infusions, 229e230 of innate immune system, 715 interaction, 438 isolation, 206, 238, 1167 migration, 456e457 morphology, 456 motility, 444e445 stimulation, 3 movement, 440 of MPS, 720 number, 1073 to organs, 399e400 phenotypes, 1060e1061 potency assays, 1302 proliferation, 445e446 senescence, 41 shape, 446 signals, 763e764 structure and composition, 392, 392f and tissue expansion systems, 1372 types, 240, 376 for cartilage repair, 939

1386 Cellebioactive surface interactions, 443e446 cell adhesion, 443e444 cell motility, 444e445 cell proliferation, self-renewal, and differentiation, 445e446 Cellebiomaterial composites, 1125 Cellecell adhesion, 398e400 from cells to organs, 399e400 changes in, 2 Cellecell interactions, 478 CelleECM adhesion, 398 changes in, 3 CelleECM interactions, 15, 19f, 28f, 437e438, 445 composition and diversity of ECM, 15e16 during healing of cutaneous wounds, 25e28 implications for regenerative medicine, 30e31 receptors for ECM molecules, 16e18 during regenerative fetal wound healing, 28e30 signal transduction events during, 18e24 CelleECM reciprocity, 457 Cellefibrin hydrogel micromolding approach, 1081 Cellepolymer scaffold, 1273 b Cells, 335 from adult stem/progenitor cells, 341e343 from pluripotent stem cells, 336e341, 337f, 337t, 338f for replacement therapy, 336 Cellescaffold combination products, 1355e1356 CELLstart, 118 Cellesubstrate interactions CelleECM matrix interactions, 437e438 cellular responses to topographical cues, 447e450 chemical properties effect, 440e442 cellular response, 441 methods of altering surface chemistry, 441e442 surface charge, 440 surface wettability, 440 CNT and graphene surfaces, 451 commonly used ligands, 442 development of bioactive surfaces, 442e446 dimensionality effects, 451e460 effect of biological properties, 442 electrically conductive substrate, 450e451 fabrication techniques, 446e447 importance of substrate, 438 physical properties, effect of, 438e440 topography effects, 446 Cellesubstrate interactions, 489 Cellular biology, 391 cardioplasty, 299

INDEX

characterization ASC, 296e297 cell source, 297 SVF, 296 copatterning to create cellular microenvironment, 478 differentiation, 445e446 fractions, 295e296 function, 1373 mathematical identification of cellular subpopulations, 101e103 mechanics, 395e396 mechanotransduction mechanisms, 397e400 through cellecell adhesions, 398e400 through celleextracellular matrix adhesions, 398 metabolism, 1067e1068 polarity, 3 proliferation, 874e875 reprogramming, 55 scale, 506 Cellular, Tissue, and Gene Therapies Advisory Committee (CTGTAC), 1352, 1358e1359 Cellular heterogeneity, 93e94 in aging, 107e108 clinical implications in tissue repair and disease, 105e108 in diabetes, 106 in fibrosis, 107 tumor cell heterogeneity and drug resistance, 108 in wound healing, 106e107 Cellular response, 441 in modifying ECM, 459e460 to three-dimensional substrates, 455e457 cell adhesion, 456 cell migration, 456e457 celleECM reciprocity, 457 to topographical cues, 447e450 Cellularized membrane, 774 Cellulose, 641e642 acetate, 494 derivatives, 1151 Cementum tissues, 907 Center for Biologics Evaluation and Research (CBER), 1347 Center for Cord Blood, 150 Center for Devices and Radiological Health (CDRH), 1347 Center for Drug Evaluation and Research (CDER), 1347 Center for International Blood and Marrow Transplant Research, 150 Center for Veterinary Medicine (CVM), 1354 Centers for Disease Control and Prevention (CDC), 369, 369f, 1351, 1353 Central nervous system (CNS), 176, 320e321, 351, 371e373, 496, 1199, 1229, 1283

case studies in tissue therapy in, 1203e1212 retinal degeneration, 1208e1212 stroke, 1203e1206 traumatic spinal cord injury, 1206e1208 therapeutic strategies in, 1201e1203 biomolecule delivery, 1201 cell therapy, 1202e1203 wound response and barriers to regeneration, 1200e1201 Centre for Commercialization of Regenerative Medicine (CCRM), 1374e1375 Ceramics, 831, 891e892 scaffolds, 507e508, 705e707 bioglass, 706e707 calcium phosphate, 706 coral, 707 Cerebral edema, 375e376 Cerebral infarction, 314 Cerebral palsy, 159e160 Cerebral perfusion pressure (CPP), 369e370 Cerebral vasospasm, 370e371 Cerebrospinal fluid (CSF), 378, 1205e1206 Cerebrovascular disease, 1203 Cethrin, 1201, 1207 CF. See Cystic fibrosis (CF) CFP. See Commissioner’s Fellowship Program (CFP) CFTR. See Cystic fibrosis transmembrane receptor (CFTR) CFU. See Colony-forming unit (CFU) CFU-erythroid (CFU-E), 924 CFU-Fs. See Colony-forming unitefibroblasts (CFU-Fs) CFU-s. See Colony-forming unitespleen (CFU-s) CGI-I. See Clinical Global Impressions scale (CGI-I) CGL. See Chronic granulocytic leukemia (CGL) CGRGDS. See Cyclic Gly-Arg-Gly-Asp-Ser (CGRGDS) CGTP. See Current Good Tissue Practice (CGTP) Chain-growth polymerizations, 560e561 Characteristic interaction parameter, 665 Charge control of orientation, 528e529 Chemical accelerators, 689e690 agents, 749 chemical-field effect transistors, 1159 cross-linked hydrogels, 637 delivery methods, 749 gene delivery methods, 749 groups, 531 mediators, 676, 677t modifications, 653 processing methods, 616 species, 525 Chemisorption, 443 Chemistry, Manufacturing, and Control (CMC), 1349e1350

INDEX

Chemokine (CeC motif) ligand 20 (CCL20), 727e729 Chemokines, 209, 297, 715 Chemometrics. See Multivariate statistical methods Chemoselective reactions, 657e659 Chemotaxis, 406, 679 Chemotherapeutic, 78, 778e779 CHI. See Chitosan (CHI) Chicken-derived erythroid cells, 192 Chickequail chimeras, 191e192 ChildePugh scores, 241 Children’s Oncology Group, 156 Chimeras, human stem cells to creating, 1335e1336 Chimeric animals, 1013 Chimeric antigen receptor T-cells (CAR Tcells), 716 China Regenerative Medicine International, 1123 ChIP-Seq. See Chromatin immunoprecipitation-sequencing (ChIP-Seq) Chitin, 538, 613 and derivatives, 641 a-Chitin, 641 Chitosan (CHI), 538e540, 613, 641, 699, 703, 940e941, 1049, 1103, 1141 in bone tissue engineering applications, 540 CHI-alginate gel-MSC-BMP-2 composites, 702 CHI-alginate hybrid scaffolds, 702 CHI-polypyrrole-alginate composite scaffold, 702 chitosan-based scaffolds, 540 processing methods, 538e540 Chondro-Gide, 616 Chondroclasts, 696 Chondrocytes, 429, 696, 939 Chondrogenesis, 406, 412 Chondroitin sulfate, 618 Chondroitin sulfate proteoglycans (CSPG), 1200, 1231e1232 Chondroitinase ABC, 1203, 1231e1232 Choriocapillaris, 352 Choroideremia, 363 Chromatin immunoprecipitationsequencing (ChIP-Seq), 52 Chromatin structure determining regulatory activity of transcription factor, 58 Chromatography, 209 Chronic bladder diseases, 1266 Chronic bone disorders, 591 Chronic CsA, 1172 Chronic granulocytic leukemia (CGL), 196 Chronic inflammation, 284, 679e680, 688, 1158 Chronic ischemic heart disease, 318 Chronic kidney disease (CKD), 1165 Chronic skin ulcers, 72 Chronic ulcers, 72 Chymotrypsin, 645

CIHR. See Canadian Institutes of Health Research (CIHR) Ciliary neurotrophic factor (CNTF), 184, 1201, 1227e1228 Ceinositol triphosphate (IP3), 425 Circulating endothelial cells (CECs), 314 CIRM. See Californian Institute for Regenerative Medicine (CIRM) Cisplatin, 878 Citric acid, 571, 600 CKD. See Chronic kidney disease (CKD) Clara cells, 142e143 CLARITY technique, 454 Class I transposons, 747 Class II transposons, 747 Class switching, 888 Claudin tight junction gene, 4 Click chemistry, 444, 580, 638, 647, 657e659 Clinical development plan in FDA, 1356e1358 Clinical Global Impressions scale (CGI-I), 162 Clinical islet transplantation, 990e993. See also Cord blood transplantation (CBT); Hepatocyte transplantation; In utero transplantation (IUT) islet transplantation procedure, 991e992 patient assessment and selection, 990e991 risks to recipient immunosuppressive therapy and complications, 992e993 surgical complications, 992 Clinical liver transplantation, 185 Clinical translation, 945e946, 1068, 1338e1340 Clonal derivation, 170e171 Clonal evolution model, 108 Clonal T cell expansion, 724 Cloning human embryos cloning, 1334e1335 reproductive cloning, 1311, 1335 therapeutic, 1311, 1334 Cloning embryos, 1310 Cloning process, 25, 747 CLP. See Collagen-like peptide (CLP) Club cell secretory protein, 10 kD (CC10), 1060e1061 “Cluster defining” genes, 103 Clustered regularly interspaced short palindromic repeats (CRISPR), 174, 285, 363, 741e746, 861, 1019e1020, 1063, 1187 CRISPR-C2c2, 745e746 CRISPR-Cas system, 741, 743 CRISPR/spCas9 system, 742 HDR, 744e745, 744f interference technology, 754 knockouts via double-strand breaks, 742e743 nickases, 743, 743f SpCas 9 variants and orthologues, 744e745 Clustering, 437e438

1387 fuzzy c-means, 103 hierarchical, 101e102, 102f partition, 102 CM. See Conditioned medium (CM) CMC. See Chemistry, Manufacturing, and Control (CMC) CMI. See Collagen Meniscal Implant (CMI) CMP. See Common myeloid progenitor (CMP) CN syndrome. See CriglereNajjar syndrome (CN syndrome) CNC cells. See Cardiac neural crest cells (CNC cells) CNS. See Central nervous system (CNS) CNTF. See Ciliary neurotrophic factor (CNTF) CNTs. See Carbon nanotubes (CNTs) CO. See Cardiac output (CO) Co-based alloys, 707 Co-microencapsulated granulosa cells, 1244 Co-printing bioinks, 816e819, 817f Coacervation, 617 COC. See Cumuluseoophorus complexes (COC) Cochlear epithelium, 874 Cohesion, 601 strategies to improve cohesion, 601 COL2. See Collagen type II (COL2) COL7A1. See Collagen a-1 (VII) chain (COL7A1) Coley’s toxin, 715e716 Collagen, 494, 540e543, 613, 615e616, 700e701, 788, 815, 940, 1103e1104, 1125, 1169, 1226, 1270 in bone tissue engineering applications, 542e543 collagen I, 529 collagen-based hydrogels, 1169 collagen-based inert matrix, 1255 collagen-based scaffolds, 542, 1172 collagen-filled vein grafts, 1224e1226 to control protein orientation, 529 and derivatives, 642e643 fibers, 1180 fibrils, 703e704 gel scaffold system, 1169 processing methods, 541e542 type I, 836 Collagen Meniscal Implant (CMI), 616 Collagen scaffold (CS), 616, 700e701, 1133, 1136, 1241 collagen scaffoldemesenchymal stromal cell study, 385 Collagen type II (COL2), 957 Collagen a-1 (VII) chain (COL7A1), 754 Collagen-based implants, 1123, 1125 Collagen-like peptide (CLP), 1123e1124 Collagenase-sensitive peptide, 573 Collective migration, 1 Colon, 1141 Colony-forming unit (CFU), 152e154, 312e313, 924 Colony-forming unitefibroblasts (CFU-Fs), 206

1388 Colony-forming unitespleen (CFU-s), 191 Combination Product, 1348 ComC. See Complement cascade (ComC) Commercial inkjet printers, 833 Commercialization, 1324e1325 Commissioner’s Fellowship Program (CFP), 1361 Common myeloid progenitor (CMP), 924 Complement cascade (ComC), 308e310 Complement component C5a, 685 Complement-activated fragment (C3b), 679 Complementary DNA (cDNA), 95e97, 752 Compliance, 439e440 Composite composite sponge-like hydrogels, 553e554 implants incorporating specific bioactive functions, 1125 scaffolds, 708 tissues, 847 Comprehensiveness, 95 Compressive strength of CPCs, 601e602 Computed tomography (CT), 370, 513, 831e832, 893, 1046 Computer numeric code machining, 1157 Computer-aided design (CAD), 453, 552, 899, 955, 1083 Computer-aided manufacturing process (CAM processes), 541, 831e832, 832f, 834 Conditioned medium (CM), 210, 1299, 1301 for hair growth, adipose-derived stem cells use and, 1299 Cone beam computed tomography, 702 Conformational stabilization for biomolecules, 527 Congenital bilateral absence of the vas deferens (CBAVD), 1255 Congenital disorders, 1263 Congenital heart disease, 1041e1042 Congestive heart failure cardiopoietic regenerative therapy study, 259e261 Connective tissue growth factor (CTGF), 75e76 Connexins (Cx), 81 connexin 32, 232 connexin 43 proteins, 1079, 1085e1086 Consensus standard, 1357 Constitutive relations, 421e423 Contact guidance, 444, 634e635, 1192, 1287e1288 Contigen, 616 Continuous intrathecal infusion, 1207 Continuous venovenous hemofiltration (CVVH), 1153e1154 Contraceptive hormones, 561 Contractile cells, 23e24 myocytes, 440 myofibrils, 273 SMC markers, 1267

INDEX

Contractures, 65 Controlled cortical impact (CCI), 371 CCI injury model, 373 Conventional approaches, 591 Conventional cell delivery routes, 378e380 pulmonary “first-pass” effect, 378 Conventional fabrication methods, 511 Conventional FACS, 99 Conventional membranes, 1152 Conventional scaffolds, 628e629 Conventional soft-lithography, 1088 Conventional therapies, 1149 Conversion coatings, 653 Copatterning to create cellular microenvironment, 478 Copolymers, 704e705 Copper-containing mesoporous bioactive glass (Cu-MBG), 709 Coral, 707 Cord blood (CB), 924e925 banking, 150 CB-derived microglial-like cells, 158e159 expansion technologies, 157 off-the-shelf therapy, 160e161 stem cells clinical uses of umbilical cord blood, 156e157 cord blood banking, 150 distributions of quality variables, 151f history, 149e150 investigations in treatment of acquired brain injuries, 159e162 public CB banking procedures, 151e156 public vs. family banks, 150e151 therapies for inherited and acquired brain diseases, 157e159 Cord blood transplantation (CBT), 149, 197. See also Hepatocyte transplantation; In utero transplantation (IUT); Islet cell transplantation for hematological malignancies, 156 for IMD, 157e159 for nonmalignant hematological diseases, 156e157 Cord blood units (CBUs), 149 characterization, 154e156 Cornea cell-free biomaterials, 1122e1124 cellebiomaterial composites, 1125 challenges, 1125e1126 composite implants incorporating specific bioactive functions, 1125 fully cell-based, self-assembled corneal constructs, 1119e1122 reconstruction, 472 regeneration of corneal layers, 1118e1119 regenerative medicine applying to keratoprosthesis development, 1116e1118 structure and function, 1115

treatment options, state of art, and need, 1115e1116 0-Cornea, 1122e1123 Corneal blindness, 1115 Corneal endothelial cells (CECs), 449, 1118e1119 Corneal endothelium, 1118e1119 Corneal epithelium, 1118 Corneal prostheses, 1116 Corneal stroma, 1119 Coronary artery bypass grafting (CABG), 261e262, 1029 Coronary artery disease (CAD), 1029 Coronary heart disease, 252e253 Corpus callosum (CC), 382e383 Corrosion process, 507e508 Corti organ. See Mammalian auditory sensory epithelia Cortical bone, 696 Corticospinal tract (CST), 382e383 Corticosteroids, 239, 987e988, 1210e1211 CosmoDerm, 616 CosmoPlast, 616 Cost-effective manufacturing, 1367 Costimulatory molecules, 716e717 Covalent coatings, 656 Covalent cross-linking mechanisms, 640, 648 COX-2+. See Cyclo-oxygenase-2 (COX-2+) CPC. See Cardiac progenitor cell (CPC) CPCs. See Calcium phosphate cements (CPCs); Calcium phosphate ceramics (CPCs) Cpf1 system, 745 CpG. See Cytosine guanine (CpG); Cytosine-guanosine oligodeoxynucleotides (CpG) CPP. See 1,3-bis(p-Carboxyphenoxy) propane (CPP); Cerebral perfusion pressure (CPP) Crack deflection, 606 Craniofacial regenerative medicine, 887 craniofacial regenerative environment, 887e890 current methods of maxillofacial reconstruction, 890e891 tissue engineering technologies, 891e899 Craniofacial region, 887 Craniofacial surgery, 891e892 CRC. See Cardiac Repair Cell (CRC) Cre enzymes, 747e748 Cricopharyngeal muscle, 971 CriglereNajjar syndrome (CN syndrome), 235 CRISPR. See Clustered regularly interspaced short palindromic repeats (CRISPR) CRISPR-associated protein 9 system (Cas9 system), 174, 285, 750, 752e755, 861, 1336 CRISPR-RNA (crRNA), 741e742 Critical path initiative, 1359 Critical Path Opportunities List, 1359 Critical-size defects, 854

1389

INDEX

Crohn disease, 223 Cross anastomosis model, 1232 Cross-linked/cross-linking, 631, 835 hydrophilic polyesters, 579 hydrophilic polymer chains, 631 PEG, 669e670 polyesters, 576e580 synthetic polymers, 576 Cross-sectional area (CSA), 1181 CRP. See C-reactive protein (CRP) crRNA. See CRISPR-RNA (crRNA) Crumbs3 polarity gene, 4 CryoArtery. See Cryopreserved arteries (CryoArtery) Cryopreservation, 153e154 of EP-isolated REC, 1156 of organoid units, 1138 Cryopreserved arteries (CryoArtery), 1032 Cryopreserved ovarian tissue, autologous transplantation of, 1243 Cryotherapy, 79 CryoVein, 1032 Crypt-villus microenvironment, 1139 Cryptorchidism, 1252 Crypts, 1135 Crystalline polymers, 439 Crystallinity, 439 CS. See Collagen scaffold (CS) CS-derived cells (CDCs), 265 CSA. See Cross-sectional area (CSA) CsA. See Cyclosporin A (CsA) CSC. See Cardiac stem cells (CSC) CSF. See Cerebrospinal fluid (CSF) CSPG. See Chondroitin sulfate proteoglycans (CSPG) CSs. See Cardiospheres (CSs) CST. See Corticospinal tract (CST) CT. See Computed tomography (CT) mCT. See Microcomputed tomography (mCT) CTGF. See Connective tissue growth factor (CTGF) CTGTAC. See Cellular, Tissue, and Gene Therapies Advisory Committee (CTGTAC) CTLA-4. See Cytotoxic T lymphocyteeassociated antigen-4 (CTLA-4) Cu-MBG. See Copper-containing mesoporous bioactive glass (Cu-MBG) Cultured epithelial autograft (CEA), 1289 Cultured epithelial cell sheets, 472 Cultured human AFMSCs, 136 Cultured MSCs, 206 Cumuluseoophorus complexes (COC), 1243 Cures Act. See 21st Century Cure’s Act Current Good Tissue Practice (CGTP), 1350e1351 Custom-built bioreactor systems, 793 Cutaneous wounds, celleECM interactions during healing of, 25e28 adhesion and migration, 25e26

apoptosis, 28 differentiation, 27e28 proliferation, 27 Cutting-edge body-on-a-chip, 779e783 ECHO platform, 779e780 organ-on-a-chip systems for personalized precision medicine, 782e783 CVD. See Cardiovascular disease (CVD) CVM. See Center for Veterinary Medicine (CVM) CVVH. See Continuous venovenous hemofiltration (CVVH) CW Bill Young Cell Transplantation Program, 150 CWQPPRARI. See Cyclic H-Trp-Gln-ProPro-Arg-Ala-Arg-Ile (CWQPPRARI) Cx. See Connexins (Cx) Cx43 antisense oligonucleotides, 81 CXCL12. See Stromal-derived growth factor-1 (SDF-1a) Cyclic adenosine monophosphate (cAMP), 426, 1201 Cyclic adenosine monophosphate response element binding protein (CBP), 53 Cyclic Gly-Arg-Gly-Asp-Ser (CGRGDS), 443e444 Cyclic H-Trp-Gln-Pro-Pro-Arg-Ala-ArgIle (CWQPPRARI), 443e444 Cyclic mechanical strains, 432 Cyclin-dependent kinase (CDK), 22 Cyclo-oxygenase-2 (COX-2+), 143e144 Cyclosporin A (CsA), 999e1000, 1201, 1204e1205 CYP3A4, 135 Cys2-His2 zinc finger motif, 747 Cystic fibrosis (CF), 174e175, 1063 Cystic fibrosis transmembrane receptor (CFTR), 174e175, 1063 Cytocompatibility, 697, 813e814 Cytokine, 209, 383, 686, 715, 788 delivery for nerve regeneration, 1227e1228 and growth factors CTGF, 75e76 FGFs, 76 ILs, 77 PDGF, 76 TGF-b superfamily, 75 VEGF, 76 Wnts, 76e77 Cytoplasm, 392 Cytoplasmic proteins, 248e250 Cytosine guanine (CpG), 55e56, 729 methylation, 55e56 Cytosine ring, 170 Cytosine-guanosine oligodeoxynucleotides (CpG), 722 Cytoskeleton, 394e396, 398 formation, 399e400 remodeling, 398 Cytotoxic edema, 375e376

lymphocytes, 727e729 myeloablation, 1010 reaction, 686 Cytotoxic T lymphocyteeassociated antigen-4 (CTLA-4), 716e717

D

DA neurons. See Dopaminergic neurons (DA neurons) Dacron material, 1093 DAG. See Diacylglycerol (DAG) DARPA. See Defense Advanced Research Projects Agency (DARPA) Days postamputation (dpa), 38 dCas9 fused with hybrid VP64-p65-Rta tripartite activator (dCas9-VPR), 754 dCas9. See Nuclease-deficient Cas9 (dCas9) DCCT. See Diabetes Control and Complications Trial (DCCT) DCD. See Donation after cardiocirculatory death (DCD) DCM. See Dilated cardiomyopathy (DCM) DCPA. See Dicalcium phosphate anhydrous (DCPA) DCPD. See Dicalcium phosphate dehydrate (DCPD) DCs. See Dendritic cells (DCs) DDRs. See Discoidin domain receptors (DDRs) Decellularization, 619e620, 790, 1030 application of decellularization techniques, 1241 decellularization/recellularization techniques, 1171 protocols, 619 of rat livers, 1102e1103 Decellularized bioscaffolds, 1046e1047 Decellularized bovine trabecular bone, 699e700 Decellularized ECM (dECM), 816, 1104e1106 as implants, 1122e1123 Decellularized esophageal tissue, 1132 Decellularized porcine ventricular myocardium, 1091 Decellularized tissue, 453e454 scaffold-based approaches, 524 Decellularized vessel grafts, 1030e1031 dECM. See Decellularized ECM (dECM) Dedifferentiated fat cells (DFATs), 703e704 Dedifferentiation, 39e40, 283 Defense Advanced Research Projects Agency (DARPA), 780 Definitive erythropoiesis, 192, 924 Deformation gradient (F), 420 DEG. See Degenerin (DEG) Degenerative process, 355 Degenerin (DEG), 393 Degradability, 514 Degradable polyesters, 572e573 Degradation, 393, 636e637, 666, 1042 profile, 507e510

1390 Degradation (Continued ) degradation mechanisms, 507e508 enzymatic, 509 factors affecting degradation rates, 508e509 surface modification for degradation control, 509e510 surface to volume ratio, 509 Degrees of freedom (DOF), 1182e1183 Delayed engraftment, 197 Delivery cargo, 748e749 DNA, 748 proteins, 748e749 RNA, 748 Delivery methods, 749e750 Delivery route for cells and patches, 1094 Demineralized bone matrixeinduced bone morphogenesis, 406 Dendrimers, 495e496, 749 Dendritic cells (DCs), 205, 344, 680, 715e716, 722e727 DC-based cancer vaccinations, 727 Density of cell injections, 976 gradient separation, 209 Dental applications of CPCs, 607 Dental caries, 907 Dental celletissue recombination approaches, 912, 912f Dental epithelialemesenchymal cell interactions, 908 Dental epithelium, 908 Dental follicle precursor cells (DFPCs), 909 Dental implants, 772e773 Dental mesenchym, 908 Dental mesenchymal cells, 908e909 Dental papilla, 909e910 Dental pulp, 913e914 Dental pulp stem cells (DPSCs), 909, 1264 Dental stem cells (DSCs), 574, 907, 909e910 Dental tissue engineering, 574, 910e917, 910f. See also Cartilage tissue engineering alveolar bone regeneration, 916e917 dental pulp and dentin regeneration, 913e914 DSCs, 909e910 periodontal regeneration, 914e915 tooth development, 908e909 whole tooth engineering, 912e913 Dentin, 908e909 regeneration, 913e914 Dentin matrix protein 1 (DMP1), 913e914 Deoxyribonucleic acid (DNA), 38, 509, 744, 748, 1180 DNA-recognition TALEs, 746 Department of Health and Human Services (DHHS), 1311e1312, 1347 Dermal fibroblasts, 107 Dermal papilla cells (DPs), 1297 Dermal papilla-like tissues (DPLTs), 1299 Dermal sheath (DS), 1304e1305 Dermal sheath cup cells (DSC cells), 1305 Dermalastyl, 617

INDEX

Dermatan sulfate, 618 Dermis, 66, 618e619, 1284, 1289 Detergents, 790, 1064 Detrusor muscle, 1263 Device (FDA definition), 1348 Dexamethasone, 762, 940, 1210e1211 DFATs. See Dedifferentiated fat cells (DFATs) DFPCs. See Dental follicle precursor cells (DFPCs) DHEW. See United States Department of Health, Education and Welfare (DHEW) DHHS. See Department of Health and Human Services (DHHS) DHT. See Dihydrotestosterone (DHT) Diabetes, 223, 344e345, 987 cellular heterogeneity in, 106 Diabetes Control and Complications Trial (DCCT), 987 Diabetes mellitus, 335 Diacrylated EH (EHD), 578 Diacylglycerol (DAG), 38e39 Dialysate, 1150 Dialysis, 1165 Dialyzers, 1150 Diamine chain extender, 567 Dicalcium phosphate anhydrous (DCPA), 592e593, 706 Dicalcium phosphate dehydrate (DCPD), 592e594 Dicarboxylic acid monomers, 574 DICE. See Dual integrase cassette exchange (DICE) DICER enzymes, 57e58 Dichloromethane, 823 Dickey amendment, 126b DickeyeWicker Amendment, 1311e1312, 1333e1334 Dickkopf-related protein 1 (DKK-1), 301e302 Dielectrophoresis array, 95e96 Differentiated germ cells, 1251e1252 Differentiation, 445e446, 788, 894, 1084e1086 celleECM interactions during healing of cutaneous wounds, 27e28 during regenerative fetal wound healing, 30 differentiation-associated genes, 49 electrical stimulation, 1085e1086 mechanical stimulation, 1084e1085 potential of rMAPCs and hMAPCs in vitro, 183 signal transduction events during celleECM interactions, 23e24 Diffuse axonal injury, 370e371 Diffusion, 787 Diffusion tensor imaging (DTI), 382 Diffusion tensor magnetic resonance imaging (DTMRI), 1081 Digestive system, 1264 Digital Imaging and Communications in Medicine format, 831e832

Dihydrotestosterone (DHT), 1298 Dilated cardiomyopathy (DCM), 473e474 DILI. See Drug-induced liver injury (DILI) Dimensionality effect, 451e460 cellular responses in modifying ECM, 459e460 to three-dimensional substrates, 455e457 effect of externally applied mechanical stimuli, 457e459 importance of three dimensions, 452 substrates for 3D culture, 452 3D culture and materials development, 454e455 3D scaffolds polymers for, 452 preparation of, 452e454 Dimethyl sulfoxide (DMSO), 149 Dipeptidyl peptidase IV antigen (DPPIV antigen), 232 Direct conversion of somatic cells, 928 Direct lineage conversion. See Transdifferentiation process Direct neuronal, 370e371 Direct reimbursable expenses, 1324 Direct remuscularization, 247, 254e256 Direct reprogramming, 172 from fibroblasts, 174 Direct transdifferentiation of cells, 172 Directed differentiation, 336 “Directed donor” programs, 150e151 Discoidin domain receptors (DDRs), 18 Disease intrinsic defects in, 283e284 modeling, 172e174, 769e770 additional, 778e779 challenges and future possibilities in, 174 “Disease-in-a-dish” modeling, 351 for retinal disorders, 362e363 induced pluripotent stem cellebased phenotyping, 363 three-dimensional retinal organoids, 363 Disk-shaped nanoparticles, 720 Displacement degenerate oligonucleotideprimed-PCR (DOP-PCR), 97 Displacement vector, 419 Distraction enterogenesis, 1135 Diversity of ECM, composition and, 15e16 Divisive analysis, 101e102 DKK-1. See Dickkopf-related protein 1 (DKK-1) DMD. See Duchenne muscular dystrophy (DMD) DMP1. See Dentin matrix protein 1 (DMP1) DMSO. See Dimethyl sulfoxide (DMSO) DNA. See Deoxyribonucleic acid (DNA) DNA demethylation, 57 DNA methylation, 100 modifications, 170 DNA methyltransferase enzymes, 56 DNMT enzymes, 56e57 DNMT1, 56

INDEX

DNMT3A, 55e56 DNMT3B, 55e56, 114 DNMT3L, 56e57 DOF. See Degrees of freedom (DOF) Domestic efforts, 1374 Donation after cardiocirculatory death (DCD), 993 Donor donor-derived mononuclear cells, 975 donor-specific HLA class I alleles, 234 donor-specific sHLA-I, 234 donor-specific tolerance, 999e1000 induction, 198e199 hepatocytes, 232 liver tissue, 237e238 microglia cells, 158 and procurement issues, 1337e1338 DOP-PCR. See Displacement degenerate oligonucleotide-primed-PCR (DOP-PCR) Dopamine dysregulation syndrome, 223 Dopaminergic neurons (DA neurons), 176 Dorsal pancreatic bud emerges, 341 Dorsal root ganglia (DRG), 1226e1227 neurons, 44 Dorsoventral (DV), 44 Double anastomosis model. See Cross anastomosis model Double CBT, 156 Double-blind study, 159 Double-strand breaks, knockouts via, 742e743 interaction of Cas9/gRNA, 743f Doxorubicin (DOX), 485, 775e776 DPCs, 1302 DPLTs. See Dermal papilla-like tissues (DPLTs) DPPIV antigen. See Dipeptidyl peptidase IV antigen (DPPIV antigen) DPs. See Dermal papilla cells (DPs) DPSCs. See Dental pulp stem cells (DPSCs) DRG. See Dorsal root ganglia (DRG) Droplet microfluidics, 96 DROSHA enzymes, 57e58 Drosophila, 1, 5 Drug delivery systems, 763e764 technologies, 574 Drug-induced liver injury (DILI), 1108 Drug(s), 81, 1348 development, 1108 organoids, 1109e1110 diffusion kinetics, 770 discovery, 174, 839 drug-eluting stents, 656 drug-loaded PLGA nanoparticles, 1207 resistance, 108 screening, 769e770, 841 targeting, 721 testing, 777e778 Drusen, 354 Dry AMD, 1208e1210 DS. See Dermal sheath (DS) DSC cells. See Dermal sheath cup cells (DSC cells)

DSCs. See Dental stem cells (DSCs) DTI. See Diffusion tensor imaging (DTI) DTMRI. See Diffusion tensor magnetic resonance imaging (DTMRI) Dual integrase cassette exchange (DICE), 748 Dual setting system, 604e605 “Dual-platform” method, 154 Dual-specific phosphatases (DUSP2/7), 54 Dual-tissue guteliver system, 775e776 DubineJohnson syndrome, 235e236 Duchenne muscular dystrophy (DMD), 273e274, 752e753, 963, 971 gene, 752 mdx mouse model, 284 Dulbecco’s Modified Eagle Medium, 212 DUSP2/7. See Dual-specific phosphatases (DUSP2/7) DV. See Dorsoventral (DV) Dye exclusion assays, 195 Dynamic compression, 424 “Dynamic reciprocity” model, 15 Dynamic tissue shear, 424 Dysfunctional wound healing, 65 Dystroglycan, 617 Dystrophin, 282e283 gene, 752 mutations, 285 protein, 273e274, 752

E

E-cadherin, 2 promoter, 4 E-selectins, 679 E2e2A and 2B class I basic helix-loophelix factor, 4 E3-ligase, 2 E8 (Thermo Fisher Scientific), 118 Ear perichondrium, 897 EB. See Electron beam (EB) EBL. See Electron beam lithography (EBL) EBP. See Elastin-binding protein (EBP) EBs. See Embryoid bodies (EBs) Eccrine glands. See Sweat glands ECFC. See Endothelial colony-forming cells (ECFC) ECHO platform. See Ex vivo console of human organoids platform (ECHO platform) ECMs. See Extracellular matrices (ECMs) Ecotropic viral integration factor 5 (Evi5), 41 ECs. See Endothelial cells (ECs) ECS. See Extracapillary space (ECS) Ectodermal lineage-derived CNC, 247 Ectomesenchyme, 908 Ectosomes, 208 ED. See Erectile dysfunction (ED) ED6. See Embryonic day 6 (ED6) EDA. See Extra domain A (EDA) EDH. See Epidural hematomas (EDH) Edmonton protocol, 989e990 EDTA. See Ethylenediaminetetraacetic acid (EDTA) EE. See Enriched environments (EE)

1391 EF1. See dZeb homeobox 1 (Zeb1) Efalizumab, 1000 Effector immune cells, 721e722 Effluvium, 1303e1304 EGF. See Endothelial growth factor (EGF); Epidermal growth factor (EGF) EGFP. See Enhanced green fluorescent protein (EGFP) Egg donors, compensating, 1323e1324 EHD. See Diacrylated EH (EHD) Eicosanoids, 888 Ejaculation, 1255e1256 Ejaculatory system, 1255e1256 engineering vas deferens, 1255 SGE, 1255e1256 urethra reconstruction, 1255 Elastatropin, 617 Elastic cartilage, 937 fibers, 1263 Elastin, 617 derivatives, 643e644 Elastin receptor complex (ERC), 18 a-Elastin, 644 k-Elastin, 644 Elastin-binding protein (EBP), 18 Elastin-like polypeptide (ELP), 644 Elastography, 799 Electrical conduction, 1086e1087 Electrical fields, 531 Electrical stimulation, 1085e1086, 1229 Electrically conductive substrate, 450e451 Electrochemical equilibrium, 371 Electroconductive scaffolds for nerve regeneration, 1229 Electroejaculation, 1255e1256 Electron beam (EB), 470e471 Electron beam lithography (EBL), 444, 447t Electronic patch, 1077 Electronics, 769 Electroporation, 497e498, 752 delivery, 749 Electrospinning, 447, 477e478, 541, 544, 546e547, 552, 633e635, 1077, 1271 Electrospun nanofiber scaffolds, 1034e1035 Electrospun recombinant human tropoelastin, 617 Electrospun scaffolds, 857, 1077 electrospun/nanofibrous scaffold, 453 Elixir sulfanilamide, 1346 ELP. See Elastin-like polypeptide (ELP) “Embrace” device, 80 Embryo research oversight committee (EMRO committee), 1337 Embryoid bodies (EBs), 118, 140, 762, 925 formation, 212 Embryonic chondrogenesis, 957 Embryonic day 6 (ED6), 128 Embryonic death, irreversibility as criterion for diagnosing, 127e128 Embryonic environment simulation, 1300e1304 expression of hair follicle-related stem cell markers, 1301f

1392 Embryonic environment simulation (Continued ) injection of HIMSC-CM, 1301f one-time injection of HSC, 1302f Embryonic mesoderm-derived mesenchymal cells, 409 Embryonic morphogenesis, 1 Embryonic myosin isoforms, 276 Embryonic progenitor cells, 939 Embryonic Stem Cell Research Oversight committee (ESCRO committee), 1317 Embryonic stem cells (ESCs), 49, 81e82, 113, 138, 169, 181, 209e210, 239e240 , 247, 255, 335e341, 337f, 762e763, 855, 909, 954, 996, 1035, 1063, 1168, 1202, 1245, 1264. See also Human embryonic stem cells (hESC) derivation, 113e114 ES-derived, cells, 240 ESC-derived cells, 255 ESC-derived exosomes, 209e210 ESC-expressed miRNAs, 58 ESCsederived cardiomyocytes transplantation, 175e176 iPSCs, 762 MSCs, 762e763 naive embryonic stem cells, 119 Embryonic wounds, 76 Embryonic-specific markers, 134e135 Embryos, 762, 1310e1311 EMBs technique. See Explanted microcirculatory beds technique (EMBs technique) EMRO committee. See Embryo research oversight committee (EMRO committee) eMSCs. See Endometrial mesenchymal stem cells (eMSCs) EMT. See Epithelial-mesenchymal transition (EMT) ENaC. See Epithelial sodium channel (ENaC) Encapsulation processes, 911 End-stage renal disease (ESRD), 1029, 1149, 1165 complete bioartificial kidney system for use in, 1159 WeBAK in preclinical, 1158 Endangered species, conservation of, 176 Endocardium, 1094 Endochondral bone formation, 696 Endochondral ossification process, 853 Endogenous repair stimulation, 254 Sox2 expression, 169e170 Endogenous stem cells, 1200e1201 factors for stimulation, 1203e1205 Endoglin, 220e221 Endometrial mesenchymal stem cells (eMSCs), 1246 Endometrium, 1240 cells, 1240 Endonucleases, 742

INDEX

Endoscopic submucosal dissection (ESD), 473 esophagus reconstruction after ESD treatment, 473 Endothelial barrier, 232 Endothelial cells (ECs), 307, 431, 475, 488, 688, 836e837, 1029e1030, 1074, 1087, 1200, 1257, 1270 differentiation, 24, 27e28 EC-induced cardiomyocyte protection, 1091 EC-specific genes, 1267 layer, 776 tight junctions, 1199 Endothelial colony-forming cells (ECFC), 311e313, 837 Endothelial growth factor (EGF), 16, 1298e1299 EGF-like repeats, 16 Endothelial NOS (eNOS), 427 activity, 311 Endothelial progenitor cells (EPC), 257, 307e308, 311e315, 317e320, 701, 841e842 angiogenesis and vasculogenesis, 317 identification and isolation, 311e312 in vitro expansion, 312e313 role in physiological and pathological neovascularization, 313e315 tissue engineering, 319e320 tissue regeneration, 317e319 Endothelialization, 1087 Endothelium, 232 Endotoxins, 515 Energy absorbed at failure, 1181 Energy-dispersive X-ray spectroscopy, 702 Engelbreth-Holm-Swarm mouse sarcoma cells, 118, 1104 Engineered/engineering. See also Tissue engineering (TE) of cell-based renal constructs, 1169e1172 in situ kidney regeneration, 1170t complex tissue constructs, 478e481 of functional vaginal tissue, 1239e1240 kidney-like constructs, 1171 nanostructured scaffolds, 487f neo-tissue, 300e301 scaffolds, 535 strain, 418 Engraftment of human ESC-derived beating cardiomyocyte, 255 of human ESC-derived cardiac cells, 255 methods to improving, 238e239 posttransplant, 998e999 Engrailed-1 gene, 75 Enhanced green fluorescent protein (EGFP), 1135, 1167 Enhanced permeation and retention effect, 718e719 Enhanced propagation method (EP method), 1156 “Enhanced specificity” SpCas9 (eSpCas9), 745

eNOS. See Endothelial NOS (eNOS) Enriched environments (EE), 1206 Enteric neuropathies, 1140 Enthesis, 960e961 Envisioned regenerative medicine manufacturing systems, 1370e1373, 1371t, 1372f Enzymatic degradation, 509 Enzymatic facilitation, 507e508 Enzymatic neutralization, 295e296 Enzyme replacement therapy (ERT), 158, 198 Enzyme(s), 790 defect, 157 enzyme-degradable PU, 573 enzyme-linked immunosorbent assay, 1302 EO. See Executive Order (EO) EP method. See Enhanced propagation method (EP method) EPC. See Endothelial progenitor cells (EPC) Epiblast, 247e248 Epiblast stem cells (EpiSCs), 50, 119 Epicardial progenitor cells, 251e252 Epicardial surface of heart, 1094 Epicardium, 251e252, 1094 Epidermal appendages, 66 Epidermal cells, 1283 Epidermal growth factor (EGF), 6, 182e183, 238e239, 667, 1116, 1134, 1185e1186, 1201 Epidermal repair, 1287e1288 Epidermal stem cells, 83e84 Epidermis, 41, 66, 1283e1284 Epidermolysis bullosa dystrophica, 754 Epidural hematomas (EDH), 370 Epigenetic changes, 96e97 memory, 170 remodeling, 170 Epimerization, 617e618 Epinephrine, 780 EpiSCs. See Epiblast stem cells (EpiSCs) Episomal vectors, 170 Epithelial cells, 1e2, 398 Epithelial injury, 1062e1063 Epithelial mesenchymal signaling, 1273e1274 Epithelial polarity, 3 Epithelial sodium channel (ENaC), 393 Epithelial-mesenchymal cell signaling, 1133 Epithelial-mesenchymal transition (EMT), 1 induction, 8, 8f molecular control, 5e8 additional signaling pathways, 7e8 ligand-receptor signaling, 6e7 transcriptional program, 4e5 posttranscriptional regulation, 5 regulation at promoter level, 5 transcription factors, 4 Epithelialemesenchymal interaction, 408 Epithelium, 38, 662e663

1393

INDEX

EPO. See Erythropoietin (EPO) Epoxides, 563e564 ePTFE. See Expanded polytetrafluoroethylene (ePTFE) ErbB2/HER-2/Neu receptor, 7 ERC. See Elastin receptor complex (ERC) Erectile dysfunction (ED), 1251 stem cell therapy for, 1258 ERK. See Extracellular signal-regulated kinase (ERK) ERK1/2. See Extracellular regulated kinase 1/2 (ERK1/2) ERT. See Enzyme replacement therapy (ERT) Erythrocytes, 924e925 Erythroid cells, 192 Erythroid progenitors, 274 Erythropoiesis, 924 Erythropoietin (EPO), 311, 924e925, 1165, 1168, 1201 Escherichia coli (E. coli), 741, 841, 1271 ESCRO committee. See Embryonic Stem Cell Research Oversight committee (ESCRO committee) ESCs. See Embryonic stem cells (ESCs) ESD. See Endoscopic submucosal dissection (ESD) Esophageal Doppler monitoring, 1159 Esophageal reconstruction, 1132 Esophagus, 1131e1134 reconstruction after ESD treatment, 473 eSpCas9. See “Enhanced specificity” SpCas9 (eSpCas9) ESRD. See End-stage renal disease (ESRD) Esrrb gene, 57 Establishment Registration rule, 1351 ET. See Excitation threshold (ET) Ethics Committee of the American Society for Reproductive Medicine, 1335 Ethylene oxide (EtO), 514 Ethylenediaminetetraacetic acid (EDTA), 815e816, 940, 1030 Etiologies of nonobstructive azoospermia, 1252 EtO. See Ethylene oxide (EtO) Ets-1 transcription factor, 3 EUROCORD, 149e150 Evaluation of Devices Used with Regenerative Medicine Advanced Therapies, 1369e1370 Evi5. See Ecotropic viral integration factor 5 (Evi5) EVLP. See Ex vivo lung perfusion (EVLP) EVs. See Extracellular vesicles (EVs) EVTs. See Extravillous cytotrophoblasts (EVTs) Ex vivo console of human organoids platform (ECHO platform), 779e780 Ex vivo environment, 1067 Ex vivo lung perfusion (EVLP), 790, 1066e1067 Excitation threshold (ET), 1076e1077 Executive Order (EO), 126b ExoCarta database, 209

Exosomal/exosomes, 205e206, 209e211, 254e255 exosome-associated proteins, 209 proteins, 209 signaling, 211 Expanded polytetrafluoroethylene (ePTFE), 1029 Expedited Programs for Regenerative Medicine Therapies for Serious Conditions, 1369e1370 Experimental autoimmune encephalitis model, 223 Explanted microcirculatory beds technique (EMBs technique), 301 Extra domain A (EDA), 23e24 Extracapillary space (ECS), 1150 Extracellular domains of syndecans, 16e18 Extracellular matrices (ECMs), 1, 15, 22, 39, 67, 135, 210e211, 273e274, 299, 308, 391e393, 395f, 397f, 405e406, 409, 423, 437, 470, 486, 489e490, 538, 560, 613e619, 628e630, 661e662, 667, 696e697, 701e702, 762, 770, 787, 808, 831e832, 853, 910e911, 937, 954, 1030, 1042e1043, 1064, 1102e1103, 1115, 1165, 1180, 1200, 1251, 1268, 1281, 1284 by cell culture and synthetic polymers, 1033e1034, 1034f cellular responses in modifying, 459e460 collagen, 615e616 component, 279e280, 621, 1226 composition and diversity, 15e16 deposition pattern, 634e635 ECMecell dynamic reciprocity, 438 ECMsebased products, 614te615t elastin, 617 fibronectin, 616e617 GAGs, 617e618 hydrogels, 620e621 integrin family, 17f laminin, 617 MBVs, 618e619 molecules, 16e18, 27e28, 31, 1224, 1231 for nerve regeneration, 1226e1227 patellar tendon healing with, 1188e1189 peptide analogs, 1123e1124 scaffolds, 621, 688, 1286 regulatory considerations, 622 substitutes and scaffolds, 80 Extracellular proteins, 618 Extracellular regulated kinase 1/2 (ERK1/ 2), 398, 489 Extracellular signal-regulated kinase (ERK), 22, 49e50 Extracellular vesicles (EVs), 205, 208, 958 Extracorporeal renal replacement, 1149 advancements in conventional renal replacement therapy devices, 1151e1152 BRECS treating acute kidney injury, 1157e1158 challenge of cell-based device, 1156

clinical experience with renal assist device, 1154 complete bioartificial kidney system, 1159 cost-effective storage and distribution for cell devices, BRECS, 1156e1157 devices used in conventional renal replacement therapy, 1150e1151 future advancements for wearable, 1159e1160 immunomodulatory effect of renal assist device, 1154e1155 RAD, 1152e1153, 1153f renal assist device therapy of acute kidney injury causing, 1153e1154 requirements of renal replacement device, 1149e1150 SCD, 1155e1156 WeBAK in preclinical end-stage renal disease model, 1158 Extracorporeal renal replacement. See also Cell-replacement therapy Extraembryonic tissues, 181 Extrahepatic biliary tree, 341 Extraovarian sources, 1245 Extravillous cytotrophoblasts (EVTs), 133e134 Extrusion-based printing, 806e807, 815, 834 Extrusion-based systems, 816e817

F

FA. See Focal adhesion (FA); Fractional anisotropy (FA) Fabricating porous scaffolds, methods for, 511 Fabricating tissue engineered vascular grafts, 1030e1038 biodegradable synthetic-based scaffolds, 1038 biological-based scaffolds, 1030e1033 hybrid scaffolds, 1033e1037 Fabrication techniques, 446e447 of thermoresponsive cell culture substrate, 471 performance of cell sheet harvesting, 472f two-dimensional substrate patterning techniques, 447t FACS. See Fluorescence-activated cell sorting (FACS) FACT. See Foundation for Accreditation of Cellular Therapy (FACT) Factor IXa (FIX), 1017 FAH. See Fumarylacetoacetate hydrolase (FAH) FAHe/emouse model, 239e240 FAK. See Focal adhesion kinase (FAK) Familial hypercholesterolemia, 1101e1102 Family banks, 150e151 Fanconi anemia, 197 FAPs. See Fibrogenic/adipogenic progenitors (FAPs) Fast Track designation, 1369e1370

1394 Fat fat-derived hMSCs, 222 stem cells from cellular characterization, 296e297 cellular fractions, 295e296 clinical delivery of adipose-derived cells, 297e300 engineered neo-tissue, 300e301 therapeutic safety of adipose-derived cells, 301e302 Fat grafting. See Autologous lipotransfer FATC. See FemureACLetibia complex (FATC) Fate mapping of c-Kit expression in heat, 248e250 fate-mapping studies, 248 FBGC. See Foreign body giant cell (FBGC) FBS. See Fetal bovine serum (FBS) FBs. See Fibroblasts (FBs) FCS. See Fetal calf serum (FCS) FD&C Act. See Food, Drug, and Cosmetic Act (FD&C Act) FDA. See US Food and Drug Administration (FDA) FDA Modernization Act (FDAMA), 1350, 1357 Fecal incontinence, 1141e1142 FED. See Fuchs endothelial dystrophy (FED) Federal Food and Drugs Act, 1346, 1357 Federal funds, 1312 Federal policy, 1313 Feeder-free culture system, 118e119 Female reproductive system. See also Male reproductive system ovaries, 1242e1245 TE applications, 1245e1246 principles, 1237e1238 uterus, 1240e1242 vagina, 1238e1240 Female SpragueeDawley rats, 1271e1272 FemureACLetibia complex (FATC), 1185 Femuregraftetibia complex, 1185 FemureMCLetibia complex (FMTC), 1179 Fertility treatment, 1318e1320 Fertilization, 399 Fetal, 1168e1169 cells, 1309e1310 deformities, 391 development and regenerative medicine, 1009e1011 ECM, 73 fibroblasts, 29e30 germline, 1019 hyaluronan, 30 liver hematopoiesis, 192e193 rat ventricular cardiomyocytes, 1076 scarless wound repair, 72e73 wound histologic sections, 74f sheep model, 1012e1013 skin development, 72 fetal scarless wound repair, 72e73 stem cells, 1266

INDEX

Fetal bovine serum (FBS), 114, 208, 325e326 Fetal calf serum (FCS), 182e183 Fetal wound healing, 80 celleECM interactions during regenerative, 28e30 adhesion and migration, 29 apoptosis, 30 differentiation, 30 proliferation, 29 FGF. See Fibroblast growth factor (FGF) FGF recombinant 1 (FGFR1), 909e910 FHF. See Primary heart field (FHF) Fiber bridging, 606 fiber-reinforced calcium phosphate cements, 605e606 reinforcement, 605e606 mechanics of fiber-reinforced calcium phosphate cements, 605e606 semipermeable membrane, 1150 Fibrillar components, 1180 Fibrin, 815, 837, 1047e1048, 1090 cable, 1225e1226 clot, 38 derivatives, 644e645 fibrin-based bioink, 837 fibrinefibronectin provisional matrix, 25 fibrineheparineNGF matrix, 1228 glues, 299, 644 sealants, 644 Fibrinogen, 644, 681, 811, 837 Fibrinolysis, 135 Fibrinopeptides A and B, 644 Fibroblast growth factor (FGF), 7, 21, 52, 76, 208, 280, 339e340, 877, 909e910, 942, 957, 1059e1060, 1087e1088 Fgf1, 43e44 FGF2, 43e44, 119, 298 FGF-5, 76 FGF-7, 76, 1300 FGF-7/KGF, 1303 FGF-9, 76 FGF-10, 76 Fibroblasts (FBs), 67, 70, 107, 171, 286, 615e616, 680, 838, 1074, 1187, 1284 Fibrocartilage, 937 Fibrocytes, 107 Fibrodysplasia ossificans progressiva, 212 Fibrogenic/adipogenic progenitors (FAPs), 273 Fibronectin (FN), 29, 279, 393, 409, 437, 523e524, 567, 616e617, 811, 911, 1270 FN-gamma, 186e187 gene, 4 patterns, 1081 subunit, 529 Fibroproliferative scarring, 68e72 hypertrophic scars, 70e72 keloids, 70 Fibroproliferative scars, 77e78 Fibrosa, 1044 Fibrosis, 68e70, 322, 391, 400, 682e684 cellular heterogeneity in, 107

Fibrotic bladder model, 1271e1273 Fibrous encapsulation, 683e684 Fibrous polyglycolic acid scaffolds, 1076e1077 Fibrous proteins, 630 Fibrous tissues, 615e616 Ficoll Paqueebased isolation, 208 FicollePaque density-gradient media, 208 FIH Studies. See First in Human Studies (FIH Studies) Filamentous (F), 394e395 Financial costs, 229e230 First in Human Studies (FIH Studies), 1349e1350 First intention wound healing, 684 First-generation biomaterials, 559e560 FISH. See Fluorescence in situ hybridization (FISH) Fistula, 1142 FIX. See Factor IXa (FIX) Flavin mononucleotide photosensitizer, 815 Flp enzymes, 747e748 Fluidic systems, 773e774 Fluorescence in situ hybridization (FISH), 117 Fluorescence microscopy, 454 Fluorescence-activated cell sorting (FACS), 95, 96f, 99 Fluorescent bar coding, 99 beads, 799 dye exclusion, 195 Fluorescently labeled albumin, 773e774 5-Fluorouracil (5-FU), 77e78, 775, 777, 780 Fms-related tyrosine kinase 3 ligand, 727e729 FMTC. See FemureMCLetibia complex (FMTC) FN. See Fibronectin (FN) FNIII7e10 fragment, 529 Focal adhesion (FA), 394e395, 437e438, 458 complexes, 635 mechanosensing, 458 Focal adhesion kinase (FAK), 19e20, 398, 437e438 FAK-induced integrin, 20 mechanosensation in FAK Y397 phosphorylation, 19e20 phosphorylation, 449 protein, 489 Focal cartilage repair, 937e938 Focal contacts, 437e438 FokI, 746 Follicular structures, 1304 Folliculogenesis, 1242 Follistatin, 1303 Follow-up Phase IIb study, 1154 Food, Drug, and Cosmetic Act (FD&C Act), 1346 for three-dimensional culture, 452 Foreign body giant cell (FBGC), 680 cell formation and interactions, 682e683

INDEX

Foreign body response, 516e517 Foreign nucleic acid gene transfer, 1186 Foreign protein, 689e690 Fo¨rster resonance energy transferebased sensors, 20 Foundation for Accreditation of Cellular Therapy (FACT), 149e150 Fovea, 352 FOXA2 gene, 339e340 Forkhead box transcription factor, 4 Fractional anisotropy (FA), 382e383 Fracture healing, 853 Fragile X mental retardation 1 gene, 754 Fragile X syndrome, 172 Frameshift mutations, 752 Francisella, 745 Free tissue transfer techniques, 890e891 Free-radical scavenging, 1150 Freeze gelation, 539e540 Freeze-drying process, 527 techniques, 538 Freezeethaw process, 453e454 Frictional sliding, 606 Friedreich ataxia, 172 Frizzled family, 6 Frozen-hydrated surface studies, 527 Frustrated phagocytosis, 679 FTE. See Functional tissue engineering (FTE) FTSG. See Full-thickness skin grafts (FTSG) 5-FU. See 5-Fluorouracil (5-FU) Fuchs endothelial dystrophy (FED), 449 Full-length ECM macromolecules, 1123 Full-thickness skin grafts (FTSG), 1285 Fully cell-based, self-assembled corneal constructs, 1119e1122 Fully integrated collaborative manufacturing systems, 1370e1372 Fully integrated modular, and automated manufacturing systems, 1370 Fumarate-based polymers, 576e577 Fumaric acid, 576e577 Fumarylacetoacetate, 750e751 Fumarylacetoacetate hydrolase (FAH), 750e751 Functional assays, 689 Functional groups, 531 Functional mucosal barrier, 1138 Functional restoration, evaluate, 432 Functional testing, 1181 Functional tissue engineering (FTE), 1179 application, 1185e1187 cell therapy, 1187 gene therapy, 1186e1187 growth factors, 1185e1186 healing of ligaments and tendons, 1183e1185, 1188e1192 normal ligaments and tendons, 1180e1183 Functional vaginal tissue engineering, 1239e1240 Fuzzy c-means clustering, 103

G

G proteinecoupled receptors (GPCR), 393, 425 G-CSF. See Granulocyte-colony stimulating factor (G-CSF) GA. See Geographic atrophy (GA); Glycolic acid (GA) GABA. See g-Aminobutyric acid (GABA) GABAergic interneurons, 1208 GAD. See Glutamic acid decarboxylase (GAD) GAGs. See Glycosaminoglycans (GAGs) Galactose-a1,3-galactose (a-gal), 690 b-Galactosidase complementation method (b-gal complementation method), 276 Galactosyl-a-1,3-galactose, 619e620 Galactosylated chitosan (GC), 1103 Galactosylceramidase lysosomal enzyme, 158 Gamma secretase inhibitor XX (GSiXX), 340 Gap junctions, 1242 targeting, 81 GAPDH. See Glyceraldehyde 3-phosphate dehydrogenase (GAPDH) Gas foaming, 552 Gastric acellular matrix, 1133 Gastric disease, 1134 Gastric submucosal space (GSMS), 998 Gastroesophageal reflux disease, 1131 Gaussian curve, 138 GC. See Galactosylated chitosan (GC) GCS. See Glasgow Coma Scale (GCS) GDF11 inhibits, 280 GDF8. See Myostatin GDFs. See Growth/differentiation factors (GDFs) GDNF. See Glial cell-line-derived neurotrophic factor (GDNF) GEF. See Guanine nucleotide exchange factor (GEF) Gel-forming polymers, 601 Gelatin, 601, 643, 815, 821e822, 836e837, 1048 gelatin-based bioink, 1079 Gelatin methacrylate (GelMA), 453e454, 808e809, 836e837, 912e913 hydrogel, 633e634 Gelatinization, 549 Gelation process, 809, 821 Gellan gum, 543e545 in bone tissue engineering applications, 544e545 processing methods, 543e544 GelMA. See Gelatin methacrylate (GelMA) GEMM. See Granulocyte, erythroid, macrophage, megakaryocyte (GEMM) Gene editing, 1336 in regenerative medicine applications, 750e755 delivery cargo, 748e749 delivery methods, 749e750 genome editing tools, 741e748

1395 strategies, 285 Gene therapy, 198, 627, 860e861, 1009e1010, 1066e1067, 1186e1187, 1354e1355, 1368 approach, 234 MSCs for, 322e325 Gene(s), 741 complementation, 972e973 delivery techniques, 752 gun delivery, 749 Generic epithelial cell markers, 1267 Genetic disorders, 907 genetic linage fate-mapping experiments in mice, 248e250 mosaicism, 97 mutation, 1081e1082 Genetic lineage-tracing approach, 248 experiments in mice, 263e264 Genetically engineered elastin-like polypeptides, 574 Genetically modified hBMSCs, 1264 Genome editing, 752, 1019e1020 tools, 741e748 other genome manipulation tools, 747e748 targetable nucleases, 741e747 Genome manipulation tools integrase, 748 recombinase, 747e748 transposons and transposase, 747 Genome-wide association studies (GWAS), 174, 354 Genomic integration, 748 genomic integration-associated insertional mutagenesis, 1018e1019 Genomic stability, 172 Genotype of cell lines, 769 Geographic atrophy (GA), 354 Germ cell lineages, 750 Germ layer cell types, 762 Germline stem cells (GSCs), 1244e1245 Geron Corporation, 1338 Gestation, 133 Gestational surrogacy, 1240 GFAP. See Glial fibrillary acidic protein (GFAP) Gfi1, transcription factors, 872 GFP. See Green fluorescent protein (GFP) GFs. See Growth factors (GFs) Ggf-2. See Glial growth factor 2 (Ggf-2) Gillmore needles method, 597e598 Glasgow Coma Scale (GCS), 370 Gli, hedgehog-activated transcription factor, 7 Glia(l), 370e371, 1206 scar, 1200 Glial cell-line-derived neurotrophic factor (GDNF), 1201, 1228 Glial fibrillary acidic protein (GFAP), 370 Glial growth factor 2 (Ggf-2), 44 Glioblastoma, 108, 186 Glioma, antitumor effects of MAPCs in, 186

1396 Glomerular filtration, 1167 Glow discharge deposition. See Plasma deposition fGlow discharge plasma deposition, 531 Glucagon-like peptide-2 (GLP-2), 1138 a-Glucan polymers, 548 N-Glucosamine, 641 Glucose lability, 989 Glutamic acid decarboxylase (GAD), 999 Glutaraldehyde, 815 Glutathione (GSH), 1149e1150 GSH-metabolizing enzymes, 1150 GlyA. See Antiglycophorin A (GlyA) Glyceraldehyde 3-phosphate dehydrogenase (GAPDH), 209 Glycerol, 571 Glycine, 640 Glycogen synthase kinase 3b (GSK-3b), 5, 52e53 Glycolic acid (GA), 568 Glycolide, 568 Glycophase glass-modified substrates, 443 Glycoproteins, 393 expression, 297 Glycosaminoglycans (GAGs), 15e16, 393, 438, 538, 580, 615, 617e618, 630, 639, 837, 956, 1045e1046, 1169, 1231e1232 Glycosylated hemoglobin (HbA1C), 987 Glycosylphosphatidylinositol (GPI), 18 Glypicans, 618 GM. See Gray matter (GM) GM-CSF. See Granulocyte macrophageecolony-stimulating factor (GM-CSF) GMP. See Good manufacturing practices (GMP) GO. See Graphene oxide (GO) Goblet cells, 1060e1061 Gold nanoparticles (AuNPs), 488, 492, 494 Golden Gate assembly, 746e747 Golden Retriever muscular dystrophy model, 286 Good Guidance Practices, 1357 Good manufacturing practices (GMP), 924e925 Goosecoid homeobox protein, 4 GPCR. See G proteinecoupled receptors (GPCR) GPI. See Glycosylphosphatidylinositol (GPI) Gr-1. See Granulocytes (Gr-1) Gradient scaffolds, anisotropic and, 511e512 Graft survival cotransplanted with cells, 185 Graft versus host disease (GvHD), 149, 184, 196, 211, 220, 309 Graft versus tumor effect (GVT effect), 196 Graft vs. host disease (GvHT) MAPCs effect on, 184 Graft-derived satellite cells formation, 975, 975f Grafted cells, 1268 “Grafting-from” approaches, 656

INDEX

“Grafting-to” approaches, 656 GRAGIL. See Groupe RhineRhoˆneeAlpes et Gene`ve Pour la Greffe d´ıˆlots de Langerhans (GRAGIL) Granular synthetic HA, 706 Granulation tissue, 26, 67, 680, 682 development, 678 Granulocyte, erythroid, macrophage, megakaryocyte (GEMM), 924 Granulocyte macrophageecolonystimulating factor (GM-CSF), 309, 716 Granulocyte-colony stimulating factor (G-CSF), 196, 241, 257, 309e310, 378, 1154, 1203e1204 Granulocytes (Gr-1), 194e195 Granulosa cell spheroids, 1243 Graphene oxide (GO), 451 Graphene surfaces, 451 Gray matter (GM), 381 Green fluorescent protein (GFP), 40, 183, 276e279, 870e871 GFP-labeled MSCs, 1241e1242 GFP-labeled transgenic mMAPCs, 183 Green strain (E), 420 gRNA. See Guide RNA (gRNA) Groupe RhineRhoˆneeAlpes et Gene`ve Pour la Greffe d´ıˆlots de Langerhans (GRAGIL), 988e989 Growth factors (GFs), 40, 339e340, 619, 680, 763e764, 788, 892, 907, 964, 1051e1052, 1116 additional studies on secretion, 1300 application of FTE, 1185e1186 in vitro studies, 1185e1186 in vivo studies, 1186 and cell signaling molecules, 80e81 growth factor-b pathway, 6 growth factorecoated PCL scaffolds, 812 growth factoreembedded scaffold materials, 1135 matrix binding with, 1271 therapy, 1298 Growth-arrested embryo, 115 Growth/differentiation factors (GDFs), 407e408 GSCs. See Germline stem cells (GSCs) GSH. See Glutathione (GSH) GSiXX. See Gamma secretase inhibitor XX (GSiXX) GSK inhibitor CHIR99021, 119 GSK-3b. See Glycogen synthase kinase 3b (GSK-3b) GSMS. See Gastric submucosal space (GSMS) GTPase. See Guanosine triphosphatase (GTPase) Guanine nucleotide exchange factor (GEF), 20 Guanosine triphosphatase (GTPase), 438, 635e636 GTPase Rab5c, 6 Guidance documents in FDA, 1347 Guide RNA (gRNA), 741e742

Guided tissue regeneration, 1133, 1136 a-L-Guluronic acid, 640e641 Gut endocrine cells, 1140e1141 GVAX vaccine, 716 GvHD. See Graft versus host disease (GvHD) GVT effect. See Graft versus tumor effect (GVT effect) GWAS. See Genome-wide association studies (GWAS)

H 3

H-T labeling index, 42 H3K27me3 marks, 55 H3K4me3 mark, 57 H3K9 methyltransferase inhibitor G9a, 55 HA. See Hemophilia A (HA); Hyaluronan (HA) HA methylcellulose (HAMC), 1205e1206 HAc. See Hyaluronic acid (HAc) hAE stem cells. See Human amnion epithelial stem cells (hAE stem cells) HAEC. See Human aortic endothelial cells (HAEC) hAECs. See Human amniotic epithelial cells (hAECs) Hair cell(s), 872 regeneration, 869, 870f clinical trial, 881 formation of new neuromasts, 877 hair cell loss, 868e869 induction of hair cell regeneration, 875e876 insights from developmental biology, 871e875 lateral line, 876e879 lateral line regeneration, 880 pathways coordinating hair cell regeneration, 879e880 regulation of cell fates during inner ear development, 873f road blocks to regeneration, 871 spontaneous hair cell regeneration, 870e871 structure of inner ear, 867e868 Hair follicle, 66, 84 cycling, 1300 formation, 1297 regeneration, 1298e1299 Hair follicle stem cells (HFSCs), 84, 1264 Hair germ progenitors, 1298 Hair growth additional studies on, 1300 ADSCs use and conditioned medium for, 1299 Hair regeneration, tissue-derived materials for, 1300 Hair shaft, 1297 Hair-stimulating complex (HSC), 1301, 1302fe1303f HALSS. See Hybrid artificial liver support system (HALSS) hAM. See Human amniotic membrane (hAM) HAMC. See HA methylcellulose (HAMC)

INDEX

hAMSCs. See Human amniotic mesenchymal stromal cells (hAMSCs) Hamstring tendon autografts, 1179 HAp. See Hydroxyapatite (HAp) Hard assignment, 103 Harsh synthetic chemistry, 647 HATs. See Histone acetyltransferases (HATs) HB. See Hemophilia B (HB) HB-EGF. See Heparin-binding endothelial growth factor (HB-EGF) HbA1C. See Glycosylated hemoglobin (HbA1C) HBB. See b Subunits of hemoglobin (HBB) hbb gene, 753 hBMSC. See Human bone marrow-derived stem cells (hBMSC) HCC. See Hepatocellular carcinoma (HCC) hCFs. See Human corneal fibroblasts (hCFs) hCG. See Human chorionic gonadotropin (hCG) HCM. See Hypoxia-derived CM (HCM) hCMPCs. See Human cardiac-derived cardiomyocyte progenitor cells (hCMPCs) hCMSCs. See Human chorionic mesenchymal stromal cells (hCMSCs) hCSSCs. See Human corneal stromal stem cells (hCSSCs) HCT/P. See Human cells, tissues, and cellular-and tissue-based products (HCT/P) hCTCs. See Human chorionic trophoblastic cells (hCTCs) HCV. See Hepatitis C virus (HCV) HD. See Hemodialysis (HD) HDACs. See Histone deacetylases (HDACs) HDE. See Humanitarian Device Exemption (HDE) HDF. See Hemodiafiltration (HDF); Human dermal fibroblasts (HDF) hDPSCs. See human dental pulp stem cells (hDPSCs) HDR. See Homology-directed repair (HDR) HE4. See Human epididymis protein 4 (HE4) Healing of cutaneous wounds, 25e28 of ligaments and tendons, 1183e1185 ACL healing, 1189e1192 anterior cruciate ligament of knee, 1184e1185 medial collateral ligament and patellar tendon healing, 1188e1189 medial collateral ligament of knee, 1184 multiple ligamentous injuries in knee, 1185 use of scaffolds, 1188 progresses, 27

Health Resources and Services Administration, 1353 Heart, 141e142, 247, 450, 846 development from cardiac stem/ progenitor cells, 247e248 regeneration, 248e250 valves, 846 cell-based therapeutics for heart disease, 252e254, 253f Heart failure (HF), 252e253, 495 Heart Failure Secondary to Myocardial Infarction trials, 258 Heart valve disease (HVD), 1041 Heart-on-a-chip model, 774, 1081 Hedgehog pathway, 7 Helixeloopehelix (HLH), 872 Helper-dependent adenoviral vectors, 749 HEMA. See 2-Hydroxyethyl methacrylate (HEMA) Hemangioblasts, 311 system, 926e927 HEMAPLA. See Polylactideehydroxyethyl methacrylate (HEMAPLA) Hematological malignancies, CBT for, 156 Hematoma, 696 Hematopoiesis, 307e308, 1010 Hematopoietic antigens, 183 Hematopoietic cancers, 931e932 Hematopoietic lineages, 307e308 markers, 258 Hematopoietic malignancies, 1253 Hematopoietic reconstitution with MAPCs, 183 Hematopoietic stem cell transplantation (HSCT), 149 for autoimmune diseases, 199 for severe combined immunodeficiency, 197e198 for tolerance induction, 198e199 Hematopoietic stem cells (HSCs), 149, 181e183, 191, 205, 219, 257, 274, 307e308, 753, 931e933, 963, 1009e1010 phenotypic properties, 194e195 properties, 191e195 therapies, 195e199 Hematopoietic system, 142, 191, 1020 Hematoxylineeosin staining, 700 Heme oxygenase-1 (HO-1), 310 Hemocompatibility, 515 Hemodiafiltration (HDF), 1149 Hemodialysis (HD), 1149 Hemodynamic forces, 424 Hemofilters, 1150 Hemofiltration (HF), 1149 b-Hemoglobinopathies, 753 Hemolysis, 515 Hemophilia A (HA), 322e323, 816, 1015, 1169 HA-based materials, 618 as model genetic disease for correction, 1015e1020 feasibility and justification for treating HA before birth, 1016e1018

1397 genome editing, 1019e1020 genomic integration-associated insertional mutagenesis, 1018e1019 potential risk to fetal germline, 1019 preclinical animal models to study in utero gene therapy, 1016 risks of in utero gene therapy, 1018 treatments, 1015e1016 Hemophilia B (HB), 1016 Hemophiliac patients, 323 Hemorrhage, 1151e1152 Hemorrhagic stroke, 1203 Hemostasis, 38e39 Hep3B spheroid, 1106 HepaChip in vitro microfluidic system, 1108e1109 Heparan sulfate, 617e618 proteoglycans, 279e280 Heparan sulfate proteoglycans (HSPGs), 16 Heparin, 1151e1152 Heparin-binding endothelial growth factor (HB-EGF), 16 Hepatectomy, 238 Hepatic diverticulum, 341 Hepatic stellate cells (HSCs), 1106 Hepatic tissue engineering, 1103e1106 alginate, 1104 biomedical applications of liver bioengineered tissues, 1105f chitosan, 1103 collagens, 1103e1104 decellularized extracellular matrix, 1104e1106 PGA, 1104 PLGA, 1104 polycaprolactone, 1104 Hepatitis C virus (HCV), 1101 Hepatocellular carcinoma (HCC), 1101 spheroids, 1106 Hepatocyte allografts, 234 Hepatocyte bridge, 233 Hepatocyte growth factor (HGF), 7, 205, 238e239, 280, 298, 339e340, 751, 1299e1300, 1302e1303 Hepatocyte transplantation, 229e230 in acute liver failure, 233e234 choice of sites for, 231 clinical, 232e237, 233t in acute liver failure, 233e234 hepatocyte bridge, 233 for metabolic liver disease, 234e237 for metabolic liver disease, 234e237 novel uses, challenges, and future directions cell transplantation immunology, 239 hepatocyte transplants for noneorgan transplant candidates, 237e238 methods to improve engraftment and repopulation, 238e239 stem cells and alternative cell sources for liver therapy, 239e241 Hepatocyte-based therapy, 241 Hepatocyte-like cells, 240 Hepatocytes, 229e230, 321e322

1398 Hepatocytes integration after transplantation, 231e232 integration of donor hepatocytes into native liver, 231t “HepatoPac” platform, 1108e1109 Herpes simplex virus serotype 1 (HSV-1), 1125 hESCs. See Human embryonic stem cells (hESCs) Heterochromatin, 55 Hexafluoro-2-propanol (HFIP), 547, 568, 699e700 Hexafluoroisopropanol. See Hexafluoro-2propanol (HFIP) Hexahistidine tags (HIS tag), 528 HF. See Heart failure (HF); Hemofiltration (HF) HFD. See High-fat diet (HFD) HFEA. See Human Fertilisation and Embryology Authority (HFEA) HFIP. See Hexafluoro-2-propanol (HFIP) HFSCs. See Hair follicle stem cells (HFSCs) HGF. See Hepatocyte growth factor (HGF) HHP. See High hydrostatic pressure (HHP) HIE. See Hypoxic-ischemic encephalopathy (HIE) Hierarchical 3D tissue equivalents, 635 Hierarchical clustering, 101e102, 102f HIF-1. See Hypoxia-inducible factor-1 (HIF-1) High hydrostatic pressure (HHP), 1241 High printing resolution, 833 High surface-to-volume ratio, 491e492 High-affinity interactions, 657 High-energy/high-temperature plasmas, 656 High-fat diet (HFD), 281 “High-flux” membranes, 1151 High-Oct4 (Oct4high), 184 High-resolution extrusion, 818 High-resolution ultrasonography, 1009 High-resolution XPS spectra, 525 High-throughput molecular techniques, 1009 High-throughput screening, 174 High-throughput sequencing (HTS), 93 Higher-level tissue function, 1065 Higher-molecular weight materials, 666 Highly functional multiorganoid systems, 779e783 HIMSCs. See Hypoxia-induced multipotent stem cells (HIMSCs) hiPCs. See Human islet-derived precursor cells (hiPCs) hiPSCs. See Human induced pluripotent stem cells (hiPSCs) Hirudin, 1151e1152 HIS tag. See Hexahistidine tags (HIS tag) Histogenesis in three-dimensional scaffolds, 661 design parameters for, 663e668 biomolecular factors, 667 cell sources, 663e664 degradation, 666 porosity, 664e665

INDEX

future directions, 670 need for replacement tissues, 661 regeneration of diseased tissues, 662e663 synthetic materials for histogenesis of new organs, 669e670 tissue components, 661e662, 662f Histolysis, 39e40 Histone acetylation, 57 modifications, 100 Histone acetyltransferases (HATs), 55 Histone deacetylases (HDACs), 54 inhibitors, 55 HLA. See Human leukocyte antigen (HLA) HLH. See Helixeloopehelix (HLH) hMAPCs. See Human MAPCs (hMAPCs) hMSCs. See Human mesenchymal stem cells (hMSCs) HNF3ß. See FOXA2 HNH domains, 742 HO-1. See Heme oxygenase-1 (HO-1) HOBs. See Human osteoblast-like cells (HOBs) Hoechst 33, 342, 195 Hollow-fiber reactors, 1109e1110 Homeostasis, 371e373 Hominis placenta (HP), 1300 Homology-directed repair (HDR), 742e745, 744f Hooke law, 422 Hormones, 859 contraceptive, 561 ovarian, 1242 PTH, 763e764, 860, 861f rhGH, 690 Host cells, 663e664 defense system, 680 response, 621, 1088e1090 RNase, 745 Hot-embossing technique, 447t Housekeeping genes, 101 HP. See Hominis placenta (HP) HPMC. See Hydroxypropyl methyl cellulose (HPMC) hPSCs. See Human pluripotent stem cells (hPSCs) HSC. See Hair-stimulating complex (HSC) HSCs. See Hematopoietic stem cells (HSCs); Hepatic stellate cells (HSCs) HSCT. See Hematopoietic stem cell transplantation (HSCT) HSPA8, 209 HSPGs. See Heparan sulfate proteoglycans (HSPGs) HSV-1. See Herpes simplex virus serotype 1 (HSV-1) HTO. See Human testis organoid (HTO) HTS. See High-throughput sequencing (HTS) HUASMCs. See Human umbilical artery smooth muscle cells (HUASMCs)

hUCBCs. See Human umbilical cord blood cells (hUCBCs) hUCMSCs. See human umbilical cord MSCs (hUCMSCs) Human allogeneic uterus transplantation, 1240 amnion, 134 cells regulation, 1350e1351 cellular therapies, 1351e1353 engineered heart tissue strips, 1092 eye, 352 fetal placental cells, 134 fibroblasts, 449 glioblastomas, 108 nonesmall cell lung cancer, 775 oligodendrocyte precursor cells, 82 placental extract. See Hominis placenta (HP) recombinant collagen, 541 saliva, 898e899 stem cells to creating chimeras, 1335e1336 therapeutic cloning, 1334 tissues engineering nanostructured scaffolds, 487f nanoscale features, 485, 486f regulation, 1350e1351 urothelial cells, 1273 Human amnion epithelial stem cells (hAE stem cells), 240 Human amniotic epithelial cells (hAECs), 133, 241, 700 Human amniotic membrane (hAM), 135, 1118 Human amniotic mesenchymal stromal cells (hAMSCs), 133 Human aortic endothelial cells (HAEC), 312, 823e824 Human bone marrow-derived stem cells (hBMSC), 703, 1264 hBMSCs plus hematopoietic stem and progenitor cells, 1264 Human cardiac-derived cardiomyocyte progenitor cells (hCMPCs), 815e816 Human cells, tissues, and cellular-and tissue-based products (HCT/P), 120, 622 Human chorionic gonadotropin (hCG), 133e134, 1203e1204, 1254 Human chorionic mesenchymal stromal cells (hCMSCs), 133 Human chorionic trophoblastic cells (hCTCs), 133 Human corneal fibroblasts (hCFs), 1125 Human corneal stromal stem cells (hCSSCs), 1125 human dental pulp stem cells (hDPSCs), 700 Human dermal fibroblasts (HDF), 299e300 Human diphtheria toxin receptor (Human Dtr), 870e871

INDEX

Human Dtr. See Human diphtheria toxin receptor (Human Dtr) Human embryonic MSC-derived exosomes, 211 Human embryonic stem cell self-renewal regulation, 52 Human embryonic stem cells (hESCs), 50, 114, 125, 169, 181, 338e339, 351, 445, 617, 762, 855, 923e927, 926fe927f, 996, 1119, 1202, 1300e1301, 1310, 1331. See also Embryonic stem cells (ESC) alternative sources, 125 SBB, 126 differentiation and manufacturing for clinical application, 119e120 evolution of hESC derivation and culture methods, 118e119 hES-NCL9, 128 hESC-derived cardiomyocytes, 1081 hESC-derived mesenchymal progenitor cells, 855 hESCederived retinal pigment epithelium, 355e356, 357f maintenance, 116e119 morphology, 118 organismically dead embryos, 127e128 sources, 114e116 Human embryos, 1331e1332 benefit from others’ destruction, 1333e1334 cloning, 1334e1335 genetically modifying, 1336 permissible to destroying, 1332e1334 Human epididymis protein 4 (HE4), 1173 Human Fertilisation and Embryology Authority (HFEA), 1320 Human hair follicle, regenerative medicine approaches for engineering, 1298f additional studies on secreted growth factors and hair growth, 1300 adipose-derived stem cells use and conditioned medium, 1299 autologous growth factors use in hair follicle regeneration, 1298e1299 bioengineering human hair follicle, 1304e1306 simulating embryonic environment, 1300e1304 tissue-derived materials for hair regeneration, 1300 Human induced pluripotent stem cells (hiPSCs), 927e928, 1080, 1107e1108 human iPSC-derived cardiomyocytes, 1081e1082 Human islet-derived precursor cells (hiPCs), 1335 Human leukocyte antigen (HLA), 115e116, 134e135, 149, 175, 182, 195e196, 315e316, 356, 688e689, 923 HLA-DR3 expression, 688e689 human leukocyte antigen-II surface expression, 298

human leukocyte antigen-matched tissue, 1118 Human MAPCs (hMAPCs), 182, 187 in vitro, 183 immunomodulatory properties, 184e187 isolation, 183 suppress T-cell proliferation, 184 Human mesenchymal stem cells (hMSCs), 205, 219, 223, 377, 444, 489, 945, 1258 hMSC-injected mice, 223 Human multipotent stromal cells. See Human mesenchymal stem cells (hMSCs) Human osteoblast-like cells (HOBs), 706e707 Human pluripotent stem cells (hPSCs), 340, 441, 450, 930 Human testis organoid (HTO), 1253e1254, 1253f Human umbilical artery smooth muscle cells (HUASMCs), 441 Human umbilical cord blood cells (hUCBCs), 376 human umbilical cord MSCs (hUCMSCs), 706 Human umbilical tissue-derived stem cells (hUTCs), 351 Human umbilical vein endothelial cells (HUVECs), 312, 441, 495, 706e707, 836e837, 913e914 HUVEC-laden patch, 1080 Humanitarian Device Exemption (HDE), 1348e1349 Humoral components, 684e685 Hurler syndrome, 158 hUTCs. See Human umbilical tissuederived stem cells (hUTCs) HUVECs. See Human umbilical vein endothelial cells (HUVECs) HVD. See Heart valve disease (HVD) Hyaline cartilage, 937 Hyaluronan (HA), 21, 29, 618, 639, 645e646, 816, 1205e1206 HA-based tubular conduits, 1226 HA-induced proliferation, 23 Hyaluronic acid (HAc), 617e618, 639e640, 699, 701e702, 810e811, 836e837, 938e940, 1048, 1169, 1226 HAc-AEMA, 702 HAc-AEMA-40 hydrogel, 702 HAcebased scaffolds, 913e914 Hybrid artificial liver support system (HALSS), 1107 Hybrid bioartificial kidney, 1159 Hybrid bioinks, 816e819, 818f Hybrid composite biomaterials functions of scaffolding and ECM, 696e697 fundamentals of bone development and defects, 696 hybrid materials, 708e709 scaffolding approaches in bone tissue engineering, 697e699 scaffolding materials, 699e709

1399 Hybrid fabrication, 834e835 Hybrid materials, 698, 708e709 metaleceramic blends, 709 metalepolymer blends, 709 polymereceramics blends, 708e709 polymerepolymer blends, 708 Hybrid scaffolds, 1033e1037 extracellular matrices formed by cell culture, 1033e1034 nature-derived polymers and synthetic polymers, 1034e1035 synthetic polymers with seeded cells, 1035e1037 Hybrid toothebone implants, 916 Hydraulic permeability of collagen-based scaffold, 701 Hydrodynamic delivery, 749 Hydrogel scaffolds, 452e453 for nerve repair, 1226 classification of nerve grafts, 1227t classification of nerve guidance conduits, 1227t Hydrogel-based bioinks for cell printing, 835e837 naturally derived hydrogels, 836e837 synthetic hydrogels, 835e836 Hydrogel(s), 285, 380, 553e554, 560, 620e621, 627, 630, 669e670, 697e699, 703e704, 763e764, 773, 819e820, 858e859, 910e911, 1075, 1208 biomaterials templates, 628e630 conventional porous solid polymer tissue engineering templates, 628f degradation, 636e637 rates, 637 hydrogel-based cardiac tissue engineering, 1076 hydrogel-forming macromonomers, 580 increasing sophistication of synthetic hydrogels for TE, 632e639 lubricates and glues, 834e835 natural biopolymers as extracellular matrixeanalog hydrogels, 639e646 naturally derived, 836e837 in regenerative medicine, 627 in situ cardiac tissue engineering via injection of cells in, 1090e1091 structureeproperty relationships in, 631e632 synthetic, 835e836 synthetic hydrogels for tissue engineering templates, 646e648 Hydrogen bonding, 645 peroxide, 552 Hydrolysis interaction, 596e597 Hydrolytic degradation products, 575 Hydrolytically degradable polymers, 669, 1050e1051 Hydrolytically stable PUs, 566e567 Hydrophilic macromolecular drugs, 579 Hydrophilic polymers, 657e659, 720e721 Hydrophilic surface, 441, 703

1400 Hydrophilicity, effects of nanoparticle surface charge and, 720 Hydrophobicity, 720 Hydrothermal method, 551e552 a-Hydroxy acids, polyesters of, 568e570 POEs, 571 polycarbonates, 572 polyesters of lactones, 570 polyesters of polyols and carboxylic acids, 571 Hydroxyapatite (HAp), 409e411, 531, 540, 551e552, 591e592, 656, 696, 858, 908e909, 955 HA for protein signal delivery and orientation control, 531 HAp-chitosanegelatin nanocomposite scaffolds, 540 microspheres, 540 nanoparticles, 488, 493 Hydroxyethyl methacrylate J2-J (HEMA), 631e632, 669 Hydroxyl group, 568 Hydroxypropyl methyl cellulose (HPMC), 600 bis(2-Hydroxypropyl) fumarate (PF), 576e577 Hyperammonemia, 750 Hyperbilirubinemia, 236 Hyperelastic bone, 813 Hypersensitivity responses, 686 Hypertrophic scars, 28, 66, 70e72 scar reduction strategies, 71f Hypertrophy process, 957 Hypogonadal disorders, 1254 Hypogonadism, 1251 Hypothalamus, 1251e1252 Hypoxia, 7e8, 232, 1286e1287 Hypoxia-derived CM (HCM), 1299 Hypoxia-induced multipotent stem cells (HIMSCs), 1298 HIMSC-CM, 1301, 1301f Hypoxia-inducible factor-1 (HIF-1), 7e8, 699e700 Hypoxic-ischemic encephalopathy (HIE), 159 Hysteresis, 1181e1182 HmREL microliver platforms, 1108e1109

I

IAC. See Immune-affinity capture (IAC) IAS. See Internal anal sphincter (IAS) IBMIR. See Instant blood-mediated inflammatory reaction (IBMIR) ICAMs. See Intracellular adhesion molecules (ICAMs) ICH. See International Council for Harmonization of Technical Requirements for Pharmaceuticals for Human Use (ICH) ICIs. See Immune checkpoint inhibitors (ICIs) ICM. See Inner cell mass (ICM) ICMVs. See Interbilayer-crosslinked multilamellar vesicles (ICMVs) ICP. See Intracranial pressure (ICP)

INDEX

ICSI. See Intracytoplasmic sperm injection (ICSI) ICU. See Intensive care unit (ICU) Id1. See Inhibitor of Differentiation 1 (Id1) IDE. See Investigational Device Exemption (IDE) IE. See Islet equivalents (IE) iENDO cells. See Induced endodermal progenitor cells (iENDO cells) IFE. See Interfollicular epidermis (IFE) IFN. See Interferon (IFN) IFs. See Intermediate filaments (IFs) IgE antibodies, 686e687 IGF. See Insulin growth factor (IGF) IGFBP. See Insulin-like growth factorbinding protein (IGFBP) IGFR1, 7 IgG. See Immunoglobulin G (IgG) aIIb/bIII integrin receptor. See Cell surface markers GPIIb/IIIa IKVAV. See Isoleucine-lysine-valinealanine-valine (IKVAV) IL. See Interleukin (IL) IL receptor (ILR), 992 ILK. See Integrin-linked kinase (ILK) ILR. See IL receptor (ILR) Imaging data, 382e383 IMD. See Inherited metabolic disorders (IMD) Imiquimod, 78 Immature follicles, 1243 Immobilization strategy, 443e444 Immune checkpoint inhibitors (ICIs), 716e717 Immune complex reaction, 686 Immune mediators, 715e716 Immune rejection, 1334 Immune system, 715, 938 Immune tolerance, 198e199 Immune Tolerance Network (ITN), 992 Immune-affinity capture (IAC), 209 Immunocompatible iPSC master cell banks development, 255e256 Immunocytochemical studies, 1257 Immunodepletion, 208 Immunodulatory effects, in vivo, 184e186 antitumor effects of MAPCs in glioma, 186 MAPCs effect on graft survival cotransplanted with cells, 185 effect on GvHD, 184 immunodulatory and/or trophic effects in ischemic disease, 185e186 role as immunomodulation in solid organ transplantation, 185 possible mechanisms of trophic effects, 186e187 Immunoediting, 724 Immunogenicity, 689e690, 748e749 Immunoglobulin G (IgG), 528e529, 679, 681 IgG2a, 731 Immunohistochemistry, 1300 Immunohistological studies, 263e264

Immunoisolation, 208 Immunologic/immunology, 1016 escape, 724 memory, 717 reactions, 687 Immunomodulation, 209, 344f, 999e1000 in bone regeneration, 861e862 macrophages, 862 T cells, 862 in solid organ transplantation, 185 Immunomodulatory agents, bolus delivery of, 721e722 Immunomodulatory effect of MSCs, 211e212 of renal assist device, 1154e1155 Immunomodulatory materials, 698 Immunostimulation, 688 Immunostimulatory effect, 184 Immunosuppressants, 1088e1089 Immunosuppression, 688, 982, 987e988 Immunosuppressive protocol, 239 Immunosuppressive therapy, 239, 790e791 and complications, 992e993 Immunotherapy nanomedicine applications in, 727 nanoparticle targeting applications in, 721 Immunotoxicity, 675, 684e691 tests, 689 Impedance spectroscopy, 597e598 Implanon, 562 Implant design goals, 1045 Implant function, 1044e1045 Implantable biomaterial scaffolds as cancer vaccines, 727e729, 728f to enhance autologous T cell therapy, 731e733 Implantable scaffolds, 891e892 ceramics, 891e892 polymers, 892 Implanted stem cells secrete factors, 1267e1268 Imprinting technique, 562e563 iMSCs. See MSCs derived from induced pluripotent stem cells (iMSCs) In situ cardiac tissue engineering via injection of cells in hydrogels, 1090e1091 In situ DC vaccination approach, 727 In situ gelation, 574, 581 In situ kidney regeneration, 1166 In situ renal regeneration, 1172e1173 In situ-forming hydrogels, 638e639 materials, 648 In utero gene therapy (IUGT), 1009, 1015e1020 In utero stem cell transplantation (IUTx), 1009 clinical experience with, 1020e1021 preclinical animal studies, 1011e1020 barriers, 1013e1015 preclinical animal studies of IUGT, 1015e1020

INDEX

In utero transplantation (IUT), 142 In vitro applications, 825e826 cell viability, 1157 culture system, 930e931 degradation profile, 1172 expansion of EPC, 312e313 of MSC, 315e316 experiments, 1152 hematopoiesis, 193e194 models, 776, 1142e1143 organoid development, 770e771 platelet production, 930e931 preclinical models, 764 spermatogenesis models, 1253e1254 studies, 73, 459e460, 1185e1186 systems, 457 ventricle chamber, 1083 tissue models, 839e841 drug screening systems, 841 tumor models, 839 In vitro fertilization (IVF), 100, 115, 1255, 1310, 1332 embryos, 1310 protocol, 1245 In vivo application, 438 3D bioprinting, 826 bioreactors, 897e898 for lung regeneration, 790e792, 792f technologies, 956 bone bioreactors for solving vascularization problem, 797e798, 798f chondrogenic bioassay, 412 conditioning and testing, 1046 implantation studies, 1169 murine models, 376 preclinical models, 764e765 studies, 1090e1094, 1168, 1186, 1191 delivery route for cells and patches, 1094 implantation of cardiac patches, 1091e1093 in situ cardiac tissue engineering via injection of cells in hydrogels, 1090e1091 in vivo integration of engineered cardiac patches, 1092f tests, 540 IND. See Investigational New Drug (IND) Indium phosphide, 490e491 Induced endodermal progenitor cells (iENDO cells), 183 Induced hematopoietic progenitors, 932e933 Induced neurons (iN), 450 Induced pluripotent stem (iPS), 49, 912e913 iPSederived muscle stem cells, 286

Induced pluripotent stem cells (iPSCs), 40, 50, 84, 169, 181, 212, 247, 255e256, 256f, 285e286, 335e341, 351, 445, 474e475, 617, 747, 750, 762, 780e782, 855e856, 923e924, 931fe932f, 954, 1063, 1168, 1203, 1245, 1264, 1286e1287, 1309, 1311, 1331 applications, 172, 173f cell therapy, 175e176 challenges and future possibilities in disease modeling, 174 conservation of endangered species, 176 disease modeling, 172e174 epigenetic remodeling, 170 future directions, 176e177 genomic stability, 172 induced transdifferentiation, 171e172 iPSC-based therapeutics, 255e256 iPSC-derived cardiomyocytes, 1076 iPSCebased phenotyping, 363 iPSCederived retinal pigment epithelium, 356e359 iPSCsederived MSCs, 212e213 mechanisms of reprogramming, 169e170 model of cardiogenesis, 248e250 systems, 175 personalized medicine, 174e175 potential therapies for reducing scar formation, 85f reprogramming techniques, 170e171, 171f Induced transdifferentiation, 171e172 Inducible nitric oxide synthase inhibitors, 945 Induction of donor-specific tolerance, 198e199 Inductive morphogenetic signals, 405e406 Indwelling urethral catheterization, 1273 Infertility, 1251 Inflammation, 297, 916, 1263 and wound healing, 676e683 acute inflammation, 679 bloodematerial interactions and initiation, 676e677 chronic inflammation, 679e680 FBGC cell formation and interactions, 682e683 granulation tissue, 680 macrophage interactions, 680e681 provisional matrix formation, 677 sequence of host reactions, 676t temporal sequence, 678, 678f Inflammatory bowel disease, 220, 223 Inflammatory phase of wound healing, 26 Inflammatory response initiation, 676e677 targeting, 73e75 Informed consent, 1338 Infrared (IR), 525e526 Inherited metabolic disorders (IMD), 157 CBT for, 157e159 Inherited retinal dystrophies, 354e355 Inhibitor of Differentiation 1 (Id1), 52

1401 Initial survival, cell-graft survival, 980e981 Injectability, 599e600 strategies to improve injectability, 599e600 particle size and shape, 600 viscosity, 600 Injectable polymer systems, 575e576 Injectable scaffolds, 858e859 hydrogels, 858e859 injectable scaffolds/controlling morphology in situ, 513 microspheres, 858 Injectable systems, 638e639 Injection molding, 447t Inkjet bioprinting, 807e808 “Inkjet” printing. See Jetting-based printing Innate or adaptive immune cells, 938 Inner cell mass (ICM), 53e54, 113e114, 181 Inner ear, 867e868, 868f Inner root sheath, 1297 INO80 chromatin remodeler, 57 Inorganic based nanobiomaterials, 492e493, 492t Inorganic materials, 916 Inosculation, 668 Inositol triphosphate (IP3), 38e39 INS. See Insulin (INS) Insertional mutagenesis, 1018 “Inside-out” signaling, 19 Insolubility of elastin, 644 Insoluble cellular fibronectin, 616 Instant blood-mediated inflammatory reaction (IBMIR), 997 Institutional Animal Care and Use Committee, 765e766 Institutional Review Board (IRB), 1310, 1337 Insulin (INS), 31, 339e340, 987 Insulin growth factor (IGF), 7, 311e312, 339e340, 942 IGF-1, 23, 280, 1116, 1300 Insulin-like growth factor-binding protein (IGFBP), 344 Intact ECM molecules, 25e26 as scaffold material decellularization, 619e620 host response, 621 hydrogels, 620e621 postprocessing, 620 tissue procurement, 619 whole-organ scaffolds, 621e622 Integra, 1287e1290 Integrase, 748 Integrated multiorganoid model systems, 778e779 Integrated system, 780 Integration-free methods, 170 a7-Integrin (ITGA), 277 Integrin-linked kinase (ILK), 5 Integrins, 3, 16, 279e280, 425, 679 b6 gene, 3 integrin-mediated cell migration, 445

1402 Integrins (Continued ) integrin-mediated ECM signaling, 19e20 integrinematrix binding, 22 b1 integrins, 279e280, 617 b4 integrins, 617 Intellectual property (IP), 231 Intelligence of thermoresponsive polymers for cell sheet engineering, 469e471 Intelligent surfaces for regulating cell orientation, 479e480 Intelligent thermoresponsive cell culture substrates, 469 Intensive care unit (ICU), 1154 Interbilayer-crosslinked multilamellar vesicles (ICMVs), 723e724 Interference RNA (RNAi), 911 Interferon (IFN), 184, 715e716 INF-gamma, 184, 383, 1154e1155 Interfollicular epidermis (IFE), 84 Interleukin (IL), 77, 184, 344, 370, 992, 1062 IL-1, 66e67, 280, 679, 888, 1087 IL-1b, 25, 799, 862 IL-2, 205, 716 IL-4, 298 IL-6, 25, 77, 280, 862, 1154e1155, 1268 IL-8, 77, 298, 1268 IL-10, 77, 205, 211, 298, 1154e1155 IL-13, 682 IL-24, 799 Intermediate filaments (IFs), 394e395 Intermolecular interactions, 655 Internal anal sphincter (IAS), 1142 International Council for Harmonization of Technical Requirements for Pharmaceuticals for Human Use (ICH), 1347 International efforts, 1374e1375 International Knockout Mouse Consortium, 747e748 International NetCord Foundation, 149e150 International Society for Cellular Therapy, 149e150, 182, 315e316 International Society for Cytotherapy, 206 International Society for Extracellular Vesicle, 208 International Society for Stem Cell Research (ISSCR), 1318, 1336 International Standard Organization (ISO), 1358, 1368 Interneuromast cells, 877 Interstitial matrix, 613e615 Intervertebral disc (IVD), 962 Intestinal microenvironment, 1139 Intestinal mucosa, 1136 Intestinal smooth muscle cells, 1139 Intestinal tissue engineering, 1136, 1138 Intestine, 143e144 Intracellular adhesion molecules (ICAMs), 679 Intracellular calcium flux, 688 Intracellular signaling molecules, 7, 16e18

INDEX

Intracoronary cardiosphere-derived autologous stem cells, 265 Intracranial pressure (ICP), 369e370 Intracytoplasmic sperm injection (ICSI), 1251 Intradiscal delivery method, 962e963 Intralesional corticosteroid injections, topical and, 77 Intramembraneous bone formation, 696 Intramembranous ossification, 853 Intramuscular accumulations of grafted SCDMs, 981 Intramuscular transplantation, 976, 977f Intramyocardial injection of MSCs, 258 of myoblasts, 261 Intraperitoneal transplantation (IP transplantation), 324 Intraportal infusion of islets, 997 Intrathecal injection, 1207 Intratumor heterogeneity, 93e94 Intravascular administration, 976 Intravenous infusion, 1061e1062 Intravital imaging (IVM), 278e279 Inverse opal, 547 Invertebrates, 538 Inverted terminal repeats (ITRs), 747 Investigational Device Exemption (IDE), 1349 Investigational New Drug (IND), 1301e1302, 1338, 1349 Iodine 125elabeled proteins (I125elabeled proteins), 527 Ion beam implantation, 653 Ion channels, 393e394 Ionic charge, 528e529 Ionic homeostasis, 370e371 Ionic polyphosphazenes, 575 IP. See Intellectual property (IP) IP transplantation. See Intraperitoneal transplantation (IP transplantation) IP3. See Inositol triphosphate (IP3) Ipilimumab, phase III clinical trials of, 716e717 iPS. See Induced pluripotent stem (iPS) iPSCs. See Induced pluripotent stem cells (iPSCs) IR. See Infrared (IR) IR injuries. See Ischemia-reperfusion injuries (IR injuries) IRB. See Institutional Review Board (IRB) Iron oxide nanoparticles, 492 g-Irradiation, 640e641 Irreversibility as criterion for diagnosing embryonic death, 127e128 Irreversibly arrested, nonviable embryos human embryonic stem cell lines deriving from, 128 morphological criteria for predicting capacity, 128 Ischemia-reperfusion injuries (IR injuries), 137e138, 186 Ischemic disease, MAPC immunodulatory/trophic effects in, 185e186

ischemiaereperfusion injury, 186 ischemic stroke, 185 myocardial infarct, 186 peripheral hind limb ischemia, 186 spinal cord injury, 185 TBI, 185e186 Ischemic necrosis, 981, 981f Ischemic stroke, 185, 1203 Isl-1+ cardioblast, 250, 263e264 Isl-1+ progenitors, 263e264 Islet cell transplantation. See also Cell transplantation; Cord blood transplantation (CBT); Hepatocyte transplantation; In utero transplantation (IUT) clinical islet transplantation, 990e993 edmonton protocol, 989e990 engraftment posttransplant, 998e999 immunomodulation, 999e1000 islet transplantation, 988e989 living donor islet transplantation, 993e994 optimal transplantation site, 997e998 overcoming tissue shortage, 993 stem cell transplantation, 995e997 xenotransplantation, 995 Islet equivalents (IE), 989 Islets, 336 cell proliferation, 343 isolation process, 997 transplantation, 988e989, 988f, 991f clinical, 990e993 patient assessment and selection, 990e991 procedure, 991e992 risks to recipient, 992e993 ISO. See International Standard Organization (ISO) Isolate morphogens, 406, 406f Isolated cells, 742 Isoleucine-lysine-valine-alanine-valine (IKVAV), 380 amino acid sequence, 635 N-Isopropylacrylamide (NiPAAm), 563e564, 648 Isoproterenol, 780 Isothermal methods, 97 Isotropic grafts, 1224 Isotropic natural materials, 1224e1225 Isotropic scaffolds for nerve regeneration ECM molecules, 1226e1227 electroconductive scaffolds, 1229 hydrogel scaffold for nerve repair, 1226 natural scaffolds, 1224e1225 neurotrophic factors and cytokine delivery, 1227e1228 seeding neuronal support cells, 1228e1229 synthetic scaffolds for nerve repair, 1225e1226 ISSCR. See International Society for Stem Cell Research (ISSCR) ITGA. See a7-Integrin (ITGA) ITN. See Immune Tolerance Network (ITN)

INDEX

ITRs. See Inverted terminal repeats (ITRs) IUGT. See In utero gene therapy (IUGT) IUT. See In utero transplantation (IUT) IUTx. See In utero stem cell transplantation (IUTx) Ivacaftor, 174e175 IVD. See Intervertebral disc (IVD) IVF. See In vitro fertilization (IVF) IVM. See Intravital imaging (IVM)

J

JAK signaling, 57 Japan, advancing regenerative medicine, 1375 Japan’s Pharmaceuticals and Medical Devices Agency, 1375 Jetting-based printing, 832e833, 1288e1289 3D multicell “pie” construct, 833f Jmjd1a and Jmjd2c. See Jumanji domain enzymes (Jmjd1a and Jmjd2c) JNK. See Jun N-terminal kinase (JNK) Joint FDAeCDC workshop, 1351 Joint function, contribution to, 1182e1183 Joint motion, 1182e1183 Jumanji domain enzymes (Jmjd1a and Jmjd2c), 55 Jun N-terminal kinase (JNK), 398 JUVENTAS trial, 83

K

K/Sr scaffolds. See Potassium and strontium ions scaffolds (K/Sr scaffolds) KADD. See N-Aminoethyl aminocaproyl dihydrocinnamoyl (KADD) Kanamycin, 870e871 Karyotype analysis of human AFSC, 138 KASH domains. See Klarsicht, ANC-1, Syne homology domains (KASH domains) KASHeSUN complex, 459 KDR. See Kinase insert domain receptor (KDR) KE. See Kinetic energy (KE) KefauvereHarris amendments to FD&C Act, 1346 Keloids, 66, 70 Keratan sulfate, 618 Keratinocyte growth factor (KGF), 1300 Keratinocytes, 26e28, 66, 320, 1283e1285, 1297 differentiation, 23 Keratins, 1283 Keratocytes, 1115 Keratoprostheses (KPros), 1116 regenerative medicine applying to, 1116e1118 Kermanite (Ca2MgSi2O7), 706e707 b-Keto nitrile tautomeric copolymers, 704 KGF. See Keratinocyte growth factor (KGF) Kidney(s), 142e143, 1149e1150 disease, 1165 regenerative medicine approaches cell-based therapy, 1166e1172

cell-free approach, 1172e1173 tissue, 1169 kidney tissueederived stem and primary cells, 1166e1168 transplantation, 1165 Kinase insert domain receptor (KDR), 311 Kinetic energy (KE), 525 Kit-CreERT2/+ reporter mice, 250e251 Klarsicht, ANC-1, Syne homology domains (KASH domains), 459 KLF2. See Kru¨ppel-like factor 2 (KLF2) Klinefelter syndrome, 1252 Knee anterior cruciate ligament, 1184e1185 medial collateral ligament, 1184 motion, 1183 multiple ligamentous injuries in, 1185 Knockout serum replacement (KSR), 114 KPros. See Keratoprostheses (KPros) Krabbe disease, 158 Kru¨ppel-like factor 2 (KLF2), 169e170 Kruppel-like factor 4 (KLF4), 50, 52, 57 Kru¨ppel-like factor 5 (KLF5), 169e170 KSR. See Knockout serum replacement (KSR) Kyphoplasty, 607

L

L-type channel blocker, 173 Lab-on-a-chip style models, 825 Labile cells, 683e684 b-Lactamases, 898e899 Lactide, 568 L-Lactic acid (L-LA), 568 Lactobacillus acidophilus (L. acidophilus), 898e899 Lactones, polyesters of, 570 Lacunae system, 133 Lagrangian strain. See Green strain Lamin proteins, 394e395 Lamina propria, 1251e1252 Laminin, 279, 393, 617, 645e646, 1270 Laminin 521 or 511, 118 “Landing pad”, 748 Langerhans’ cells, 1283 LangmuireBlodgett deposition method, 655 Large animal lungs regeneration, bioreactors for, 790, 791f Large animal models, 765 Large-diameter tissue-engineered vascular graft fabricating tissue engineered vascular grafts, 1030e1038 biodegradable synthetic-based scaffolds, 1038 biological-based scaffolds, 1030e1033 hybrid scaffolds, 1033e1037 Large-scale tissue construction, vascularization in cell sheets for, 475e477 Laser ablation technique, 670 Laser patterning, 772 Laser therapy, 78 Laser-assisted bioprinting, 846

1403 Laser-assisted printing, 834 Laser-based bioprinting technology, 773 Laser-based patterning of hydrogels, 670 Laser-enabled analysis and processing, 95 Lateral line, 867, 876f hair cell regeneration using, 876e879 models of progenitor cells, 878f pathways coordinating hair cell regeneration in, 879e880 regeneration, 880 Latex protein, 689e690 Layer-by-layer method (LbL method), 655 LbL method. See Layer-by-layer method (LbL method) LCST. See Lower critical solution temperature (LCST) Lef-1. See Lymphoid enhancer-binding factor-1 (Lef-1) LEF/TCF. See Lymphoid enhancerbinding factor/T-cell factor (LEF/ TCF) Left ventricular (LV), 253 dysfunction, 253 function, 1087e1088 Left ventricular assist devices (LVADs), 688 Left ventricular ejection fraction (LVEF), 257 Lentivirus vectors, 749 Leptotrichia shahii (L. shahii), 745e746 Leukemia inhibitory factor (LIF), 49e52, 114, 119, 181e183 Leukocyte(s), 679 extravasation, 25 leukocyteeendothelial cell interactions, 679 migration, 445 Leukodystrophies, 158 Leukotriene B4 (LTB4), 677 Lheureux model, 44e45 LIF. See Leukemia inhibitory factor (LIF) Ligament, 959e960 healing of ligaments and tendons, 1183e1185, 1188e1192 mechanical properties of ligament substance, 1181 normal ligaments and tendons biology, 1180 biomechanics, 1180e1183 Ligand-receptor signaling, 6e7 growth factor-b pathway, 6 Hedgehog pathway, 7 notch pathway, 7 signaling by receptor tyrosine kinase ligands, 7 Wnt pathway, 6e7 Ligands, 442, 685 Light transmission aggregometry (LTA), 930 LIM-homeodomain transcription Nkx2e5, 247e248 Limb interaction of cells from opposite sides of limb circumference, 44e45 ischemia, 345

1404 Limb (Continued ) limb-girdle muscular dystrophy, 753 regeneration, 37 Limbal stem cell biopsy-derived stromal cells, 1119 Limbal stem transplantation, 1116 line. See Lineage negative (line) Line/Sca-1+/c-Kit+ cells (LSK cells), 194e195 Lin28, 50, 114 Lineage negative (line), 194e195, 253 Lineage reprogramming. See Transdifferentiation process LINK. See Linker of nucleoskeleton and cytoskeleton (LINK) 1,4-Linked b-D-mannuronic acid, 640e641 Linker of nucleoskeleton and cytoskeleton (LINK), 396, 459 Lipid(s), 209 proteins immobilization in lipid layers, 530 Lipofectamine, 497e498 15-Lipooxygenase, 888 Lipopolysaccharide (LPS), 186e187 Liposuctioned fat, 222 Liquid-to-powder ratio (LPR), 600 Live cells, 95 Live imaging, 279 tracking muscle stem cell behavior through, 278e279 Liver, 750e751, 1101 allografts, 239 biopsy, 239 cell therapy of liver disease background studies, 231e232 clinical hepatocyte transplantation, 232e237, 233t current treatments for liver disease, 230t hepatocyte transplantation, 237e241 cell transplantation, 1101e1102 ECMederived hydrogels, 779e780 failure, 321 liver-derived cell lines, 773 liver-on-a-chip, 773 liver-specific hepatospheres, 1106 markers, 779e780 organoids, 769, 773, 779e780 partial degradation pathway of tyrosine, 751f spheroids, 1106e1110 stem cells and alternative cell sources for liver therapy, 239e241 Liver transplantation (LT), 1101 clinical, 239 Living biological systems, 523 Living donor islet transplantation, 993e994 Living radical polymerization processes (LRP processes), 471, 479e480 LMW compounds. See Lowemolecular weight compounds (LMW compounds) Loadeelongation curve, 1181, 1181f Logistic(s), 1368, 1373

INDEX

Long-term HSCs (LT-HSCs), 107 Long-term survival, 981e982 Longitudinal outcome measures, 381 Lower critical solution temperature (LCST), 469e470, 563e564, 648 Lower urinary tract reconstruction, 1263 Lowemolecular weight compounds (LMW compounds), 689e690, 940 LOXL. See Lysyl oxidase-like proteins (LOXL) loxP site, 747e748 LPR. See Liquid-to-powder ratio (LPR) LPS. See Lipopolysaccharide (LPS) LRP processes. See Living radical polymerization processes (LRP processes) LSK cells. See Line/Sca-1+/c-Kit+ cells (LSK cells) LT. See Liver transplantation (LT) LT-HSCs. See Long-term HSCs (LT-HSCs) LTA. See Light transmission aggregometry (LTA) Lubrication, 413 Luciferase-expressing satellite cells, 278e279 Ludwig scale, 1305 Lumacaftor, 174e175 Lung, 143 bioreactors, 788e794 bioengineering functional lungs, 788e789 evaluation of bioengineered lungs, 793e794 for regeneration of large animal lungs, 790, 791f for regeneration of small animal lungs, 789e790, 789f for study of lung biology, 792e793 in vivo bioreactors for lung regeneration, 790e792, 792f cancer spheroids, 775 development, 1059e1060, 1060f lung-on-a-chip, 774 microvascular endothelial cells, 1065e1066 novel cell populations for lung repair, 1061e1063 regeneration, 1061, 1065, 1067f in vivo bioreactors for, 790e792, 792f transplantation, 1068 trapping, 324 tumor models, 775 LV. See Left ventricular (LV) LVADs. See Left ventricular assist devices (LVADs) LVEF. See Left ventricular ejection fraction (LVEF) Lymph nodeetargeting nanoparticles, 723e724 Lymphocytes, 680, 731, 1284 Lymphoid enhancer-binding factor-1 (Lef-1), 5 Lymphoid enhancer-binding factor/T-cell factor (LEF/TCF), 4 Lymphoid progenitor cells, 192

Lymphoid tissues, 724 Lyophilization, 620 Lyophilized scaffolds, 620 Lysine, 573, 640 Lysine-diisocyanate, 572e573 Lysyl oxidase-like proteins (LOXL), 5 Lysyl oxidases, 617

M

MACI. See Matrix-induced ACI (MACI) Macrophage mannose receptor (MMR), 682 Macrophage(s), 39, 106e107, 185, 506, 678, 680, 715, 862, 1200, 1228, 1284e1285 behavior, 516e517 in blastema formation, 41 depletion, 41 interactions, 680e681 Mac-1, 194e195 Macropores, 891e892 Macroporosity, 891e892 Macroporous alginate scaffold, 702 Macroporous Cryogel scaffolds, 729 Macroporous Cryogel-based cancer vaccine, 729 Macroporous scaffolds, 727, 731e733 Macroscale biomaterials, 731 implantable biomaterial scaffolds as cancer vaccines, 727e729, 728f to enhance autologous T cell therapy, 731e733 scaffolds for cancer immunotherapy, 727e733 injectable biomaterial systems as cancer vaccines, 729e731 Macroscale drug delivery biomaterial platforms, 727 Macular degeneration, 176 Mad. See Mothers against decapentaplegic (Mad) MadineDarby Canine Kidney cells (MDCK cells), 4 MafA. See Musculoaponeurotic fibrosarcoma oncogene homolog A (MafA) MAG. See Myelin-associated glycoproteins (MAG) Magnesium phosphate (MP), 708e709 Magnetic nanoparticles (MNPs), 491 Magnetic resonance imaging (MRI), 381, 491, 799, 831e832, 961 Maintenance, hESC, 117e118 Major histocompatibility complex (MHC), 135, 258, 323e324, 362, 685, 1011, 1258 MHC class II proteins, 688e689 MALBAC. See Multiple annealing and looping-based amplification cycles (MALBAC) Male reproductive system. See also Female reproductive system ejaculatory system, 1255e1256 penis, 1257e1258 testes, 1251e1254 Male urethra, 1255

INDEX

Maleylacetoacetate, 750e751 Mammalian auditory sensory epithelia, 867 Mammalian cardiogenesis, 247 Mammalian vestibular organs, spontaneous hair cell regeneration in, 870e871 Mammography Quality Standards Act program, 1347 Manufacturability, 517 MAP. See Mean arterial pressure (MAP) MAP2. See Microtubule-associated protein 2 (MAP2) MAPCs. See Multipotent adult progenitor cells (MAPCs) MAPK. See Mitogen-activated protein kinase (MAPK) MARCKS family, 41 Marfan syndrome, 212 Marine corals, 552e553 Marrow isolated adult lineage inducible cells (MIAMI), 182 Marrow stromal cells (MSC), 307, 578 Mass cytometry, 99 MASTERGRAFT, 897e898 Material surface property-dependent blood protein adsorption, 681 Material-forming processes, 695e696 Materials-based cancer immunotherapies advantages and disadvantages, 717e718 macroscale biomaterial scaffolds, 727e733 nanoparticle biomaterials, 718e727 overview, 715e717 MATES. See Multi-Agency Tissue Engineering Science (MATES) Mathematical identification of cellular subpopulations, 101e103 Matriderm, 1287e1288 Matrigel, 118, 668, 1090, 1104, 1205e1206 Matrix binding with growth factors, 1271 mechanics, 635e636 molecules, 16, 22 or matrix-mimicking bioinks, 810e821 cell-laden bioinks, 819e821, 820f co-printing and hybrid bioinks, 816e819, 817f natural materials, 813e816, 814f synthetic materials, 811e813, 812f Matrix metalloproteinases (MMPs), 2e3, 39, 579, 637, 666, 1200 MMP-2, 22, 232, 431e432 MMP-degradable hydrogels, 666 Matrix-bound vesicles (MBVs), 618e619 Matrix-induced ACI (MACI), 945e946 Mature cells, 664 Mature epidermis, 1283e1284 Mature stratum corneum, 1283 Maxillofacial reconstruction, current methods of, 890e891 Maximum capture rate (MCR), 1076e1077 MayereRokitanskyeKu¨stereHauser syndrome (MRKHS), 1237

MBG. See Mesoporous bioactive glass (MBG) MBVs. See Matrix-bound vesicles (MBVs) MC3T3 fibroblasts, 702 MC3T3-E1 osteoblastic cells, 439 MCAM. See Melanoma cell adhesion molecule (MCAM) MCAO. See Middle cerebral artery occlusion (MCAO) MCL. See Medial collateral ligament (MCL) MCPM. See Monocalcium phosphate monohydrate (MCPM) MCR. See Maximum capture rate (MCR) MCT. See Monocrotaline (MCT) MD. See Mean diffusivity (MD) MDCK cells. See MadineDarby Canine Kidney cells (MDCK cells) MDR. See Multidrug resistance (MDR) MDSCs. See Muscle-derived stem cells (MDSCs); Myeloid-derived suppressor cells (MDSCs) Mean arterial pressure (MAP), 376, 1153e1154 Mean diffusivity (MD), 382e383 Mechanical augmentation, 1190e1191 combined biological and, 1191e1192 Mechanical colony dispersion, 117 Mechanical deformations, 424 Mechanical elasticity and strength development, 1086 Mechanical failure, 505e506 Mechanical forces, 787 Mechanical picking, 117 Mechanical regulation of vascularized tissue regeneration, 427e432 mechanical stimulation in vitro, 429e432 blood vessel bioreactors, 431e432 bone bioreactors, 429e430 cartilage bioreactors, 429 mechanical stimulation in vivo, 427 Mechanical stimulation, 664, 1084e1085 Mechanical stimuli, 423e427 externally applied effect, 457e459 mechanical stress on cellular behavior, 458f mechanotransduction, 458e459 mechanotransduction, 424e427 tissue remodeling, 423e424 Mechanical stress(es), 636e637 Mechanical support, 505e506, 506f Mechanobiology, 391, 417, 1286 affects tissue development and function, 399e400 Mechanoreceptors, 393e394 Mechanosensation in FAK Y397 phosphorylation, 19e20 Mechanosensitive (MS), 393 ion channels, 425 Mechanosensory lateral line, 877 transduction, 392e393 Mechanotransduction, 397, 424e427, 458e459, 635e636

1405 through cellecell adhesions, 398e400 through celleextracellular matrix adhesions, 398 mechanisms and major effectors, 392e396 cell structure and composition, 392 cytoskeleton, 394e396 ECM, 392e393 ion channels and mechanoreceptors, 393e394 Medial collateral ligament (MCL), 1179, 1184 Medial collateral ligament, 1188e1189 of knee, 1184 Median normalization, 101 Medical Device Amendments to FD&C Act, 1346 Medical Research Council Adjuvant Gastric Infusional Chemotherapy trial, 261 Medical Technology Enterprise Consortium (MTEC), 1374 Medical therapy, 1131 Medicinal signaling cells, 344 MEF. See Mouse embryonic fibroblasts (MEF) Megakaryocytes (MK), 923 megakaryopoiesis, 929 and platelets from adult stem cells and somatic cells, 929 and platelets generating from human pluripotent stem cells, 930 Megakaryopoiesis, 929 MEK. See Mitogen activated protein kinase/Extracellular signalregulated kinase (MEK) MEK inhibitor PD0325901, 119 Melanocytes, 1283e1284 Melanoma cell adhesion molecule (MCAM), 206 MELD scores. See Model for End-Stage Liver Disease scores (MELD scores) Melt spinning technique, 553 Melt-derived approach, 553 Membranous labyrinth, 867 Memoranda of Understanding (MOU), 1361 Menaflex, 616 Meniscus, 961e962 Menstrual MSCs (MenSCs), 210 mEpiSCs. See Mouse epiblast stem cells (mEpiSCs) Merkel cells, 1283 mESCs. See Mouse embryonic stem cells (mESCs) Mesenchymal cells, 1 Mesenchymal cellular condensation, 696 Mesenchymal stem/stromal cells (MSCs), 82e83, 135, 182, 205e206, 240e241, 250e251, 258e261, 259f, 295, 315e316, 320e322, 335e336, 344, 371e373, 376, 383, 409, 441, 489, 540, 699e700, 762e763, 772, 799, 812, 823e824, 841e842, 853e854, 894, 909, 939, 953e954, 996, 1062, 1118, 1168e1169, 1187, 1202, 1251, 1264, 1284e1285, 1299, 1360

1406 Mesenchymal stem/stromal cells (MSCs) (Continued ) cell therapy with allogeneic, 262f exosomes, 208e211 for gene therapy, 322e325 identification, isolation, characterization, and in vitro expansion, 315e316 immunomodulatory effects, 211e212 induced pluripotent stem cellederived, 212e213 isolation techniques, 208 to modulating immunity, 344e345, 344f MSC-loaded collagen hydrogel collagen, 700e701 neurologic stem cell treatment study, 385 in regenerative medicine, 220e223, 221f asthma, 223 clinical trials, 224 clinically relevant therapies using MSCs, 222e223 diabetes, 223 inflammatory bowel disease, 223 mesengenic process, 220f multiple sclerosis, 223 new MSCs, 224 stroke and acute myocardial infarct, 223 stem cell nature, 206e207 subacute TIB trial, 383 rationale for using MSC, 383 results, 383 tissues containing, 207 transendocardial injection of autologous human cells, 260fe261f Mesenchymaleepithelial transition (MET), 1 Mesenchyme, 1136 Mesoangioblasts, 976 Mesoderm posterior 1 gene (Mesp1 gene), 247e248 Mesp1-rtTA transgene, 248 Mesp1+ progenitors, 248 Mesoporous bioactive glass (MBG), 709 Mesoporous silica nanobiomaterials, 492e493 Mesoporous silica rods (MSRs), 729e731, 731f Mesp1 gene. See Mesoderm posterior 1 gene (Mesp1 gene) Messenger ribonucleoprotein granules (mRNP granules), 275e276 Messenger RNA (mRNA), 50, 76, 98e99, 170, 275e276, 745e746, 928, 972e973, 1187 MET. See Mesenchymaleepithelial transition (MET) Metabolic activity, 1157 Metabolic cues, 281 Metabolic liver disease(s), 237 clinical transplants for, 235t hepatocyte transplantation for, 234e237 Metallic/metal(s), 831 biomaterials, 507e508

INDEX

metal-catalyzed oxidation, 567 metaleceramic blends, 709 metalepolymer blends, 709 nanobiomaterials, 492, 492t scaffolds, 509, 513, 707 Methacrylate, 579 Methacrylated hyaluronic acid, 808e809 Methacrylates, 559e560 2-Methacryloyloxyethyl phosphorylcholine (MPC), 1123 N-Methyl-D-aspartate (NMDA), 394 MHC. See Major histocompatibility complex (MHC) MI. See Myocardial infarction (MI) MIAMI. See Marrow isolated adult lineage inducible cells (MIAMI) Mibefradil, 780 Michael-type addition reactions, 638, 647e648 Microarray post detectors, 1081e1082 Microbial cellulose, 641e642 Microbiofabrication, 772 Microcomputed tomography (mCT), 699e700, 799 Microcontact printing, 659, 772 Microencapsulated ovarian cells, 1244 Microengineered/microengineering, 772e773 lung models, 792 tissue constructs, 770 Microfabricated intelligent surface for engineering complex tissue constructs copatterning to create cellular microenvironment, 478 regulating cell orientation in cell sheet engineering, 479e481 Microfabrication techniques, 95e96, 769e770, 772e775 Microfilaments, 395 Microfluidic organoid for drug screening platform (MODS platform), 1109 Microfluidic(s), 769, 773e774 cell culture devices, 1142e1143 chips, 457 circuit, 769e770 device, 776, 1243e1244 fluid exposure, 659 platforms, 454e455 systems, 773e775 techniques, 774e775 technologies, 95e96, 772 Microglia, 371e373, 1200 b2-Microglobulin, 930, 1151 Microgrooved polydimethylsiloxane substrates, 479e480 Microinjection delivery, 749 Microparticulate ceramic materials, 577 Micropatterning, 478, 772 approaches, 479e480 of PEG hydrogels, 670 technologies, 772 Microphotopatterning lines, 457 Microphysiological system, 1082 Microporous PLGA scaffolds, 670

MicroRNA (miRNA), 39e40, 50, 208, 860, 862f integrating with cell signaling and transcription factors, 57e58 miR-9, 5 Microspheres, 858, 1172 Microtubule-associated protein 2 (MAP2), 448 Microtubules (MTs), 394e395 Microvasculature, importance of, 667e668 Microvesicles (MVs), 205, 208 Middle cerebral artery occlusion (MCAO), 1204 Migration, 788 celleECM interactions during healing of cutaneous wounds, 25e26 during regenerative fetal wound healing, 29 signal transduction events during celleECM interactions, 19e22 Miniaturization approaches, 773 Miniorgan, 1304 Minocycline, 1206e1207 miR-17e92 miRNA cluster, 55 miR302 cluster, 57e58 miRNA. See MicroRNA (miRNA) Mitogen activated protein kinase/ Extracellular signal-regulated kinase (MEK), 49e50 Mitogen-activated protein kinase (MAPK), 6, 19e20, 49e50, 284, 398 Mitogenesis, 406 Mitogenic response, 408 Mixed chimerism, 198e199 Mixed metalloproteinase 8 (MMP 8), 72 MK. See Megakaryocytes (MK) MLP. See Muscle LIM protein (MLP) mMAPC-VP. See mMAPCs predifferentiated to vascular progenitors (mMAPC-VP) mMAPCs. See Murine MAPCs (mMAPCs) mMAPCs predifferentiated to vascular progenitors (mMAPC-VP), 186 MMP 8. See Mixed metalloproteinase 8 (MMP 8) MMPs. See Matrix metalloproteinases (MMPs) MMR. See Macrophage mannose receptor (MMR) MNCs. See Mononuclear cells (MNCs) mNPCs. See Murine NPCs (mNPCs) MNPs. See Magnetic nanoparticles (MNPs) MO. See Monocyte (MO) Mobilized PB (mPB), 309 MOD. See Multiple organ dysfunction (MOD) Model for End-Stage Liver Disease scores (MELD scores), 241 Modern surface analysis methods, 524e525 Modified cells, 381 Modified SB623 cells trial, 384

INDEX

MODS platform. See Microfluidic organoid for drug screening platform (MODS platform) MOF. See Multiple organ failure (MOF) Molecular cell biology, 405e406 Molecular circuitry underlying pluripotency, 49 BMP, 50e52 chromatin structure determining regulatory activity, 58 ground state and primed embryonic stem cells, 49e50 induced pluripotent stem cells, 50 LIF, 50e52 MicroRNAS integrating with cell signaling and transcription factors, 57e58 Myc linking cell signaling to pluripotency gene regulation, 54e55 OCT4, SOX2, NANOG, 53e54 specific epigenetic program helping maintain pluripotency, 55e57 TGFb and fibroblast growth factor signaling pathways, 52 WNT signaling, 52e53 Molecular imaging techniques, 788 Molecular imprinting systems, 510, 511f Molecular organization of cells, 1e4 changes in cell polarity and stimulation of cell motility, 3 changes in cellecell adhesion, 2 changes in celleextracellular matrix adhesion, 3 epithelial vs. mesenchymal, 2f epithelialemesenchymal transition transcriptional program, 4e5 invasion of basal lamina, 3e4 molecular control of epithelialemesenchymal transition, 5e8 Molecular self-assembly, 635 Molecular signaling factors, 280 Molecular weight (MW), 1149 Molecular weight cutoff (MWCO), 1149e1150 Monkey’s biceps brachium, 977e979 Monocalcium phosphate monohydrate (MCPM), 592e593 Monocrotaline (MCT), 319 Monocyte (MO), 185, 373, 678e679, 1155e1156 Mononuclear cells (MNCs), 208, 311 Mononuclear phagocyte/phagocytic system (MPS), 488, 490, 680, 681t, 720 Monophosphoryl lipid A (MPL-A), 722 Monotherapy, 80e81 “Mop-end” injury model, 1184 Morphogenesis, 405e406 of cartilage, 412 Morula, 115 Mosquito HTS, 95 Mothers against decapentaplegic (Mad), 426

MOU. See Memoranda of Understanding (MOU) Mouse 4T1 breast tumor resection model, 731e733 Mouse cytokeratin 18, 240e241 Mouse embryonic carcinoma cells, 113 Mouse embryonic fibroblasts (MEF), 117 Mouse embryonic stem cells (mESCs), 49e50, 113e114, 119, 181, 445, 451, 617 LIF and BMP signaling pathways regulating mESc self-renewal, 50e52 mouse ES-derived cells, 239e240 Mouse epiblast stem cells (mEpiSCs), 181 Mouse femoral bone marrow, 193 Mouse Flk-1, 311 Mouse MC3T3-E1 cells, 704e705 Mouse skin fibroblasts, 762 Mouse Vasa homolog (MVH), 1244e1245 MP. See Magnesium phosphate (MP) mPB. See Mobilized PB (mPB) MPC. See 2-Methacryloyloxyethyl phosphorylcholine (MPC) MPL-A. See Monophosphoryl lipid A (MPL-A) MPS. See Mononuclear phagocyte/ phagocytic system (MPS); Mucopolysaccharidosis (MPS) MRFs. See Myogenic regulatory factors (MRFs) MRI. See Magnetic resonance imaging (MRI) MRKHS. See MayereRokitanskyeKu¨stere Hauser syndrome (MRKHS) mRNA. See Messenger RNA (mRNA) mRNP granules. See Messenger ribonucleoprotein granules (mRNP granules) MRTFs. See Myocardin-related transcription factors (MRTFs) MS. See Mechanosensitive (MS); Multiple sclerosis (MS) MSC. See Marrow stromal cells (MSC) MSC-derived extracellular vesicles (MSC-EVs), 210 MSC-EVs. See MSC-derived extracellular vesicles (MSC-EVs) MSCs. See Mesenchymal stem/stromal cells (MSCs) MSCs derived from induced pluripotent stem cells (iMSCs), 206, 213 MSRs. See Mesoporous silica rods (MSRs) Msx1, 39e40 MTEC. See Medical Technology Enterprise Consortium (MTEC) mTR mouse, 284 MTs. See Microtubules (MTs) MTU. See Muscleetendon unit (MTU) Mucin, 70 Mucopolysaccharides, 639 Mucopolysaccharidosis (MPS), 158 Mucosa of neointestine, 1136e1137 Mu¨ller glia, 1200 Mu¨llerian ducts, 1237

1407 Multi-Agency Tissue Engineering Science (MATES), 1361 Multiarmed PEG, 580 Multicellular tissues, 1 Multicenter bone marrow mononuclear cell pediatric trial, 384 Multidisciplinary studies, 1188e1189 Multidrug resistance (MDR), 195 MDR-1, 261e262 MDR-2, 232 Multifunctionalized systems, 911 Multiorgan systems and future applications, 775e783 cutting-edge body-on-a-chip, 779e783 importance of multiorganoid integration, 776e779 Multiorganoid body-on-a-chip platform, 779, 783 integration importance, 776e779 additional disease modeling, 778e779 cancer, 776 drug testing and toxicology, 777e778 studies, 780 systems, 783 Multiphasic scaffolds, 634e635 Multiphoton-excited 3D printing method, 1080 Multiple annealing and looping-based amplification cycles (MALBAC), 97 Multiple commercial 3D coculture platforms, 1108e1109 Multiple disease mechanisms, 174e175 Multiple disorders of mesenchymal tissues, 212 Multiple displacement amplification, 97 Multiple ligamentous injuries in knee, 1185 Multiple organ dysfunction (MOD), 1150 Multiple organ failure (MOF), 1154 Multiple sclerosis (MS), 220, 223 experimental autoimmune encephalitis model, 223 Multipotent adult progenitor cells (MAPCs), 182, 371e373 ASCs, 181e182 differentiation potential of rMAPCs and hMAPCs in vitro, 183 future directions, 187 isolation of hMAPCs, 183 isolation of rMAPCs, 182e183 MAPC-induced tolerogenicity, 185 regenerative capacities, 183e187 hematopoietic reconstitution, 183 immunomodulatory properties of low-oct4 mMAPCs and hMAPCs, 184e187 stem cells, 181 Multipotent cardiovascular stem cell, 248e250 Multipotent progenitors, formation of new neuromasts from, 877 Multipotentiality, 1266e1267 of MSCs, 207 multiple modes of action assigned to adult stem cells, 1267t

1408 Multiprotein surface, 523e524 Multistep perfusion system, 1105 Multisystem organoids, 174 Multivariate methods for analyzing surface molecular information, 527 Multivariate statistical methods, 527 Multiwalled CNT-soaked collagen (MWCNT-soaked collagen), 451 Murine AFSC, 138 Murine cardiomyocyte-like tumor cells, 253 Murine embryonic stem cells, 664 Murine fibroblasts, 445 Murine MAPCs (mMAPCs), 182 immunomodulatory properties of lowoct4 in vitro effects of MAPCs on T cells, B cells, and NK cells, 184 in vivo immunodulatory effects, 184e186 Murine NPCs (mNPCs), 448 Musashi-1, 1135 Muscle, 751e753 cells, 1273 fibers, 279 interstitial cells, 273 precursor cells, 971 regeneration, 40, 281, 972 regenerative defect, 277 stem cell types within, 285e286 iPSederived muscle stem cells, 286 tissue, 494e495 Muscle LIM protein (MLP), 41 Muscle stem cells (MuSCs), 273, 279e281, 282f, 283e284 activation, 280 functional characteristics, 276e277 isolation, 277 molecular characteristics, 274e276 regulation, 279e281 satellite cells, 274 Muscle-derived stem cells (MDSCs), 963 Muscleetendon unit (MTU), 817, 845e846, 845f Muscleevein combined grafts, 1224e1225 MuSCs. See Muscle stem cells (MuSCs) Muscular dystrophy, 287, 751e753, 752t, 963 DMD, 752e753 limb-girdle muscular dystrophy, 753 Muscular layers, 1238 Muscular thin films, 1080 Muscularis mucosa, 1135 Muscularis propria, 1135 Musculoaponeurotic fibrosarcoma oncogene homolog A (MafA), 336e337 Musculoskeletal diseases, 953 Musculoskeletal system, 141, 953 Mussel-inspired surface modification strategies, 655 MVH. See Mouse Vasa homolog (MVH) MVs. See Microvesicles (MVs) MW. See Molecular weight (MW)

INDEX

MWCNT-soaked collagen. See Multiwalled CNT-soaked collagen (MWCNT-soaked collagen) MWCO. See Molecular weight cutoff (MWCO) Myc, 5, 170 linking cell signaling to pluripotency gene regulation, 54e55 Myelin-associated glycoproteins (MAG), 1200e1201 Myeloblasts, 775 Myeloid-derived suppressor cells (MDSCs), 715 Myelosuppressive effects of chemotherapy, 196 Myf5 gene, 275e276 Myoblasts, 261, 263f, 273, 480e481, 971 culture, 440 progeny, 274 sheets, 473e474 transplantation in skeletal muscles cell administration, 976e980 cell-graft survival, 980e982 relevant properties of SCDMs, 972e975 SCDMs, 971e975 Myocardial infarct, 186 Myocardial infarction (MI), 141e142, 175e176, 186, 252e253, 495, 1073 Myocardial tissue engineering, 636 Myocardin-related transcription factors (MRTFs), 5 Myocardium, 489e490, 1094 regeneration, 473e474 Myocyte renewal, 253 MYOD+ cells, 282e283 MYOD1, 276 Myofiber(s) basal lamina, 279 formation of new, 973e974 fragment, 39 Myofibroblast(s), 1, 30, 680 differentiation, 23e24, 27 Myogenesis, 971e972 molecular characteristics of muscle stem cells, 274e276, 275f Myogenic cells, 285e286 Myogenic regulatory factors (MRFs), 275 Myogenic transcription factors, 286 Myoglobin, 1151 Myometrium cells, 1240 Myopathy, 972 Myoseverin, 40 Myosins, 276, 299 Myostatin, 280 Myriad Genetics, 1325

N

Na+-glucose cotransporter 1 (SGLT1), 1138 NAAGS. See N-Acetyl aspartyl-glutamate synthetase (NAAGS) NAD+. See Nicotinamide adenine dinucleotide (NAD+) NaHCO3. See Sodium bicarbonate (NaHCO3)

Naive embryonic stem cells, 119 “Naive” state, 181 NAM Forum. See National Academy of Medicine Forum (NAM Forum) Nano-HA (n-HA), 704e705 Nanoapatite (nAp), 708 Nanobiomaterials, 491e493, 492t Nanoengineering techniques, 1288 Nanofibrous PLLA (NF-PLLA), 857 Nanofibrous scaffolds, 496 for bone tissue engineering, 857 nanofibrous silk scaffolds, 548 Nanofibrous spongy microsphere (NF-SMS), 858 Nanog, 50, 52e54, 57, 114, 138 Nanoimprinting, 447t, 489 Nanolipogels (NLGs), 722, 723f Nanomaterial properties, 486e491 chemical properties, 490 self-assembly, 490 surface chemistry, 490 electrical properties, 491 magnetic properties, 491 optical properties, 490e491 physical properties, 486e490 shape, 488e489 size, 487e488 surface topography, 489e490 physicochemical properties, 488f Nanomaterial-based drug delivery, 495e496 Nanomedicine applications in immunotherapy, 727 in cancer, 718e719 nanoparticle platforms, 719f Nanoparticle(s), 718e719 biomaterials for cancer immunotherapy, 718e727 effects of nanoparticle size and shape, 720 effects of nanoparticle surface charge and hydrophilicity, 720 effects of nanoparticle surface functionalization, 720e721 nanomedicine applications in immunotherapy, 727 nanomedicine in cancer, 718e719 delivery systems, 722 effects of nanoparticle size and shape, 720 effects of nanoparticle surface charge and hydrophilicity, 720 nanoparticle vaccine strategy, 724 nanoparticle-based drug delivery, 494 surface functionalization effects, 720e721 systems, 721 targeting amphiphilic peptide and adjuvants targeting lymph nodes, 725f of APCs, 722e727 applications in immunotherapy, 721 of tumor microenvironment, 721e722 Nanoscale spinning process, 1151 Nanospheres, 208

INDEX

Nanostructured biomaterials, 491e492 Nanostructured CaP biomaterials and scaffolds, 706 Nanostructured scaffolds, 496 Nanotechnology, 485, 718e719 applications for stem cell therapy, 497f nanotechnology-based stem cell therapy, 497e498 nanotechnology applications for, 497f stem cell delivery, 498 stem cell expansion, 498 stem cell transfection, 497e498 nanotechnology-based strategies in regenerative medicine bone tissue, 493e494 cartilage, bladder, and skin, 496 muscle tissue, 494e495 neural tissue, 496 vascular tissue, 495e496 Nanotopography, 448 nAp. See Nanoapatite (nAp) NAS. See US National Academies of Science (NAS) National Academy of Medicine Forum (NAM Forum), 1362 National cord blood inventory (NCBI), 150 National Heart, Lung, and Blood Institute’s Cardiovascular Cell Therapy Research Network, 265, 1361 National Institute for Innovation in Manufacturing Biopharmaceuticals (NIIMBL), 1374 National Institute of Neurological Disorders and Stroke, 1361 National Institute of Standards and Technology (NIST), 1355, 1368 National Institutes of Health (NIH), 126, 126b, 169, 440, 780e782, 817, 1311e1312, 1333e1334, 1353 NIH 3T3 fibroblasts, 440, 773 NIH-negotiated licenses, 1325 National Marrow Donor Program (NMDP), 150 National Nanotechnology Initiative, 718e719 National Network for Manufacturing Innovation, 1374 National Research Council, 1331e1332 Native intestinal tract, 1136e1137 Native starch, 549 Natural biopolymers as extracellular matrixeanalog hydrogels, 639e646 polysaccharides, 639e642 proteins and peptides, 642e646 Natural collagen matrix, 1270e1271 acellular tissue matrices, 1270 collagen, 1270 matrix binding with growth factors, 1271 silk, 1271 Natural decellularized matrices allogenic matrices, 1032 xenogenic matrices, 1030e1032

Natural killer cells (NK cells), 184, 196, 211 , 698, 715e716, 1013 in vitro effects of MAPCs on, 184 Natural materials, 813e816, 814f alginate, 1048 chitosan, 1049 fibrin, 1047e1048 gelatin, 1048 HA, 1048 PHAs, 1049 for TEHVs, 1047e1049 Natural nerve grafts, 1231e1232 Natural polymers, 699e704, 892. See also Synthetic polymers alginate, 702 CHI, 703 collagen, 700e701 HAc, 701e702 peptide hydrogels, 703e704 silk, 699e700 Natural scaffolds, 1205e1206, 1224e1225 alginates, 940 chitosan, 940e941 collagen, 940 HA, 940 Natural-based bioceramics, 550 Natural-based polymers, 538 Naturally derived hydrogels, 836e837 Naturally derived scaffolds, 763 Nature-derived polymers, 1032, 1169 Nature-derived polymers and synthetic polymers, 1034e1035 NC. See Neural crest (NC) NCAM. See Neural cell adhesion molecule (NCAM) NCAs. See N-carboxy-anhydrides (NCAs) NCBI. See National cord blood inventory (NCBI) NCM. See Normoxic conditions (NCM) NDA. See New Drug Application (NDA) NDD. See Neurological determination of death (NDD) NDMA receptor-mediated neuronal death, 1203 NE. See Nuclear envelope (NE) NEBs. See Neuroepithelial bodies (NEBs) NEC. See Necrotizing enterocolitis (NEC) Necrosis, 980 Necrotizing enterocolitis (NEC), 143 Needlestack platform, 104 Negative-pressure wound therapy, 79 Neointestine, 1137 Neomucosa, 1136 Neomycin, 870e871 Neonatal donor livers, 229e230 Neonatal HIE, 159 Neonatal murine-only skin substitutes, 1304 Neonatal rat cardiac cells, 1085 Neonatal rat cardiomyocytes, 1076 Neonatal rat heart cells, 1083e1084 Neovascularization, 68, 107, 209e210, 475e477, 680 Nephila clavipes (N. clavipes), 645 Nephron-like segments, 1171

1409 Nephropathy, 345 Nerve, 754 autografts, 1224, 1231e1232 fibers, 1140e1141 hydrogel scaffold for nerve repair, 1226 injury, 1227e1228 synthetic scaffolds for nerve repair, 1225e1226 tissue, 496 Nerve growth factor (NGF), 667, 1201, 1226e1227 Nerve guidance channels (NGCs), 1223e1224 Nerve regeneration anisotropic scaffolds, 1229e1231 isotropic scaffolds, 1224e1229 ECM molecules, 1226e1227 electroconductive scaffolds, 1229 hydrogel scaffold for nerve repair, 1226 natural scaffolds, 1224e1225 neurotrophic factors and cytokine delivery, 1227e1228 seeding neuronal support cells, 1228e1229 synthetic scaffolds for nerve repair, 1225e1226 Nervous system, 141, 172e173 NES. See Nuclear export sequence (NES) Nesprins, 459 Neural cell adhesion molecule (NCAM), 276e277 Neural crest (NC), 1119 neural crest cellederived dental mesenchyme, 908e909 neural crestespecific myelin P0 Crereporter mouse line, 250e251 Neural ectoderm, 1283 Neural progenitor cells (NPCs), 362, 448, 1205e1206 Neural retina, 1208e1210 Neural stem cells (NSCs), 103, 181e182, 376, 448, 1200e1201, 1205e1206, 1300, 1360 Neural stem-progenitor cells (NSPCs), 1202 Neural tissue, 496, 847 NeuroD1. See Neurogenic differentiation factor 1 (NeuroD1) Neurodevelopmental disorder, 173 Neuroectoderm, 399 Neuroepithelial bodies (NEBs), 1061 Neurogenesis, 378 Neurogenic differentiation factor 1 (NeuroD1), 336e337, 1205 Neurogenin 3 (NGN3), 336e337, 339e340 Neuroinflammation, 371e373, 372f multipotent adult progenitor cell treatment, 374f Neurological determination of death (NDD), 993 Neurological diseases, 172e173, 496 Neuromasts, 876e877 formation from multipotent progenitors, 877

1410 Neuronal apoptosis, 371 Neuronal lineage, 1208 Neurons, 496, 1199 Neuroprotection, 361, 1201 Neuroregeneration, 1201 Neurospheres, 265 Neurotrophic factors, 184, 351, 1227e1228, 1230 Neurotrophin-3 (NT-3), 1201, 1229 NeutrAvidin Protein A complexes, 531 Neutrophils, 25, 72, 106, 373, 676e679 New Drug Application (NDA), 1348e1349 New Jersey Stem Cell Research Grant Program, 1314 New York Blood Center, 149 Newt regeneration blastema, 39e40 Next-generation DNA sequencing, 97 NF-AT. See Nuclear factor of activated T cells (NF-AT) NF-PLLA. See Nanofibrous PLLA (NF-PLLA) NF-SMS. See Nanofibrous spongy microsphere (NF-SMS) NFAT. See Nuclear translocation of nuclear factor (NFAT) NGCs. See Nerve guidance channels (NGCs) NGF. See Nerve growth factor (NGF) NGN3. See Neurogenin 3 (NGN3) NHEJ. See Nonhomologous end joining (NHEJ) Ni-NTA. See Nickelenitrilotriacetic complex (Ni-NTA) NICD. See Notch intracellular domain (NICD) Niche, regulation of muscle stem cells by, 279e281 biophysical cues, 280e281 extracellular matrix components, 279e280 metabolic cues, 281 molecular signaling factors, 280 Nickases, 743, 743f Nickelenitrilotriacetic complex (Ni-NTA), 528, 528f “Nicks”, 743 NiCord. See Nicotinamide-based expansion approach (NiCord) Nicotinamide adenine dinucleotide (NAD+), 281 Nicotinamide-based expansion approach (NiCord), 157 NIH. See National Institutes of Health (NIH) NIIMBL. See National Institute for Innovation in Manufacturing Biopharmaceuticals (NIIMBL) Niobium-coated carbon disks, 1156e1157 NiPAAm. See N-Isopropylacrylamide (NiPAAm) NIST. See National Institute of Standards and Technology (NIST) Nitric oxide (NO), 310, 426, 880 Nitric oxide synthase (NOS), 426 NK cells. See Natural killer cells (NK cells)

INDEX

NK2 homeobox 1 transcription factor (Nkx2.1), 1063 NK2 transcription related, locus 5 (Nkx2e5), 247e248 NK6 homeobox 1 (NKX6.1), 338e339 Nkx2.1. See NK2 homeobox 1 transcription factor (Nkx2.1) Nkx2e5. See NK2 transcription related, locus 5 (Nkx2e5) NKX6. 1. See NK6 homeobox 1 (NKX6. 1) NLGs. See Nanolipogels (NLGs) NMDA. See N-Methyl-D-aspartate (NMDA) NMDP. See National Marrow Donor Program (NMDP) NO. See Nitric oxide (NO) NOD. See Nonobese diabetic (NOD) NogoA, 1201 Noise reduction in single-cell data, 100e101 Nonchemically modified ECM scaffold materials, 620 Nonchimeric animals, 1013 Noncoding RNAs, 5, 57 Noncovalent bonds, 635 Noncovalent coatings, 655 Noncultured dissociated murine skin cells, 1304 Nondegradable polymers, 559 Nondegradable synthetic polymers, 561e567 polymers witheCeCebackbone, 561e567 hydrolytically stable PUs, 566e567 other nondegradable polymers, 566e567 poly(ethylene terephthalate), 564f, 566 poly(meth)acrylates and polyacrylamides, 562e564 polyethers, 564e565 polyethylene and derivatives, 561e562 polysiloxanes, 565 Nondestructive quality control systems, 1372 Nonfibrillar components, 1180 Nonfibrous proteins, 437 Nonhealing wounds, 72 Nonhematopoietic stem cells, 257, 307 Nonhomologous end joining (NHEJ), 742e743, 745 Nonhuman primates, 765e766, 971 Nonlinear viscoelastic models, 1181e1182 Nonmalignant hematological diseases, CBT for, 156e157 Nonneuronal stem cells, 1202 Nonobese diabetic (NOD), 999 Noneorgan transplant candidates, hepatocyte transplants for, 237e238 Nonepatient derived immunotherapies, 717 Nonpeptide amino acidebased polymers, 573 Nonperfused bioreactors for bone regeneration, 795

Nonresorbable materials, 1136 Nonseeded scaffold bladders, 1274 Nonsolid cancer survivors, 1253 Nonspecific chemical reactions, 653 Nonsteroidal antiinflammatory drugs, 780, 888, 947 Nonsulfated glycosaminoglycan, 1048 Nonsurgical vaginal dilation, 1238 Nonsynaptic layer, 353e354 Nontraumatic peripheral nerve injuries, 1223 Nonvascular autogenous bone grafts, 890e891 Normal cells in human tissues, 697 Normal ligaments and tendons, 1180e1183 Normal rat kidney (NRK), 440 Normal stress, 421 Normalization, single-cell data, 101 Normoxic conditions (NCM), 1299 Northern white rhinoceros (NWR), 176 fibroblasts, 176 Norwood Hamilton scale, 1305 NOS. See Nitric oxide synthase (NOS) Notch, 39e40 Notch1, 7 pathway genes, 7, 874 signaling pathway, 872 Notch intracellular domain (NICD), 872 Novel cell delivery routes, 380e381 Novel cell populations for lung repair, 1061e1063 Novel CRISPR system, 745 Novel decellularized ECM bioink, 1079 Novel three-component biomimetic hydrogel, 704e705 NP. See Nucleus pulposus (NP) NPCs. See Neural progenitor cells (NPCs) Nrad expression, 39e40 NRK. See Normal rat kidney (NRK) NSCs. See Neural stem cells (NSCs) NSPCs. See Neural stem-progenitor cells (NSPCs) NT-3. See Neurotrophin-3 (NT-3) Nuclear envelope (NE), 396 Nuclear export sequence (NES), 5 Nuclear factor kB (NF-kB), 5, 378, 426 Nuclear factor of activated T cells (NF-AT), 427 Nuclear lamina, 396e397 Nuclear magnetic resonance spectroscopy, 531 Nuclear translocation of nuclear factor (NFAT), 688 Nuclease-deficient Cas9 (dCas9), 754 Nucleic acids, 859e860 Nucleophilic additions reactions, 638 Nucleotide delivery, 860e861 Nucleus as central organelle in regulating mechanotransduction, 396e397 Nucleus pulposus (NP), 299, 962 NutriStem, 118 NuvaRing, 562 NVP. See N-Vinyl pyrrolidone (NVP)

1411

INDEX

NWR. See Northern white rhinoceros (NWR) Nylon, 559e561

O

OA. See Osteoarthritis (OA) Occludin tight junction gene, 4 OCN. See Osteocalcin (OCN) OCP. See Octacalcium phosphate (OCP) OCPs. See Office of Combination Products (OCPs) OCT4. See Octamer binding protein 4 (OCT4) Octacalcium phosphate (OCP), 592e593 Octamer binding protein 4 (OCT4), 50, 52e54, 114, 138 ODN. See Oligodeoxynucleotides (ODN) Odontoblasts, 908e909 OEG. See Oligo(ethylene glycol) (OEG) Off-target effects, 745 “Off-the shelf” cancer drug, 717, 1289 Office for Human Research Protections, 1310 Office of Combination Products (OCPs), 1348 OGP. See Osteogenic growth peptide (OGP) Oligo(ethylene glycol) (OEG), 443 Oligo(poly[ethylene glycol] fumarat) (OPF), 578 Oligodendrocyte, 378 Oligodendrocyte cell line (OLN93), 184 Oligodendrocyte progenitor cells (OPCs), 1202 Oligodeoxynucleotides (ODN), 1187 Oligomeric biodegradable domains, 579 Oligonucleotide exon-skipping strategy, 752 Oligosaccharide(s), 641 residues, 297 Ologen Biocornea, 1122e1123, 1122f OLT. See Orthotopic liver transplantation (OLT) OM. See Osteogenic media (OM) Omentum, 997e998, 1136 On-a-chip technologies, 775e776 tissue models, 1109 On-demand release, 510e511 Oncogenes, 253 One-dimension (1D) Green strain, 420e421 migration, 457 Ontogeny, 1014 Oogonial stem cells (OSCs). See Germline stem cells (GSCs) OPCs. See Oligodendrocyte progenitor cells (OPCs) OPF. See Oligo(poly[ethylene glycol] fumarat) (OPF) OPN. See Osteopontin (OPN) Opsonins, 679 Optic nerve, 352 Optical imaging, 799 Optical techniques, 801

Optimal transplantation site, 997e998 Organ bath studies, 1257 Organ homeostasis, 316 Organ printing technique, 301 Organ Procurement and Transplantation Network, 1101 Organ replacement therapies, 927e928 Organ-on-a-chip engineering, 1080e1082, 1082f systems for personalized precision medicine, 782e783, 782f technologies and applications, 772e775 cancer-on-a-chip, 774e775 heart-on-a-chip, 774 liver-on-a-chip, 773 lung-on-a-chip, 774 microengineering and biofabrication, 772e773 vessel-on-a-chip, 773e774 Organ-specific stimuli, 621e622 Organic acids, 599 Organic additives, 600 Organic molecules, 593 Organismically dead embryos, 127e128 human embryonic stem cell lines, 128 irreversibility as criterion, 127e128 Organogenesis, 665 Organoid(s), 174, 769e770, 1106e1110, 1135 in drug development, 1109e1110 system, 1253e1254 Orientation control, 531 Ornithine transcarbamylase deficiency (OTC), 230, 236 Orphan diseases, 954 ORS. See Outer root sheath (ORS) Orthoester, 571 Orthopedic applications of CPCs, 607 vertebroplasty and kyphoplasty, 607 Orthopedic fixation devices, 567 Orthotopic liver transplantation (OLT), 229e230 OS. See Overall survival (OS) Osseointegration, 494, 907 Osteo-induced iPSCs, 762 Osteoarthritis (OA), 391, 413, 937, 958, 961 Osteoblast differentiation, 664 Osteoblastic niche, 193 Osteoblasts, 299, 764 Osteocalcin (OCN), 699e700, 909e910 Osteochondral autografting or mosaicplasty, 937e938 defects, 211 progenitors, 219 tissue, 958e959, 959f Osteoclasts, 39 Osteogenesis, 426, 706 Osteogenic cells, 762 Osteogenic differentiation markers, 855t of MSCs, 854 Osteogenic growth peptide (OGP), 444 Osteogenic media (OM), 703 Osteoinductive materials, 698 Osteomyelitis, 955

Osteonecrosis, 955 Osteoodonto-keratoprosthesis, 1116 Osteopontin (OPN), 529, 699e700, 811 Osteoporosis, 391, 591, 860, 955 Osteoprogenitor cells, 853 Osteopromotion, 888e889 OTC. See Ornithine transcarbamylase deficiency (OTC) Outer root sheath (ORS), 1297 “Outside-in” signaling, 19e20 Ovalbumin (OVA), 223, 723e724 Ovarian cells, 1243e1244 Ovarian cortical tissue, 1243 Ovarian follicle, 1242 Ovarian function, pathological loss of, 1242e1243 Ovarian hormones, 1242 Ovariectomy, 113 Ovaries, 1242e1245 regenerating ovarian tissue from stem cells, 1244e1245 tissue engineered ovarian follicles, 1243e1244 Overall survival (OS), 154 Overcoating alterations, 651, 652f technologies, 655e656 covalent coatings, 656 noncovalent coatings, 655 Ovine fetuses, 210 Oxygen supply, 1083e1084 transport, 1083 Oxygenation, 770 Oxytocin, 284 Ozurdex, 1210e1211

P

P-collagen. See Polymerized collagen (Pcollagen) P(EG-co-LA) diacrylate, 579 P(HEMA/MMA). See Poly (2-hydroxyethyl methacrylate-comethyl methacrylate) (P(HEMA/ MMA)) P[PF-co-EG]. See Poly(propylene fumarate-co-ethylene glycol) (P[PFco-EG]) p120-catenin, 2 P4HB. See Poly-4-hydroxybutyrate (P4HB) p63. See Protein 63 (p63) PA. See Polyanhydride (PA) PAcMo. See Poly(N-acryloylmorpholine) (PAcMo) PAD. See Peripheral artery diseases (PAD) PAH. See Pulmonary arterial hypertension (PAH) PAM. See Partitioning around medoids (PAM) PAM sequence. See Protospacer-adjacent motif sequence (PAM sequence) PAN. See Polyacrylonitrile (PAN) PAN/PVC. See Polyacrylonitrile/ polyvinylchloride (PAN/PVC) Pancreas, 341e343

1412 Pancreatic and duodenal homeobox (Pdx-1), 336e337 Pancreatic endoderm cells (PECs), 996 Pancreatic islet b cells, 335 Pancreatic progenitors, 338e339 Papillary dermis, 66 Papio ursinus (P. ursinus), 707 Paracrine effects, 1267e1268 Parallel-plate design, 429 flow experiments, 429e430 Paramesonephric ducts. See Mu¨llerian ducts Parathyroid hormone (PTH), 763e764, 860, 861f Parkinson disease (PD), 174, 176 Parthenogenesis, 115e116 Partial hepatectomy, 238e239 Partial hepatic resection, 238e239 Particulate leaching, 553 Partition clustering, 102 Partitioning around medoids (PAM), 102 Parylene C, 1139e1140 PAs. See Peptide amphiphiles (PAs) Passive physisorption of biomacromolecules, 657 Patellar tendon (PT), 1179 healing with ECM, 1188e1189 Pathogens, 715 Pathological loss of ovarian function, 1242e1243 Pathological scarring, 66 Patient assessment and selection, 990e991 Patient iPSC-derived sensory neuron, 175 Patient-derived cells, 717 Patient-specific technology, 899 Pattern recognition receptor (PRR), 727 Patterned substrates, 659 Pax7 expression, 277 PAX7ecells, 282e283 Paxillin, phosphorylation of, 20 PB. See Peripheral blood (PB) PBCA. See Poly(n-butylcyanoacrylate) (PBCA) PBGs. See Peribiliary glands (PBGs) PBMA. See Poly(n-butyl methacrylate) (PBMA) PBMCs. See Peripheral blood mononuclear cells (PBMCs) PBMNCs. See Peripheral blood mononuclear cells (PBMNCs) PBS. See Phosphate-buffered saline (PBS) PBSCs. See Peripheral blood stem cells (PBSCs) PBT. See Polybutylene terephthalate (PBT) PCA. See Principal components analysis (PCA) PCK-26epositive esophagus epithelial cells, 1133 PCL. See Poly(ε-caprolactone) (PCL) PCLA. See Poly(ε-caprolactone-co-lactide) (PCLA) PCP pathway. See Planar cell polarity pathway (PCP pathway) PCR. See Polymerase chain reaction (PCR)

INDEX

PCs. See Pluripotent cells (PCs) PCTMC. See Poly(glycolide-cotrimethylene carbonate) (PCTMC) PD. See Parkinson disease (PD); Peritoneal dialysis (PD); Population doubling (PD) PD translation. See Proximodistal translation (PD translation) PD-1. See Programmed cell death protein-1 (PD-1) PDCD6IP, 209 PDGF. See Platelet-derived growth factor (PDGF) PDL. See Periodontal ligament (PDL) PDL progenitor (PDLP), 914e915 PDL stem cells (PDLSCs), 909 PDLCL. See Poly(D,L-lactide) and ε-caprolactone (PDLCL) PDLP. See PDL progenitor (PDLP) PDLSCs. See PDL stem cells (PDLSCs); Periodontal ligament stem cells (PDLSCs) PDMS. See Polydimethylsiloxane (PDMS) PDS. See Pyridyl disulfide (PDS) PDX expression. See Podocalyxin expression (PDX expression) Pdx-1. See Pancreatic and duodenal homeobox (Pdx-1) PE. See Poly(ethylene) (PE) PEA-g-TA. See Phenyl amino end-capped tetraaniline (PEA-g-TA) PEC. See Polyelectrolyte complexation (PEC) PECAM-1. See Platelet-endothelial cell adhesion molecule-1 (PECAM-1) PECs. See Pancreatic endoderm cells (PECs) Pediatric Intensity Level of Therapy (PILOT), 381, 382f PEEK. See Poly(ether ether ketone) (PEEK) PEG. See Poly(ethylene glycol) (PEG) PEG-ADA. See Polyethylene glycolconjugated adenosine deaminase (PEG-ADA) PEG-dimethacrylate hydrogels, 580 PEGDA. See Poly(ethylene glycol) diacrylate (PEGDA) PEGylation, 647, 720e721 Pelnec, 1287e1288 Pelvic organ prolapse (POP), 1245e1246 Penetrating keratoplasty (PKP), 1115, 1122e1123 Penile vibratory stimulation (PVS), 1255e1256 Penis penile reconstruction, 1257 penile transplantation, 1257e1258 stem cell therapy for ED, 1258 Penumbra, 1203 PEO. See Polyethylene oxide (PEO); Proepicardial organ (PEO) Pepsin, 620e621 Peptide amphiphiles (PAs), 496, 703e704, 1123 Peptide(s), 24, 573e574

analogs of extracellular matrix, 1123e1124 hydrogels, 703e704 peptide-amphiphile nanofibers, 574 Perfluorocarbon emulsion (PFC emulsion), 1083e1084 Perfusion bioreactor, 1142e1143 for bone regeneration, 795e797, 796fe797f Perianal fistulas, 1142 Peribiliary glands (PBGs), 341, 342f Pericyte(s), 207, 224, 285e286, 295, 315, 1200 differentiation, 27e28 progenitors, 320 Periderm cells, 1283 Periocular mesenchymal precursor (POMP), 1119 Periodontal diseases, 907 Periodontal ligament (PDL), 474, 908e909 regeneration, 474 Periodontal ligament stem cells (PDLSCs), 909e910 Periodontal regeneration, 914e915 Periodontium, 888 Peripheral artery diseases (PAD), 1029 Peripheral blood (PB), 307, 924e925 types and source of stem cells in, 307e310 mobilization of bone marrow cells, 308e310 Peripheral blood mononuclear cells (PBMCs), 257 Peripheral blood mononuclear cells (PBMNCs), 1268 Peripheral blood stem cells (PBSCs), 196 cell surface markers expressed on progenitor and mature endothelial cells, 313t EPC, 311e315 future directions, 325e326 MSCs, 315e316 therapeutic applications, 317e325 EPC, 317e320 MSCs, 320e322 MSCs for gene therapy, 322e325 types and source of stem cells in PB, 307e310 Peripheral hind limb ischemia, 186 Peripheral nerve regeneration animal models, 1232 anisotropic scaffolds for nerve regeneration, 1229e1231 current strategies for, 1224, 1225f historical background, 1223e1224 isotropic scaffolds for nerve regeneration, 1224e1229 natural nerve grafts, 1231e1232 problems and challenges with peripheral nerve injuries, 1223 Peripheral nervous system (PNS), 496, 1223 injuries, 1223 Peripheral waveform analysis, 1159

INDEX

Perisinusoidal penile progenitor cells, 1258 Peritoneal dialysate, 1158 Peritoneal dialysis (PD), 1149 Perlecan, 618 Permeability, 701 Persistent vegetative state (PVS), 383 PET. See Positron-emission tomography (PET) PEVAc. See Poly(ethylene-co-vinyl acetate) (PEVAc) Peyer patches, 1135 PF. See bis(2-Hydroxypropyl) fumarate (PF) PF-diacrylate (PF-DA), 576e577 PFC emulsion. See Perfluorocarbon emulsion (PFC emulsion) PFIC. See Progressive familial intrahepatic cholestasis (PFIC) PGA. See Polyglycolic acid (PGA) PGC. See Primordial germ cells (PGC) PGCLCs. See Primordial germ cell-like cells (PGCLCs) PGD. See Preimplantation genetic diagnosis (PGD) PGE2. See Prostaglandin E2 (PGE2) PGS. See Poly(glycerol sebacate) (PGS) pH-sensitive chemistry, 724 pH-sensitive materials, 454 Phagocytosis, 679, 721 Pharmaceutical Affairs Act. See Pharmaceutical and Medical Device Act Pharmaceutical and Medical Device Act, 1375 Pharmacological immunosuppression, 982 Pharmacological therapy, stroke, 1203 PHAs. See Poly(hydroxy-alkanoates) (PHAs); Polyhydroxyalkanoates (PHAs) Phase separation, 633e634 PHB. See Poly 3-hydroxybutyrate (PHB) PHB-HV. See Polyhydroxybutyratehydroxyvalerate (PHB-HV) PHBHHx. See Poly 3-hydroxybutyrate-co3-hydroxyhexanoate (PHBHHx) PHBV. See Poly 3-hydroxyvalerate (PHBV) PhelaneMcDermid syndrome, 173 PHEMA. See Poly(2-hydroxyethyl methacrylate) (PHEMA) Phenyl amino end-capped tetraaniline (PEA-g-TA), 704e705 Phenylketonuria (PKU), 237 Phenylmethylsulfonyl fluoride, 1030 phiC31 phage integrases, 748 PHO. See Polyhydroxyoctanoate (PHO) Phosphate-buffered saline (PBS), 143, 211 Phosphatidylinositol 3’kinase (PI3K), 3, 393e394 Phosphatidylinositol bisphosphate (PIP2), 38e39 Photocross-linkable GelMA, 836e837 Photocross-linkable hydrogels, 633 Photocross-linking extrusion, 818

Photocured HAc hydrogel, 702 Photoelectrons, 525 Photogelation, 579e580 Photoinitiation, 669 Photolithography, 446, 447t, 633, 653e655 Photopatterned PEG hydrogels, 670 Photopatterning, 772 Photopolymerizable polyanhydrides, 575 Photopolymerization techniques, 580, 670 Photopolymerized (meth)acrylated biodegradable hydrogels, 579 Photopolymerized methacrylated polymer networks, 579e580 Photoreceptor layer, 352e353 transplantation, 360e361, 361f Photosensitive peptides or proteins, 670 Phototransduction, 352 PHS. See US Public Health Service (PHS) PHS Act. See Public Health Service Act (PHS Act) PHSRN. See Pro-His-Ser-Arg-Asn (PHSRN) Physical entrapment methods, 657 Physical stress evaluating functional restoration, 432 mechanical environment, 418e423, 418f boundary value problems, 423 constitutive relations, 421e423 strain and stress, 418e421 mechanical regulation of vascularized tissue regeneration, 427e432 role of mechanical stimuli, 423e427 Physically cross-linked hydrogels, 637 Physicochemical modifications, 651, 652f Physicochemical surface modifications, 653e655 chemical modifications, 653 topographical modifications, 653e655 PI3K. See Phosphatidylinositol 3’kinase (PI3K) PIDs. See Primary immunodeficiencies (PIDs) Piezo receptors, 394 Piezoelectric inkjet printers, 832e833 PiggyBac transposon system, 750, 754, 1203 PILOT. See Pediatric Intensity Level of Therapy (PILOT) PiolaeKirchoff stress tensor, 421 PIP2. See Phosphatidylinositol bisphosphate (PIP2) PIPAAm. See Poly(Nisopropylacrylamide) (PIPAAm) PKC. See Protein kinase C (PKC) PKP. See Penetrating keratoplasty (PKP) PKU. See Phenylketonuria (PKU) PLA. See Poly-L-arginine (PLA); Poly(lactic acid) (PLA) Placebo injection sites, 690 Placenta, function, origin, and composition, 133e134 Placental growth factor (PlGF), 311, 1301e1303 Placental hematopoiesis, 192

1413 Planar cell polarity pathway (PCP pathway), 283, 399 Planned Parenthood, 1310 Plasma cells, 680 deposition, 656 INF-gamma, 1154e1155 membranes, 392, 425 Plasma rich in growth factors (PRGFs), 1298e1299 Plasmacytoid DCs, 727e729 Plasmid, 748 Plasminogen activator, 3 Plastic contact lenses, 559 Platelet endothelial cell adhesion molecule-1. See Cell surface antigen CD31 Platelet-derived growth factor (PDGF), 20, 66, 76, 182e183, 221, 311, 438, 667, 677, 892e893, 914, 1183e1184, 1298e1299 PDGF-A chains, 893 PDGF-B chains, 893 PDGF-BB, 1091, 1298e1299 PDGFR-a, 251e252 Platelet-endothelial cell adhesion molecule-1 (PECAM-1), 296 Platelet-rich plasma (PRP), 413, 703, 942e943, 1186, 1298e1299 Platelet(s), 923, 929e931 activation, 998 from adult stem cells and somatic cells, 929 biogenesis, 929 efficiency for in vitro platelet production, 930e931 factor 4, 66 lysate, 208 megakaryocytes and platelets generating from hPSCs, 930 “Platform” technologies, 531 PLCL. See Poly(L-lactide and ε-caprolactone) (PLCL) Plerixafor, 310 PLGA. See Poly(lactic-co-glycolic acid) (PLGA) PlGF. See Placental growth factor (PlGF) PLLA. See Poly(L-lactic acid) (PLLA) PLO. See Poly-L-ornithine (PLO) “Plug and play” design philosophy, 1370 Pluripotent cells (PCs), 963 Pluripotent stem cells (PSCs), 50, 181, 255e256, 336e341, 454e455, 909, 923, 1134, 1168e1169, 1251, 1309e1310 Pluronic F127, 821 Pluronics. See Poly(propylene oxide) (PPO) PMA. See Premarket Approval Application (PMA) PMCs. See Primary mesenchyme cells (PMCs) PMMA. See Poly(methyl methacrylate) (PMMA)

1414 PNiPAAm. See Poly(Nisopropylacrylamide) (PNiPAAm) PNS. See Peripheral nervous system (PNS) POC. See Poly(1, 8-octanediol-co-citrate) (POC) Podocalyxin expression (PDX expression), 1167 Podocytes, 143, 1149e1150 POEs. See Polyorthoesters (POEs) Poiseuille flow, 429 Poisson ratio, 419 Poloxamers. See Poly(propylene oxide) (PPO) Poly 3-hydroxybutyrate (PHB), 545, 1226 Poly 3-hydroxybutyrate-co-3hydroxyhexanoate (PHBHHx), 545 Poly 3-hydroxyvalerate (PHBV), 545 Poly I:C. See Polyinosinicepolycytidylic acid (poly I:C) Poly-4-hydroxybutyrate (P4HB), 319e320 Poly-L-arginine (PLA), 702 Poly-L-lysine, 491 Poly-L-ornithine (PLO), 702 Poly-lactide-co-glycolide. See Polylactic-coglycolic acid (PLGA) Poly(1,8-octanediol-co-citrate) (POC), 569f, 571 Poly(2-hydroxyethyl methacrylate-comethyl methacrylate) (P(HEMA/ MMA)), 1208 Poly(2-hydroxyethyl methacrylate) (PHEMA), 562e563, 631e632 Poly(amide carbonate), 569f Poly(amino acids), 573e574 Poly(anhydride), 569f Poly(anhydrides-co-imides), 574e575 Poly(bis[methoxyethoxyethoxy] phosphazene), 575 Poly(butylene succinate), 571 Poly(caprolactone). See Poly(εcaprolactone) (PCL) Poly(D,L-lactic acid) (PD,LLA), 568, 569f, 860 Poly(D,L-lactide-co-glycolide). See Poly(lactic-co-glycolic acid) (PLGA) Poly(D,L-lactide) and ε-caprolactone (PDLCL), 1134 Poly(di[carboxylatophenoxy] phosphazene), 575 Poly(diol citrates), 571 Poly(DTE carbonate), 573 Poly(ether ether ketone) (PEEK), 899 Poly(ethylene glycol) (PEG), 443, 490, 512, 561, 564, 564f, 579, 632, 647, 657e659, 720e721, 788, 811, 1044, 1116 bioactive forms as exemplars of increasing sophistication, 632e633 block copolymers of polyesters or polyamides with, 572 PEG-based hydrogels, 285, 579, 637e638, 666, 835e836, 1049, 1172 surfaces, 531

INDEX

Poly(ethylene glycol) diacrylate (PEGDA), 808, 811, 834e835 Poly(ethylene terephthalate), 319, 561, 564f, 566 Poly(ethylene-co-vinyl acetate) (PEVAc), 561e562, 562f Poly(ethylene) (PE), 561e562, 562f and derivatives, 561e562 tubes, 1255 Poly(glycerol adipate), 569f, 571 Poly(glycerol dicarboxylate), 569f Poly(glycerol sebacate) (PGS), 569f, 571, 1050, 1077 Poly(glycolide-co-trimethylene carbonate) (PCTMC), 439 Poly(hydroxy-alkanoates) (PHAs), 570 Poly(L-lactic acid) (PLLA), 439, 568, 811, 813, 857, 857f Poly(L-lactide and ε-caprolactone) (PLCL), 1035e1036 Poly(L-lactide). See Poly(L-lactic acid) (PLLA) Poly(lactic acid) (PLA), 493e494, 568, 569f, 666, 941, 1045, 1049, 1077, 1226, 1268e1269 Poly(lactic-co-glycolic acid) (PLGA), 359, 452, 494, 568, 569f, 669, 690, 708, 727, 811, 813, 858, 892, 941e942, 964, 1048, 1104, 1118, 1204e1205, 1240, 1268e1269 PLGAeb-TCP skeleton, 708e709 scaffolds, 617 Poly(lactide-co-caprolactone), 1226 Poly(lactide). See Poly(lactic acid) (PLA) Poly(LLA-co-CL). See Poly(L-lactide-coε-caprolactone) (Poly(LLA-co-CL)) Poly(LLA-co-DXO). See Poly(L-lactide-co1,5-dioxepan-2-one) (Poly(LLA-coDXO)) Poly(meth)acrylates, 562e564 Poly(methyl methacrylate) (PMMA), 447, 559, 562e563, 562f, 575e576 Poly(N-acryloylmorpholine) (PAcMo), 479e480 Poly(n-butyl methacrylate) (PBMA), 478 Poly(n-butylcyanoacrylate( (PBCA), 496 Poly(N-isopropylacrylamide) (PIPAAm), 470e471 Poly(N-isopropylacrylamide) (PNiPAAm), 469e470, 495, 562e564, 648 Poly(NiPAAm-copolylactideehydroxyethyl methacrylate-co-acrylic acid-co-Nacryloxysuccinimide), 563e564 Poly(octanediol citrate), 569f Poly(ortho ester), 569f Poly(p-dioxanone), 569f, 570 Poly(phosphanzene), 569f Poly(propylene fumarate-co-ethylene glycol) (P[PF-co-EG]), 576f, 577 Poly(propylene fumarate) (PPF), 561, 576e577, 576f, 806e807, 811, 813

Poly(propylene oxide) (PPO), 565 Poly(propylene) (PP), 561e562, 562f Poly(sodium styrene sulfonate) (polyNaSS), 709 Poly(styrene) (PS), 561e562, 562f Poly(tetrafluoroethylene) (PTFE), 562, 562f Poly(trimethylene carbonate) (PTMC), 569f, 573 Poly(vinyl alcohol) (PVA), 632, 788 Poly(a-hydroxy acids), 568, 570 Poly(ε-caprolactone-co-lactide) (PCLA), 1230 Poly(ε-caprolactone) (PCL), 141, 439, 494, 549e550, 569f, 570, 700e701, 811e812, 823, 834e835, 857, 892, 941e942, 1034e1035, 1044 poly(ε-caprolactone/D, L-lactide)-based scaffolds, 554 scaffold, 488, 806e807, 1104, 1268e1269 Polyacrylamide(s), 562e564 hydrogels, 669 Polyacrylic acid, 601 Polyacrylonitrile (PAN), 1151 Polyacrylonitrile/polyvinylchloride (PAN/PVC), 1226 Polyamides, 561 with PEG, 572 Polyanhydride (PA), 561, 574, 669, 860 Polyaniline, 450 Polybutylene terephthalate (PBT), 704 Polycarbonates, 572 Polycomb repressive complex 2 (PRC2), 55 Polydimethylsiloxane (PDMS), 441, 1080, 1139e1140 Polydopamine coating, 441 Polyelectrolyte complexation (PEC), 702 Polyelectrolyte multilayers, 441 Polyesters, 508e509, 567e572, 705 block copolymers of polyesters or polyamides with PEG, 572 fibers, 568 of a-hydroxy acids, 568e570 of lactones, 570 POEs, 571 polycarbonates, 572 of polyols and carboxylic acids, 571 Polyethers, 561, 564e565 Polyethylene glycol-conjugated adenosine deaminase (PEG-ADA), 198 Polyethylene oxide (PEO), 440, 704 Polyglycerols, 565 Polyglycidols. See Polyglycerols Polyglycolic acid (PGA), 319e320, 439, 568, 569f, 666, 813, 910e911, 940e941, 1033e1034, 1104, 1169e1170, 1226, 1255, 1268e1269 Polyglycolide. See Polyglycolic acid (PGA) Polyhydroxyalkanoates (PHAs), 545e546, 1049 advantages and disadvantages of natural biomaterial, 1050t

INDEX

in bone tissue engineering applications, 546 processing methods, 545e546 Polyhydroxybutyrate-hydroxyvalerate (PHB-HV), 546 Polyhydroxyhexanoate, 545 Polyhydroxyoctanoate (PHO), 545, 1045e1046 Polyhydroxyvalerate, 545 Polyiminocarbonates, 572 Polyinosinicepolycytidylic acid (Poly I:C), 722 Polylactideehydroxyethyl methacrylate (HEMAPLA), 563e564 Polymerase chain reaction (PCR), 93, 1016e1017 amplification, 97 Polymeric/polymer, 545, 749, 831, 892 biomaterials, 507e508 blends, 544 chains, 507e508 coatings, 655 containing acrylate, 579 demixing, 447 liposome system, 722 membrane, 687 nanobiomaterials, 493 polymer-based precipitation, 209 polymer-ceramic blends, 698, 708e709 polymeric-based biomaterials, 892 polymeric-based nanobiomaterials, 492t, 493 polymerepolymer blends, 708 synthesis, 560e561, 811 witheCeCebackbone, 561e567 Polymerization mechanisms, 637e638 Polymerized collagen (P-collagen), 1172 Polymorphonuclear leukocytes, 676e677 polyNaSS. See Poly(sodium styrene sulfonate) (polyNaSS) Polyolefins, 636e637 Polyols, polyesters of, 571 Polyorthoesters (POEs), 571 Polypeptides, 632e633 Polyphosphazene(s), 575, 708 polyphosphazeneepolyester blends, 705 Polypyrrole (PPY), 450, 1229 Polys, 676e677 Polysaccharide(s), 601, 639e642, 816 agarose, 641 alginate, 640e641 cellulose, 613, 641e642 chitin and derivatives, 641 hyaluronic acid, 639e640 mucosal layer, 440 Polysiloxane(s), 565 gels, 565 Polystyrene microbeads, 1118 Polysulfone hollow fibers, 1156e1157 Polysurgery approach, 1076 Polytetrafluoroethylene (PTFE), 319, 1029 Polyurethanes (PUs), 559e560, 566, 566f, 572e573, 812, 845e846, 1050 hydrolytically stable PUs, 566e567

polyurethane-urea matrices, 573 Polyvinyl alcohol (PVA), 494, 544, 1049e1050 POMP. See Periocular mesenchymal precursor (POMP) POP. See Pelvic organ prolapse (POP) Population assays, 94e95, 94f Population doubling (PD), 1266 Porcine lung xenografts for transplantation, 1065 Porcine small intestinal submucosa extracellular matrix (SIS-ECM), 619 Porcine type I collagen, 616 Porcine urinary bladder matrix extracellular matrix (UBM-ECM), 619 Pore size, 1151 Porites Goniopora coral, 707 Porogens, 539 Porosity, 664e665, 1226, 1268, 1270 of CPCs, 602e604 Porous and fibrous 3D scaffold, 1075 Porous bioactive glass-ceramic, 706e707 Porous CaSiO3, 706e707 Porous HA-CHI-alginate composite scaffolds, 702 Porous scaffolds, 550, 628e629, 670, 1076 Porous solids, 630 Porous structures effect, 510 “Positive selection” technique, 208 Positron-emission tomography (PET), 371 Postmitotic organ, 247, 1264 Postprocessing, 620 Posttraumatic OA (PTOA), 937 Potassium and strontium ions scaffolds (K/Sr scaffolds), 709 Pou4f3 expression, 872 PP. See Poly(propylene) (PP) PPF. See Poly(propylene fumarate) (PPF) PPO. See Poly(propylene oxide) (PPO) PPY. See Polypyrrole (PPY) PRC2. See Polycomb repressive complex 2 (PRC2) Precision, 95 engineering, 95e96 options to control proteins at interphases with, 531 protein signaling, 523 techniques and technologies for precision immobilization, 527e531 collagen to control protein orientation, 529 HA for protein signal delivery and orientation control, 531 hexahistidine tags, 528 ionic charge and charge control of orientation, 528e529 proteins immobilization in lipid layers and tethered lipid bilayers, 530 streptavidin for biomolecular orientation control, 530 Preclinical animal studies of in utero stem cell transplantation, 1011e1020

1415 Preclinical bone repair models in regenerative medicine biomineralization and bone regeneration, 761e762 cell sources, 762 embryonic stem cells, 762e763 preclinical models of bone tissue regeneration, 764e766 scaffolds, 763e764 Preclinical data supporting stem cell therapies for TBI cell types, 376 conventional cell delivery routes, 378e380 mechanisms of action, 376e377 novel cell delivery routes, 380e381 timing of infusion, 377 Preclinical development plan in FDA, 1356 Preclinical translation, 945 Prednisone, 999e1000 Preimplantation genetic diagnosis (PGD), 116, 126, 1332 Premarket Approval Application (PMA), 1348e1349 Prenatal period, 1009 Prenatal regenerative medicine, 1009 Preservation agents, 527 Pressure dressings and negative-pressure wound therapy, 79 Preterax and Diamicron Modified Release Controlled Evaluation trials, 258 Prevascularized skeletal muscle, 1088 Prevention of Contrast Renal Injury With Different Hydration Strategies trials, 258 Prevotella, 745 PRGFs. See Plasma rich in growth factors (PRGFs) Primary brain injury, 370e376 bloodebrain barrier permeability, 373e375 cerebral edema, 375e376 neuroinflammation, 371e373 Primary cell culture techniques, 1169 Primary heart field (FHF), 247e248 progenitors, 248 Primary Immune Deficiency Treatment Consortium of 41 North American centers, 198 Primary immunodeficiencies (PIDs), 156 Primary mesenchyme cells (PMCs), 1 Primary shear stressedriven signaling pathway, 636 Primed embryonic stem cells, 49e50 Primed mESCs, 50 Primitive CFU-s, 194 Primitive ectoderm, 247e248 Primitive erythropoiesis, 924 Primitive streak, 247e248 Primordial germ cell-like cells (PGCLCs), 1245 Primordial germ cells (PGC), 1019 Principal components analysis (PCA), 527

1416 Printability, 835 Printing process, 816e817, 1288 Private banks. See Family banks Private funding, 1313e1316 Pro Osteon 200R, 551 Pro-His-Ser-Arg-Asn (PHSRN), 443 Processing methods, 541e542 Prod-1, 45 Proepicardial organ (PEO), 247e248 progenitors, 248 Proerythroblast. See Pronormoblast Professional antigen-presenting cells, 684 Progenitor(s), 247e248 cells, 307 mobilization, 308 and stem cell heterogeneity, 93 acquiring single-cell data, 96e100 analyzing single-cell data, 100e103 clinical implications of cellular heterogeneity, 105e108 population vs. single-cell assays, 94f single-cell isolation, 95e96 subpopulation determination, 103e105 Programmed cell death protein-1 (PD-1), 716e717 Progressive familial intrahepatic cholestasis (PFIC), 235 Prohibition of Human Cloning for Reproduction and Regulation of Human Embryo Research Amendment Bill 2006, 1321 Proinflammatory cytokines, 41, 721 Proliferation, 788 celleECM interactions during healing of cutaneous wounds, 27 during regenerative fetal wound healing, 29 signal transduction events during celleECM interactions, 22e23 Proline, 640 Pronormoblast, 924 Proof-of-concept system, 1082 toxicity, 174 Prophylactic vaccination, 727e729 Propranolol, 780 Prospective Randomized Study of Mesenchymal Stem Cell Therapy in Patients Undergoing Cardiac Surgery trial, 258 Prostaglandin E2 (PGE2), 7e8, 205, 425 Prostaglandin-2, 298 Prostatectomy procedures, 1223 Prosthetic silicone implants, 565 Proteases, 645 Protein 63 (p63), 1060e1061 Protein kinase C (PKC), 38e39 Protein zero (P0), 250e251 Protein(s), 440, 523e525, 529, 748e749, 816, 976, 1267 adsorption, 525, 657

INDEX

collagen to control protein orientation, 529 controlled with precision methods and supporting tools, 525e527 precision control of proteins at interfaces, 523e524 and role in precision delivery of biological signals, 524e525 surface analysis techniques and technologies for precision immobilization, 527e531 immobilization in lipid layers and tethered lipid bilayers, 530 parameters, 657 and peptides, 642e646 collagen and derivatives, 642e643 elastin derivatives, 643e644 fibrin derivatives, 644e645 self-assembled peptides, 645e646 silk, 645 protein G, 531 protein-replacement therapy, 1015e1016 protein-resistant brush polymers, 656 proteinebiomaterial interfaces, 524e525 signal delivery, 523 HA for, 531 stability, 5 Proteoglycans, 26, 393, 617e618, 1180 Proteolytic enzymes, 308e309, 637 Protospacer-adjacent motif sequence (PAM sequence), 742 Prototype model, 405e406 Provisional matrix formation, 677 Proximodistal translation (PD translation), 43 PRP. See Platelet-rich plasma (PRP) PRR. See Pattern recognition receptor (PRR) PS. See Poly(styrene) (PS) PSCs. See Pluripotent stem cells (PSCs) Pseudoelasticity, 422 Pseudomonas elodea (P. elodea), 543 Psychiatric diseases, 172e173 PT. See Patellar tendon (PT) PTFE. See Poly(tetrafluoroethylene) (PTFE); Polytetrafluoroethylene (PTFE) PTH. See Parathyroid hormone (PTH) PTMC. See Poly(trimethylene carbonate) (PTMC) PTOA. See Posttraumatic OA (PTOA) Public banks, 150e151 Public cord blood banking procedures, 151e156 collection techniques, 152 cord blood unit characterization, 154e156 donor recruitment and consent, 151 processing and cryopreservation, 153e154 volume and cell count considerations, 152e153 Public Health Service Act (PHS Act), 1346 Pulmonary “first-pass” effect, 378, 379f

Pulmonary arterial hypertension (PAH), 319 Pulmonary fibrosis, 107 Pulp tissue devitalization, 907 Pulpedentin regeneration, 913 Pulsatile intramural pressures, 424 Pulsed dye laser therapy, 78 PuraMatrix, 703e704 Pure inorganic materials, 605 Purified collagen, 642e643 Purkinje fibers, 450 PUs. See Polyurethanes (PUs) PVA. See Poly(vinyl alcohol) (PVA); Polyvinyl alcohol (PVA) PVS. See Penile vibratory stimulation (PVS); Persistent vegetative state (PVS) Pyridinoline cross-linked carboxyterminal telopeptide of type I collagen, 893 Pyridyl disulfide (PDS), 724 PDSeCpG nanoparticles, 724

Q

QCM-D. See Quartz crystal microbalance with dissipation monitoring (QCMD) QD. See Quantum dots (QD) QHREDGS peptide, 1090 qPCR. See Quantitative polymerase chain reaction (qPCR) Quail-donor Hensen node cells transplantation, 248 Quality control, 101, 1369 Quantitative polymerase chain reaction (qPCR), 97 Quantum dots (QD), 490e491 Quartz crystal microbalance with dissipation monitoring (QCM-D), 526e527 Quasilinear viscoelastic theory, 1181e1182 Quiescent satellite cells, 274

R

R-Smads, 52 RA. See Retinoic acid (RA) Rabbit models, 765 Rabbit penile defect model, 1257 RAC. See Recombinant DNA Advisory Committee (RAC) Rac1 protein, 772 Rac1. See Ras-related C3 botulinum toxin substrate 1 (Rac1) Rac1b, 3 RAD. See Renal assist device (RAD) Radiation grafting, 656 Radiation therapy, 79 Radio-frequency glow discharge plasma deposition, 656 RAFT. See Reversible additionfragmentation chain transfer (RAFT) Rapamycin, 1172 Rapid prototyping techniques, 95e96, 552 Ras-related C3 botulinum toxin substrate 1 (Rac1), 398

INDEX

Rat extraembryonic endodermal precursor (rXENP), 182e183 RBCs. See Red blood cells (RBCs) Reactive astrocytes, 1200 Reactive cyclic acetal polymers, 578 Reactive oxygen species (ROS), 3, 636e637, 880, 1200 Real-time sensor technologies, 1159 Rebuilding functional lung tissue, 1065e1068 ReCell kit, 1289e1290 Recellularization, 790, 1105 of biological materials, 1270 of organ scaffolds, 621e622 Receptor for hyaluronate-mediated motility (RHAMM), 16e18 Receptor tyrosine kinase ligands (RTK ligands), 6 signaling by, 7 Receptor tyrosine kinases, 18 Recognized consensus standard, 1357 Recombinant DNA Advisory Committee (RAC), 1355 Recombinant human BMPs (rhBMPs), 893 rhBMP-2, 702 Recombinant human collagen hydrogels (RHCIII hydrogels), 1116 Recombinant human growth hormone (rhGH), 690 Recombinase, 747e748 Recombination human bone morphogenetic protein-2 (rhBMP-2), 540, 894 Recommended ionized calcium range (riCa range), 1155 Recruitment factors, 727e729, 732f RECs. See Renal epithelial cells (RECs) Red blood cells (RBCs), 153e154, 924 erythropoiesis, 924 generating from adult stem cells in vitro, 924e925 generating from hESCs, 925e927 generating from human iPSCs, 927e928 generation by direct conversion of somatic cells, 928 manufacture of safe and effective RBC substitutes, 928e929 universal blood generation by modifying RBC surface antigens, 924 Reduced GO (rGO), 451 Reepithelialization, 26, 38e39, 68 Refinement process, 761 Regeneration, 683e684, 1059e1060, 1281 of articular cartilage surface and lubrication, 413 of corneal layers, 1118e1119 of diseased tissues, 662e663 in native lung, 1060e1061 ovarian tissue from stem cells, 1244e1245 road blocks to, 871 of skin, 1285 Regenerative healing cytokines and growth factors, 75e77

targeting inflammatory response, 73e75 Regenerative medicine, 229, 295, 321e322, 405e406, 407f, 411, 485, 550, 559, 4613, 627, 628f, 907, 1009, 1061, 1131, 1142, 1165 adult populations, 1041 applying to keratoprosthesis development, 1116e1118 biodegradable synthetic polymers for, 567e580 biologic scaffolds composed of ECM, 613e626 biomaterial interfaces in, 651 bioreactors in, 787e803 of bladder, 1263e1279 challenges in use of satellite cells in, 284e285 clinical experience with in utero stem cell transplantation, 1020e1021 clinical options, 1042 biological, 1042, 1043t mechanical, 1042, 1043t for engineering human hair follicle, 1297e1308 for female reproductive system, 1237e1250 fetal development and, 1009e1011 gene editing in, 741e759 HVD, 1041 implications for, 30e31 for kidney, 1165e1177 for male reproductive system, 1237e1250 MSCs in, 219e228 nanobiomaterials, 491e493 nanomaterial properties, 486e491 nanotechnology-based stem cell therapy, 497e498 nanotechnology-based strategies in, 493e496 for personalized medicine, 769e786 precision control of proteins in, 523e524 preclinical animal studies of in utero stem cell transplantation, 1011e1020 preclinical bone repair models in, 761e768 regulations and guidance of special interest for, 1350e1356 of respiratory tract, 1059e1072 and surgery of articular cartilage, 412 TEHVs, 1042e1046 young populations, 1041e1042 Regenerative Medicine Advanced Therapy (RMAT), 1370, 1374 Regenerative medicine manufacturing, 1367 bioprinting, 1369 current challenges, 1368t current opportunities, 1370t envisioned manufacturing systems of future, 1370e1373, 1371t, 1372f global landscape, 1373e1375, 1373t domestic efforts, 1374 international efforts, 1374e1375

1417 technical societies, 1373e1374 lack of standards, 1368 logistics, 1368 primary challenges for widespread adoption, 1367e1368 scale-up and automation, 1369e1370 sensors and quality control systems, 1369 Regenerative Medicine Manufacturing Society (RMMS), 1373e1374 Regenerative signals, 894 Regenerative therapies, 252e253 RegenMed Development Organization, 1372 Regulation cell orientation in cell sheet engineering arrangement of 3D orientation, 480 intelligent surfaces for, 479e480 skeletal muscle tissue engineering, 480e481 of human cells and tissues, 1350e1351 Regulatory T cells, 1268 Rehabilitation, 1206 Renal assist device (RAD), 1152e1153, 1153f immunomodulatory effect, 1154e1155 therapy of acute kidney injury causing by sepsis, 1153e1154 Renal cells, 1168 Renal epithelial cells (RECs), 1154e1155 Renal fibrosis, 107 Renal papilla, 1167 Renal proximal tubule cells (RPTC), 1150 Renal replacement device, requirements of, 1149e1150 Renal replacement therapies (RRT), 1149e1150 Renal tubule, 1167 REPAIR-AMI study, 257 Repeat variable di-residue (RVD), 746 Replacement, 761. See also Cellreplacement therapy; Extracorporeal renal replacement b cells for replacement therapy, 336 androgen replacement therapy, 1254, 1254f ERT, 158, 198 need for replacement tissues, 661 organ replacement therapies, 927e928 Repopulation, methods to improving, 238e239 Representative tests, 689, 689t Reproductive cloning, 1311, 1335 Reprogramming, 58, 336 factors, 170 methods, 50, 169e171, 171f Request for Designation, 1348 Research cloning. See Therapeutic cloning Research conduct, 1338 Resident cells, 613e615 Resorbable bioceramics, 550 Resorbable polymers, 559e560 Resorbable tricalcium phosphate, 891e892

1418 Respiratory tract, regenerative medicine of advances in rebuilding functional lung tissue, 1065e1068 biological scaffolds to support regeneration, 1063e1065 clinical translation and future considerations, 1068 lung development, 1059e1060 novel cell populations for lung repair, 1061e1063 repair and regeneration in native lung, 1060e1061 Responding stem cells, 405e406 Reticular dermis, 66 Reticuloendothelial systems, 236 Retina, 352e354, 352f, 754e755 Retinal degeneration, 1208e1212 biomolecule delivery, 1210e1211 cell transplantation, 1211e1212 Retinal disorders, 351 Retinal ganglion cells transmit neuronal stimuli, 352e353 Retinal organoids, three-dimensional, 363 Retinal pigment epithelium (RPE), 176, 356e359, 358f, 1208e1210 cells, 119, 1202 transplantation, 359e360 scaffolds for, 359e360 surgical techniques for, 360 Retinal progenitor cells (RPCs), 351, 362, 1211 Retinal stem cells (RSCs), 1200e1201, 1212 Retinitis pigmentosa (RP), 351, 354e355, 754e755, 1199 Retinitis pigmentosa GTPase regulator gene, 754e755 Retinoic acid (RA), 339e340, 879e880 Retrotransposons. See Class I transposons Retroviral transduction, 170 Rett syndrome, 173 Reverse process of MET, 1 Reverse transcriptaseepolymerase chain reaction (RT-PCR), 97, 700 Reversible addition-fragmentation chain transfer (RAFT), 471 Reversine, 40 Rex1, 52e53, 114 Rezulin, 780 RGD. See Arginine-glycine-aspartate (RGD) RGDS. See Arginine-glycine-aspartic acidserine (RGDS) rGO. See Reduced GO (rGO) RHAMM. See Receptor for hyaluronatemediated motility (RHAMM) rhBMP-2. See Recombination human bone morphogenetic protein-2 (rhBMP-2) rhBMPs. See Recombinant human BMPs (rhBMPs) RHCIII hydrogels. See Recombinant human collagen hydrogels (RHCIII hydrogels) RHD. See Rheumatic heart disease (RHD) Rhesus monkeys, 690

INDEX

Rheumatic heart disease (RHD), 1041 Rheumatoid arthritis, 317 rhGH. See Recombinant human growth hormone (rhGH) Rho GTPases, 459 Rho-associated, coiled-coil containing protein kinase (ROCK), 398, 1118e1119, 1201 Rho-associated, coiled-coil containing protein kinase inhibitor (ROCKi), 1118e1119 Y27632, 119 Rho123. See Rhodamine 123 (Rho123) Rhodamine 123 (Rho123), 195 rhPDGF-A-BB, 893 Ribonucleoprotein complexes, 748e749, 753 riCa range. See Recommended ionized calcium range (riCa range) Riluzole, 1206e1207 Ring-opening polymerizations (ROPs), 561 rMAPCs. See Rodent MAPCs (rMAPCs) RMAT. See Regenerative Medicine Advanced Therapy (RMAT) RMMS. See Regenerative Medicine Manufacturing Society (RMMS) RNA, 39, 75, 209, 748 RNA-guided RNase, 745e746 RNA-induced silencing complex, 57 virus, 170 RNA liposome complex (RNA-LPX), 724, 726f RNA seq. See RNA sequencing (RNA seq) RNA sequencing (RNA seq), 95 RNA-LPX. See RNA liposome complex (RNA-LPX) RNAi. See Interference RNA (RNAi) RNases, 748 Roboticeuniversal force-moment sensor (UFS), 1183 Robust cell source, 1156 Robust protocol mimicking neuronal development, 176 ROCK. See Rho-associated, coiled-coil containing protein kinase (ROCK) Rod-shaped nanoparticles, 720 Rodent MAPCs (rMAPCs), 182 characteristics, 183 express primitive endoderm markers, 182 isolation of, 182e183 rMAPCs in vitro, differentiation potential of, 183 Rodent models, 175, 1271 Rodents, 765 Rofecoxib, 780 Root canal treatment or extraction, 907 Root segments (RSs), 913 ROPs. See Ring-opening polymerizations (ROPs) ROS. See Reactive oxygen species (ROS) Roscovitine, 173 Rotary jet spinning, 1077 RP. See Retinitis pigmentosa (RP)

RPCs. See Retinal progenitor cells (RPCs) RPE. See Retinal pigment epithelium (RPE) RPTC. See Renal proximal tubule cells (RPTC) RRT. See Renal replacement therapies (RRT) RSCs. See Retinal stem cells (RSCs) RSs. See Root segments (RSs) RT-PCR. See Reverse transcriptaseepolymerase chain reaction (RT-PCR) RTK ligands. See Receptor tyrosine kinase ligands (RTK ligands) Rubriblast. See Pronormoblast Runt-related transcription factor 2 (Runx2), 425, 450, 703, 854, 957 Runx2. See Runt-related transcription factor 2 (Runx2) RUVBL-1 and-2, 55 RuvC domains, 742 RVD. See Repeat variable di-residue (RVD) rXENP. See Rat extraembryonic endodermal precursor (rXENP)

S

S-E ratio. See Stimulator to effector ratio (S-E ratio) S. aureus Cas9 (SaCas9), 749e750 Saccule, 867 Sacrificial bioinks, 821e823, 822f Sacrificial Pluronic F127, 836e837 Sad1p, UNC-84 domain (SUN domain), 459 Safety, 1202 and biocompatibility requirements for biomaterial scaffolds, 513e517 foreign body response, 516e517 hemocompatibility, 515 infection and sterilization, 514e515 toxicity, 515 Safety of Regenerative Medicine Act, 1375 SAH. See Subarachnoid hemorrhage (SAH) SAL. See Sterility assurance level (SAL) Saline, 1303e1304 Salt leaching, 547 SAMs. See Self-assembled monolayers (SAMs) Sarcolemma, 273 Sarcomeric banding, 1081 Sarcomeric structures, 1079 Sarcopenia, 283 SASP. See Senescence-associated secretory phenotype (SASP) Satellite cell(s), 273, 279e280 satellite cell-associated transcription factors, 277 self-renewal mechanisms, 281e283 Satellite cellederived myoblasts (SCDMs), 971e975, 973fe975f, 979f formation of graft-derived satellite cells, 975 formation of new myofibers, 973e974 gene complementation, 972e973 relevant properties, 972e975

INDEX

Sauerbrey model, 526e527 SBB. See Single blastomere biopsy (SBB) SBS. See Synthetic bone substitute (SBS) SC-b cells. See Stem cellederived b cells (SC-b; cells) SCA1. See Stem cell antigen-1 (SCA1) Scaffold(ing), 505, 510, 559, 675, 763e764, 788, 799, 910e911, 939, 1268, 1288 approaches in bone tissue engineering, 697e699 biochemical signaling, 763e764 combination of cell sheet engineering and scaffold-based engineering, 477e478 degradation, 507 design, 856, 1044 of ECM and biomimetic biomaterials, 409 ECM substitutes and, 80 functions, 696e697 hemocompatibility, 515 lacked fibrous microstructure, 1051 material, 359, 699e709 ceramic scaffolds, 705e707 intact and solubilized ECMs as, 619e622 metallic scaffolds, 707 natural polymers, 699e704 synthetic polymers, 704e705 microstructure, 535 morphology, 509, 511e513 anisotropic and gradient scaffolds, 511e512 injectable scaffolds/controlling morphology in situ, 513 methods for fabricating porous scaffolds, 511 surface feature manipulation, 512e513 porous and highly interconnecting, 856 for retinal pigment epithelium transplantation, 359e360 scaffold-based therapy, 1075 scaffold-free cell printing, 838 scaffold-infiltrating DCs, 727e729 sequence, 742 surface modification and functionalization, 859 systems, 1169 tracking, 513 use, 1188 Scale-up, 1369e1370 Scanning electron microscopy, 793e794 Scar, 662 formation, 30, 66e68 classic stages of wound repair, 69f inflammatory cell recruitment to site of tissue damage, 68f reduction theory, 73e77 cytokines and growth factors, 75e77 targeting inflammatory response, 73e75 tissue, 68 “Scar in jar” system, 1288 Scarless

fetal wounds, 29 wound healing adult skin, 66e72 current therapeutic interventions, 77e80 fetal skin, 72e73 future therapeutic interventions, 80e84 perspective, 85e86 regenerative healing and scar reduction theory, 73e77 Scarring, 65, 1285 Scatter factor. See Hepatocyte growth factor (HGF) scBS-seq. See Single-cell genome-wide bisulfate sequencing (scBS-seq) SCD. See Selective cytopheretic device (SCD); Sickle cell disease (SCD) SCDMs. See Satellite cellederived myoblasts (SCDMs) SCF. See Stem cell factor (SCF) Schwann cells (SCs), 635, 1224e1225 SCI. See Spinal cord injury (SCI) SCID. See Severe combined immunodeficiency (SCID) SCIPIO trial. See Cardiac stem cells in patients with ischemic cardiomyopathy trial (SCIPIO trial) SCNT. See Somatic cell nuclear transfer (SCNT) SCRO committee. See Stem cell research oversight committee (SCRO committee) SCs. See Schwann cells (SCs) SCS. See Silicified collagen scaffold (SCS) SDF-1. See Stromal cell-derived factor-1 (SDF-1) SDF-1a. See Stromal-derived growth factor-1 (SDF-1a) SDFs. See Syngeneic dermal fibroblasts (SDFs) SDOs. See Standards Development Organizations (SDOs) SDS. See Sodium dodecyl sulfate (SDS) Seager Electroejaculator, 1255e1256 SeaPrep agarose hydrogel, 1226 Seaweeds, 640 Second intention wound healing, 684 Second-generation biomaterials, 559e560 Secondary brain injury, 370e376 bloodebrain barrier permeability, 373e375 cerebral edema, 375e376 neuroinflammation, 371e373 Secondary heart field (SHF), 247e248 progenitors, 248 Secondary ion mass spectrometry analysis (SIMS analysis), 525, 527e529, 529f Secreted protein acidic and rich in cysteine protein (SPARC protein), 21 Secreted proteome of MAPCs, 186e187

1419 Secretoglobulin 1A1. See Club cell secretory protein, 10 kD (CC10) Seeding neuronal support cells for nerve regeneration, 1228e1229 Selective cytopheretic device (SCD), 1155e1156. See also Renal assist device (RAD) Selective laser sintering (SLS), 807 Self-antigens, 715 Self-assembled monolayers (SAMs), 441, 528e529, 656 Self-assembled peptides, 645e646, 698e699 self-assembled peptide-amphiphiles, 574 Self-assembling RAD16-I, 698e699 Self-assembly, 490, 633e634 Self-renewal, 281e282 cellebioactive surface interactions, 445e446 Semen cryopreservation, 1252 Seminiferous tubules, 1251e1252 Semiporous membrane, 773 Sendai virus, 170 Senescence-associated secretory phenotype (SASP), 283 Sensitivity, 95 Sensors, 1369 Sepsis, renal assist device therapy of acute kidney injury causing by, 1153e1154 Serial dilution series, 1109 Sericin, 699 Serine protease inhibitors, 1030 Serum-free culture system, 925e926 Setting/hardening mechanism of CaP, 595e599 chemical reaction, 596e597 acidebase interaction, 596 hydrolysis interaction, 596e597 setting times, 597e598 strategies to improve setting times, 598e599 Severe combined immunodeficiency (SCID), 197e198, 1013, 1020 beige mice, 135 hematopoietic stem cell transplantation for, 197e198 SF. See Silk fibroin (SF) SFF technologies. See Solid free-form fabrication technologies (SFF technologies) SFG spectroscopy. See Sum frequency generation spectroscopy (SFG spectroscopy) SGE. See Spinal ejaculation generator (SGE) SGLT1. See Na+-glucose cotransporter 1 (SGLT1) SH2, 220e221 SH2B3 gene, 754 SH3, 220e221 SH4, 220e221 Shape of nanomaterials, 488e489 Shear injury, 370e371

1420 Shear strain, 419 Shear stress (SS), 421, 429e431, 431f, 1155 SHEDs. See Stem cells from human exfoliated deciduous teeth (SHEDs) b-Sheets, 645e646 SHF. See Secondary heart field (SHF) shh gene. See Sonic hedgehog gene (shh gene) sHLA-I. See Soluble human leukocyte antigen-I (sHLA-I) Short bowel syndrome, 1135 Short interfering RNAs (siRNAs), 721, 915 Short-term-HSCs (ST-HSCs), 107 Sickle cell disease (SCD), 157 Side population (SP), 195 Signal transducer and activator of transcription 3 (STAT3), 49e50, 52, 117, 722 expression, 284 inhibitor NSC74859, 119 Signal for bone, 405e406 transduction events during celleECM interactions, 18e24 adhesion and migration, 19e22 apoptosis, 24 differentiation, 23e24 proliferation and survival, 22e23 Signaling cascades, 391 molecules, 339e340 Silanol groups, 550 Silica, 605 Silicate ceramics, 553e554 in bone tissue engineering applications, 553e554 processing methods, 553 Silicate-substituted calcium phosphate, 554 Silicified collagen scaffold (SCS), 701 Silicone(s), 559e560, 565 elastomers, 565 gel sheets, 79 Silk, 645, 699e700, 1125, 1271 powders, 699e700 Silk-I, 547 Silk-II, 547 Silk fibroin (SF), 547e548, 645, 699, 1125, 1270e1271 in bone tissue engineering applications, 547e548 processing methods, 547 Silkworm (Bombyx mori), 547 Silver nanoparticles (AgNPs), 492, 1125 SIM. See Simvastatin (SIM) Simple hydrogel, 632 Simple limbal epithelial transplantation (SLET), 1116 SIMS analysis. See Secondary ion mass spectrometry analysis (SIMS analysis) Simvastatin (SIM), 702 Single blastomere, 116 Single blastomere biopsy (SBB), 125e126

INDEX

Single celletype systems, 783 Single gene mutation, 157 Single integral finite strain theory, 1181e1182 Single nucleotide polymorphism risk variant, 174 Single transplants of hepatocytes, 236 Single-anastomosis model, 1232 Single-cell approaches, 169e170 epigenetics, 100 genomics, 97 isolation, 95e96 proteomics, 99e100 qPCR, 107e108 RNA seq, 103, 107e108 RT-PCR analysis, 101 technologies, 93 transcriptomics, 97e99, 98f Single-cell data acquiring, 96e100 single-cell epigenetics, 100 single-cell genomics, 97 single-cell proteomics, 99e100 single-cell transcriptomics, 97e99, 98f analyzing, 100e103 mathematical identification of cellular subpopulations, 101e103 noise reduction in single-cell data, 100e101 normalization, 101 Single-cell genome-wide bisulfate sequencing (scBS-seq), 100 “Single-platform” approaches, 154 Single-stranded breaks, 743 Single-walled carbon nanotubes (SWNT), 451, 576e577 Sipuleucel-T, 716, 727 siRNAs. See Short interfering RNAs (siRNAs) Sirolimus, 998 SIRS. See Systemic inflammatory response syndrome (SIRS) SIS. See Small intestinal submucosa (SIS); Small-intestinal submucosa (SIS) SIS-ECM. See Porcine small intestinal submucosa extracellular matrix (SIS-ECM) Six1 homeoprotein, 4 Size of nanomaterials, 487e488 SJS. See StevenseJohnson syndrome (SJS) Skeletal muscle, 248e250, 273, 281, 494, 845e846, 963e964 cell transplantation in, 972e975 clinical trials of cell therapies in muscular dystrophies, 965t skeletal muscleederived stem cells, 261 stem cells, 296 challenges in use of satellite cells in regenerative medicine, 284e285 functional characteristics of muscle stem cells, 276e277 gene editing strategies, 285 isolation of muscle stem cells, 277

molecular characteristics of muscle stem cells, 274e276, 275f muscle stem celleintrinsic defects in aging and disease, 283e284 regulation of muscle stem cells by niche, 279e281 satellite cell self-renewal mechanisms, 281e283 satellite cells, 274 tracking muscle stem cell behavior through live imaging, 278e279 types within muscle, 285e286 tissue engineering, 480e481 transplantation in, 971e975 Skeletal myoblasts, 261, 473e474 Skin, 496, 754, 846, 1281 regeneration, 81 repair, 1282, 1287 tissue engineering current TE skin technologies, 1287e1289 development, anatomy, and function of skin, 1283e1285 future, 1290e1291 potential prerequisite requirements, 1285e1287 TE skin solutions in clinical practice, 1289e1290 wound healing, 66 SLET. See Simple limbal epithelial transplantation (SLET) Slow-resorbing hydroxyapatite, 891e892 SLS. See Selective laser sintering (SLS) Smad gene, 42, 409, 426 SMAD4, 52 Smad-interacting protein-1 (Sip1). See Zeb homeobox 2 (Zeb2) Small animal lungs regeneration, bioreactors for, 789e790, 789f Small bowel transplantation, 1135 Small intestinal submucosa (SIS), 618e619, 690, 1030, 1132, 1136, 1188e1190, 1190f, 1267, 1270 Small intestine, 1135e1141 Small-intestinal submucosa (SIS), 1239 Smart manufacturing, 1370e1372 SMC. See Smooth muscle cells (SMC) Smooth muscle, 494 differentiation markers, 1267 smooth muscle-specific marker, 1267 tissues, 1263 Smooth muscle cells (SMC), 314, 477e478, 478f, 495, 838, 1029e1030, 1257, 1264 Smooth-surface implants, 682 Smurf1 E3 ubiquitin ligase, 3 Snail family of zinc-finger transcription factors, 4 Snail-1 or Snail-2 activator, 4e5 SNCA gene, 754 Sodium alginate, 601 Sodium bicarbonate (NaHCO3), 539, 836 Sodium dodecyl sulfate (SDS), 1046e1047, 1104, 1241 Soft assignment, 103 Soft hydrogel, 834e835

INDEX

Soft lithography, 447, 447t, 633, 659 Solegel method, 551e553 Solid free-form fabrication technologies (SFF technologies), 511, 763 Solid organ transplantation, MAPCs as immunomodulation in, 185 Solid scaffolds, 550 Solid tumor cancers, 721e722 Solubilized ECMs as scaffold material, 619e622 decellularization, 619e620 host response, 621 hydrogels, 620e621 postprocessing, 620 tissue procurement, 619 whole-organ scaffolds, 621e622 Soluble human leukocyte antigen-I (sHLA-I), 234 Soluble mediators, 689 Soluble plasma fibronectin, 616 Soluble polymers, 657e659 Solvent evaporation, 545 Somatic cell nuclear transfer (SCNT), 115, 1311 Somatic cells, 929, 1332 direct conversion, 928 Sonic hedgehog gene (shh gene), 45, 339e340, 911 Sorbitol, 527 South Korea, stem cell research in, 1321 SOX17. See SRY-related HMG-box 17 (SOX17) Sox2, 50, 114 SOX2. See SRY-box 2 (SOX2) Sox9-EGFP multipotent intestinal epithelial stem cells, 1135 SP. See Side population (SP) “Spacer sequence”, 742 SPARC protein. See Secreted protein acidic and rich in cysteine protein (SPARC protein) “Spare” embryos, 1310 Spatial heterogeneity, 633e635 Spatiotemporal delivery, 727 SpCas 9 components of multiplex gene preeclustered, 746f protein, 742 SpCas9-HF, 745 SpCas9-HF1, 745 SpCas9/gRNA, 742 variants and orthologues, 744e745 Sperm banking, 1252 Spermatogonia, 1252 Spermatogonial stem cell technology (SSC technology), 1251e1254, 1252f Spheroid reservoir bioartificial liver (SRBAL), 1107 Spheroids, 773 Spider silk, 645 Spinal cord identity, 1208 Spinal cord injury (SCI), 185, 1199, 1255e1256 Spinal cord regeneration, 1208

Spinal ejaculation generator (SGE), 1255e1256 SPIONs. See Superparamagnetic iron oxide nanoparticles (SPIONs) Spleen, 231 Split-liver procedures, 237e238 Split-thickness skin grafting (SSG), 1285, 1289 Sponges, 552 Spongiosa, 1044 Sponsors, 1345 SPR. See Surface plasmon resonance (SPR) SR1. See Stem-regenin (SR1) SRBAL. See Spheroid reservoir bioartificial liver (SRBAL) Src-mediated FAK phosphorylation, 19e20 SRY-box 2 (SOX2), 52e54, 426 SRY-related HMG-box 17 (SOX17), 339e340 SS. See Shear stress (SS) SSC technology. See Spermatogonial stem cell technology (SSC technology) SSEAs. See Stage-specific antigens (SSEAs) SSG. See Split-thickness skin grafting (SSG) ST-HSCs. See Short-term-HSCs (ST-HSCs) Stage-specific antigens (SSEAs), 134e135 SSEA-1, 114 Stainless steel, 559e560 Stainless steel 316 L, 707 Standardization, 1370 Standardized uptake values (SUV), 373, 375f Standards Development Organizations (SDOs), 1357e1358 Standards development program in FDA, 1357e1358 Staphylococcus aureus (S. aureus), 898, 1125 Staphylococcus pyogenes (S. pyogenes), 745 Starch, 548e550 in bone tissue engineering applications, 550 processing methods, 549e550 starch-based blends, 549 Stargardt macular dystrophy (STGD), 351, 354 STAT3. See Signal transducer and activator of transcription 3 (STAT3) State policy, 1313e1316 Stem Cell Act (2005), 150 Stem cell antigen-1 (SCA1), 251e252 markers, 277 Sca-1+ CSCs, 263e264 Stem cell factor (SCF), 924e925 Stem Cell Institute of New Jersey, 1314 Stem Cell Oversight Committee, 1320e1321 Stem Cell Research Enhancement Act, 1312e1313 Stem cell research oversight committee (SCRO committee), 1337 Stem Cell Research Oversight/Embryo Research Oversight Committee Review, 1337

1421 Stem cell therapy, 954 for erectile dysfunction, 1258 for musculoskeletal diseases, 953e964 articular cartilage, 956e958 bone, 955e956 challenges and prospects, 964e966 IVD, 962e963 meniscus, 961e962 osteochondral tissue, 958e959 regulatory and financial challenges to stem cell therapies, 953e954 skeletal muscle, 963e964 tendon and ligament, 959e960 tendonebone interface, 960e961 nanotechnology-based, 497e498 Stem cell(s), 81e84, 82f, 409, 750, 762, 1044, 1136, 1168e1169, 1251, 1267 and alternative cell sources for liver therapy, 239e241 in bone tissue engineering, 854e856 in clinical research and clinical applications, 1337e1340 clinical application, 1340 clinical translation, 1338e1340 donor and procurement issues, 1337e1338 research conduct, 1338 SCRO committee/EMRO committee, 1337 combined stem cell therapeutics, 265e266 delivery, 498 differentiation, 445e446 for diseases of retina AMD, 354 cell-based neuroprotection, 361e362 cell-replacement therapy, 355e361 disease-in-a-dish modeling for retinal disorders, 362e363 retina, 352e354 retinitis pigmentosa, 354e355 STGD, 354 epidermal stem cells, 83e84 ES cells, 81e82 exceptionalism, 1337 expansion, 498 iPS cells, 84 MSCs, 82e83 nature of MSCs, 206e207 niche, 629 regenerating ovarian tissue from, 1244e1245 research, 1312 animalehuman chimeras, 1325e1326 commercialization and access to treatments, 1324e1325 compensating egg donors, 1323e1324 ethical, legal, social, and policy questions of, 1321e1326 research guidelines, 1316e1318 ISSCR, 1318 NAS, 1316e1318 SC-seeded venous grafts, 1224e1225 sources, 1264e1266, 1265t, 1309e1311

1422 Stem cell(s) (Continued ) adult cells, 1309 embryos, 1310e1311 fetal cells, 1309e1310 iPScs, 1311 stem cellebased therapies, 354, 376, 1241 stem cellerelated genes, 255e256 transfection, 497e498 transplantation, 995e997, 1009e1010 Stem cellederived b cells (SC-b cells), 336, 340 Stem cells from human exfoliated deciduous teeth (SHEDs), 909 Stem hematopoietic cells, 240 Stem-regenin (SR1), 157 Stemness genes, 39e40, 206 Stereocilia, 867e868 Stereolithography, 805e806, 808, 819e820 Sterility assurance level (SAL), 514 Sterilization, 514e515 methods, 620e621 patient risk factors, 514f Steroid methylprednisolone, 1206e1207 StevenseJohnson syndrome (SJS), 472, 1118 STGD. See Stargardt macular dystrophy (STGD) Stiffness, 439e440, 1181 matrix, 422 Stimulation of endogenous repair, 247 Stimulator to effector ratio (S-E ratio), 184 Stomach, 1134e1135 Storage starch, 548 Strain, 418e421, 419f, 1045, 1181 energy density, 1181 Stratum corneum production, 1284 Streptavidin for biomolecular orientation control, 530 Streptococcus mutans (S. mutans), 898e899 Streptococcus pyogenes (S. pyogenes), 742 Stress, 418e421, 1181 relaxation, 1181e1182 stressestrain curve, 1181, 1182f relationship, 1179 Stretch-activated ion channels. See Mechanosensitive ion channels STRO1, 206 Stroke, 160e161, 223, 1203e1206 cell transplantation, 1205e1206 factors for endogenous stem cell stimulation, 1203e1205 pharmacological therapy, 1203 Stroma, 662e663 Stromal cell-derived factor-1 (SDF-1), 193, 250e251, 280, 701, 913e914 SDF1/CXCR4, 250e251 Stromal cell-surface markers, 258 Stromal vascular fraction (SVF), 222, 295e296 Stromal-derived growth factor-1 (SDF-1a), 308 Structural preservation, 527 Structural protein, 529

INDEX

Structureeproperty relationships in hydrogels, 631e632 Subarachnoid hemorrhage (SAH), 370, 371f Submucosa, 1135 Subpopulation determination, 103e105, 104f cell-based therapies development, 104e105 isolating best cell for given clinical application, 105f Substrates importance, 438 Subtotal cystectomy reservoirs, 1274 Subtunic penile progenitor cells, 1258 Subunit vaccines, 716 b Subunits of hemoglobin (HBB), 753 Subventricular zone (SVZ), 181e182, 1200e1201 Sulfation, 617e618 Sum frequency generation spectroscopy (SFG spectroscopy), 525e526 SUN domain. See Sad1p, UNC-84 domain (SUN domain) Supercritical assisted phase-in version process, 549 Supercritical CO2 sterilization, 620e621 Superparamagnetic iron oxide nanoparticles (SPIONs), 491 Supply chain, 1373 Supporting baths, 823e825 Supporting bioinks, 823e825 Supporting cells, 867 Suppressive immune cell populations, 715 Suprapubic catheterization, 1273 Sural nerve, 1231 Surface analysis methods and supporting tools, 525e527 AFM, 527 conformational stabilization for biomolecules, 527 iodine 125elabeled proteins, 527 multivariate, 527 QCM-D, 526e527 SFG spectroscopy, 525e526 SPR, 526 ToF-SIMS, 525 XPS, 525 and role in precision delivery of biological signals, 524e525 technologies, 653, 654t Surface charge, 440 Surface chemical patterning, 659 Surface chemistry, 441, 490 methods of altering, 441e442 Surface coatings, 651 Surface feature manipulation, 512e513 Surface modification for degradation control, 509e510 strategies, 651e653 Surface nanotopography, 494 Surface plasmon resonance (SPR), 526 Surface roughness, 653e655, 655f Surface stability, 651e653 Surface to volume ratio, 509

Surface topography, 489e490, 653e655, 655f Surface wettability, 440 Surface-eroding polymers, 574 Surfactant production, 1059e1060 Surgery, 80 Surgical complications, 992 Surgical nerve injuries, 1223 Surgical techniques, 237e238, 1132 for retinal pigment epithelium transplantation, 360 “Surplus” embryos, 1310 SUSD2. See Sushi domain containing-2 (SUSD2) Sushi domain containing-2 (SUSD2), 1246 SUV. See Standardized uptake values (SUV) SVF. See Stromal vascular fraction (SVF) SVZ. See Subventricular zone (SVZ) Sweat glands, 66 SWNT. See Single-walled carbon nanotubes (SWNT) Syndecans, 21, 618 extracellular domains, 16e18 Syngeneic dermal fibroblasts (SDFs), 915 Syngenic nerve grafts, 1232 Synthetic 3D scaffolds, 453 Synthetic biomaterials, 1049, 1052t PEG hydrogels, 1049 Synthetic bone substitute (SBS), 893 Synthetic HAP, 493, 551 Synthetic hydrogels, 835e836 increasing sophistication of synthetic hydrogels for TE, 632e639 bioactive forms of poly(ethylene glycol) as exemplars, 632e633 hydrogel degradation, 636e637 injectable systems, 638e639 matrix mechanics, 635e636 polymerization mechanisms, 637e638 spatial heterogeneity, 633e635 systems, 1172 for TE templates, 646e648 Synthetic materials, 559, 811e813, 812f for histogenesis of new organs, 669e670 hydrogels, 669e670 hydrolytically degradable polymers, 669 and tissue grafting, 661 Synthetic membranes, 1151 Synthetic NGCs, 1226 Synthetic peptides, 574 Synthetic polymers, 452, 559, 567, 704e705, 837, 892, 1033e1035, 1151, 1169e1170, 1255, 1268e1269. See also Natural polymers applications, 580 applications of synthetic polymers, 580 biodegradable synthetic polymers, 567e580 copolymers, 704e705 nondegradable synthetic polymers, 561e567 polyesters, 705 polymer synthesis, 560e561

INDEX

with seeded cells, 1035e1037 Synthetic scaffold(s), 763, 941e942, 1268e1270 biodegradable properties, 1269e1270 materials, 613 for nerve repair, 1225e1226 porosity, 1270 Systemic inflammatory response syndrome (SIRS), 1150

T

T cell(s), 205, 862 antigen receptors, 198 depletion, 196 implantable biomaterial scaffolds to enhancing autologous T cell therapy, 731e733 macroporous alginate scaffold, 732f T cell-stimulating cytokines, 722 T cellemediated process, 239 therapies, 716 in vitro effects of MAPCs on, 184 T lymphocytes, 685, 688 T-cell factor 4 (TCF4), 210 “T-cell piggyback” nanoparticle systems, 727 T-helper (Th), 344 immune response, 690, 715 Th1, 184, 620, 715 Th17 cytokine production, 184 Th2, 715 helper lymphocytes, 682 immune response, 715 T-regulatory cells (T-reg cells), 344, 715e716, 1013 T1D. See Type 1 diabetes (T1D) T2D. See Type 2 diabetes (T2D) 3T3 fibroblasts gels, 817 TA. See Tetraaniline (TA) Tacrolimus, 998 TALE. See Transcription activator-like effector (TALE) TALENs. See Transcription activator-like effector nucleases (TALENs) Tamoxifen-inducible Cre-reporter mouse lines, 248e250 TAMs. See Tumor-associated macrophages (TAMs) Target Fragile X syndrome, 172e173 Target Product Profile (TPP), 1356e1357 Targetable nucleases, 741e747 CRISPR, 741e746 TALENs, 746e747 ZFNs, 747 Targeted a-synuclein, 754 Targeting moieties, 720e721 TAZ. See Transcriptional coactivator with PDZ-binding domain (TAZ) TBI. See Traumatic brain injury (TBI) TBTE. See Thread-based tissue engineering (TBTE) Tbx18, 251e252 TC. See TCP-CHI composite (TC); Technical Committee (TC) TCF4. See T-cell factor 4 (TCF4)

TCP. See Tricalcium phosphate (TCP) a-TCP. See a-Tricalcium phosphate (a-TCP) b-TCP. See b-Tricalcium phosphate (b-TCP) TCP-CHI composite (TC), 703 TCPC. See Total cavopulmonary connection (TCPC) TCPS. See Tissue culture polystyrene (TCPS) tdLNs. See Tumor-draining lymph nodes (tdLNs) TE. See Tissue engineering (TE) TEA domain family (TEAD1e4), 426 TEBVs. See Tissue-engineered blood vessels (TEBVs) Technical Committee (TC), 1358 Technical societies, 1373e1374 TEER. See Transepithelial electrical resistance (TEER) Teflon, 559e560 Tegafur, 777 TEHVs. See Tissue engineered heart valves (TEHVs) “Template”, 629 Temporal sequence of inflammation and wound healing, 678, 678f Temporomandibular joint (TMJ), 897 TEMPs. See Tissue engineered medical products (TEMPs) Tenascin (TN), 393 TNC, 279 Tendinopathy, 391 Tendon(s), 845e846 healing of ligaments and, 1183e1185, 1188e1192 and ligament, 959e960 normal ligaments and biology, 1180 biomechanics, 1180e1183 tendonebone interface, 960e961 Tensile strength, 1181 Tension offloading, 80 Teratocarcinoma stem cells, 113e114 Terodiline, 780 Terpolymers, 572 TESA. See Tissue engineering by selfassembly (TESA) TeSR media, 118 Testes, 1251e1254 androgen replacement therapy, 1254, 1254f SSC technology, 1252e1254, 1252f Testosterone, 1254 Tethered lipid bilayers, proteins immobilization in, 530 TetOn sequence, 751 Tetraaniline (TA), 704e705 Tetracalcium phosphate (TTCP), 592e594, 706 TEVG. See Tissue engineered vascular graft (TEVG) TFs. See Transcription factors (TFs) TGF. See Transforming growth factor (TGF) Th. See T-helper (Th)

1423 Thalassemia, 157, 753 Therapeutic cloning, 1311 Therapeutic effects of MSCs, 222t Therapeutic intensity, 381 Therapeutic interventions, 77e80 bleomycin, 78 cryotherapy, 79 extracellular matrix substitutes and scaffolds, 80 5-FU, 77e78 growth factors and cell signaling molecules, 80e81 imiquimod, 78 laser therapy, 78 other drugs and biologics, 81 pressure dressings and negativepressure wound therapy, 79 radiation therapy, 79 silicone gel sheets, 79 stem cells, 81e84 surgery, 80 targeting gap junctions and Cx, 81 tension offloading, 80 topical and intralesional corticosteroid injections, 77 Thermally induced phase separation (TIPS), 856 Thermodynamics second law, 422 Thermoresponsive cell culture substrate, fabrication techniques of, 471 Thermoresponsive hydrogels, 648 Thermoresponsive liposome nanoparticles (TSL nanoparticles), 485 Thermoresponsive pentablock copolymers, 573 Thermoresponsive polymers for cell sheet engineering controlled grafting of thermoresponsive polymer on culture substrates, 470e471 thermoresponsive polymer for biomedical applications, 469e470 thermoresponsive surface for regulating cell adhesion and detachment, 470 variety of fabrication techniques of thermoresponsive cell culture substrate, 471 Thermoresponsive scaffolds, 454 Third-generation biomaterials, 559 Thiry-Vella loop, 1136 Thread-based tissue engineering (TBTE), 1032e1033 Three-dimension (3D), 762, 787 arrangement of 3D orientation using cell sheet layering techniques, 480 cardiac constructs, 774 cardiac tissue constructs, 1076 cell culture technology, 770 cell spheroids, 793 cellular responses to, 455e457 cervical-like tissue constructs, 1242 culture materials, 454, 455t models, 1106e1107

1424 Three-dimension (3D) (Continued ) systems, 1243 technologies, 454e455, 771 ECM, 1 environment, 437, 469, 1102e1103 folding, 615 human intestinal tissue, 1137 importance, 452 InSight Human Liver Microtissues of InSphero, 1108e1109 lamellar-like stromal tissue, 1123 lung scaffold, 789 microfabrication, 670 migration, 457 organoids, 779e780 plotting, 547 poly(carbonate) urethane scaffolds, 617 porous scaffold, 535 porous SF scaffolds, 699e700 printable Matrigeleagarose system, 1140 printed scaffolds, 453, 859 printing, 454, 763, 773, 831 reconstructions, 913 renal reconstructs, 1171 renal structures, 1165 retinal organoids, 363, 364f scaffolds, 542, 695, 955, 1268e1270 decellularized tissue, 453e454 electrospun/nanofibrous scaffold, 453 hydrogel scaffolds, 452e453 polymers for, 452 preparation, 452e454 three-dimensional printed scaffolds, 453 space, 419 stamping technique, 1088 structure, 58, 1073 substrates for three-dimensional culture, 452 techniques, 628e629 testis organoid system, 1253e1254 tissue constructs cell sheet layering technique, 474e475 coculture system based on cell sheet layering, 475 vascularization in cell sheets for largescale tissue construction, 475e477 tumor organoids, 775 Three-dimensional bioprinting (3D bioprinting), 453, 806f, 808, 811, 831e832, 832f, 1078e1079. See also Bioprinting mechanisms 3D bioprinted vascular structures, 838e839, 838f current translation of three-dimensional bioprinting, 825e826 supportive baths, 824f in vitro applications, 825e826 in vivo applications, 826 future perspectives, 847e848 tissue engineering applications, 841e847 for tissue regeneration applications, 843te844t

INDEX

variables critical to, 836f in vitro tissue models, 839e841 Three-dimensional printing (3DP), 805, 1044, 1281 bioinks, 808e826 fundamentals, 805e808 extrusion-based printing, 806e807 inkjet bioprinting, 807e808 SLS, 807 stereolithography, 808 future directions, 826 Thrombin, 815 Thrombogenicity, 515, 1087 Thrombopoiesis, 929 Thrombopoietin (TPO), 929 Thrombus formation, 676 Thy-1.1 antibodies, 194e195 Thy-1.2 antibodies, 194e195 Thymosin b-4 (Tb-4), 141e142 Tie6Ale4V, 707 Tight junction (TJ), 393, 1200 markers, 1267 Time-of-flight secondary ion mass spectrometry (ToF-SIMS), 525 Timing of infusion, 377 Timothy syndrome (TS), 173 TIMP3. See Tissue inhibitor of matrix metalloproteinase-3 (TIMP3) TIMPs. See Tissue inhibitors of metalloproteinases (TIMPs) TIPS. See Thermally induced phase separation (TIPS) Tissue culture polystyrene (TCPS), 439, 470e471, 523e524 Tissue engineered heart valves (TEHVs), 1042e1046 biomaterials for, 1046e1053 bioreactors, 1046 cell source, 1043e1044 implant design goals, 1045 implant function, 1044e1045 scaffold design, 1044 testing TEHV function, 1045e1046 in vivo conditioning and testing, 1046 Tissue engineered medical products (TEMPs), 1357e1358 Tissue engineered vascular graft (TEVG), 1029e1030, 1036fe1037f Tissue engineering (TE), 319e320, 505, 769, 772, 1251, 1263, 1281, 1368, 1372 applications, 841e847, 855 bone, 841e842 cardiac tissue and heart valves, 846 cartilage, 842 other tissue types, 847 POP, 1245e1246 skeletal muscle and tendon, 845e846 skin, 846 approaches, 254, 1165, 1271 for cartilage repair, 938e946 biological factors, 942e943 bioreactors, 943e946 bioscaffolds in cartilage repair, 939e942

cartilage surface modification, 938e939 cell types for cartilage repair, 939 revised approach to tissue engineering triad, 938f increasing sophistication of synthetic hydrogels for, 632e639 precision control of proteins in, 523e524 principles, 1237e1238, 1238f scaffolds, 629, 859, 899 skin, 1281 current TE skin technologies, 1287e1289 development, anatomy, and function of skin, 1283e1285 future, 1290e1291 potential prerequisite requirements for tissue engineered skin solutions, 1285e1287 solutions in clinical practice, 1289e1290 TE skin solutions in clinical practice, 1289e1290 technologies, 1287e1289 strategies, 662, 1074e1075 synthetic hydrogels for tissue engineering templates, 646e648 technologies, 770, 788e789, 891e899, 907 adjuvant therapies, 898e899 bioactive molecules, 892e893 bioreactors, 897e898 BMAC technique, 894e897 implantable scaffolds, 891e892 translation, 825 triad, 695 Tissue engineering by self-assembly (TESA), 1032e1033 Tissue inhibitor of matrix metalloproteinase-3 (TIMP3), 373e375 Tissue inhibitors of metalloproteinases (TIMPs), 39 Tissue plasminogen activator (tPA), 1200, 1203 Tissue Reference Group (TRG), 622 Tissue-engineered blood vessels MSCs, 207 culture systems, 429e430 deficiencies, 505 development, 391 cellular mechanotransduction mechanisms, 397e400 mechanotransduction mechanisms and major effectors, 392e396 nucleus as central organelle in regulating mechanotransduction, 396e397 engineered ovarian follicles, 1243e1244, 1244f enzymes, 636e637 explants, 793 fibrosis, 107 function, 317e318 injury, 277

INDEX

loss, 505 monitoring environment and tissue development, 799 muscle, 494e495 neural, 496 organoid types, 779 procurement, 619 regeneration, 317e319, 485, 490, 493, 498, 1263 3D bioprinting technologies, 843te844t models, 1271 remodeling, 423e424 repair, 662 in diabetes, 344e345 rules, 1350e1351 tissue-based therapy, 469 tissue-derived materials for hair regeneration, 1300 tissue-engineered constructs, 688 tissue-engineered implant, 675 tissue-guided regeneration, 1287e1288 tissue-material interactions, 675 tissue-specific, ECM-based bioinks, 837 tissue-specific stem cells, 274, 1251 vascular, 495e496 vascularization, 1088 Titanium, 559e560 TJ. See Tight junction (TJ) TLR. See Toll-like receptor (TLR) TMC. See Trimethylene carbonate (TMC) TMJ. See Temporomandibular joint (TMJ) TMPRSS4 type II serine protease, 3 TN. See Tenascin (TN) TNC. See Total nucleated cell (TNC) TNCC. See Total nucleated cell count (TNCC) TNF. See Tumor necrosis factor (TNF) TNF aestimulated gene 6 product (TSG-6 product), 377 ToF-SIMS. See Time-of-flight secondary ion mass spectrometry (ToF-SIMS) Tolerance induction, 1000 hematopoietic stem cell transplantation for, 198e199 Toll-like receptor (TLR), 721 Tooth development, 908e909, 908f Topical and intralesional corticosteroid injections, 77 Topographical cues, 446 cellular responses to, 447e450 Topographical features, 1287e1288 Topographical modifications, 653e655 Topography effect, 446 Total cavopulmonary connection (TCPC), 1036 Total heart transplantation, 495 Total nucleated cell (TNC), 149 Total nucleated cell count (TNCC), 150e151 Total parenteral nutrition, 1135 Totipotent stem cells, 181 Toxicity, 515 Toxicology, 777e778, 779f, 1108 screening, 779 Toxins, 869

tPA. See Tissue plasminogen activator (tPA) TPO. See Thrombopoietin (TPO) TPP. See Target Product Profile (TPP) Trabeculae, 696 Traceability and imaging, 513 Trachea, 791e792, 847 Traditional salt leaching technique, 578 Transcription activator-like effector (TALE), 746 Transcription activator-like effector nucleases (TALENs), 741, 746e747 identity of variable sequences in, 746t Transcription factors (TFs), 169, 171, 181, 210, 299, 394e395, 722, 927e928 TCF3, 53 transcription factorebased reprogramming, 171 Transcriptional coactivator with PDZbinding domain (TAZ), 426 Transcriptional enhancer factor domain family member, 394e395 Transdetermination, 336 Transdifferentiation, 376 of oral mucosa, 1118 process, 171, 869 Transendocardial Injection of Autologous Human Cells in Chronic Ischemic Left Ventricular Dysfunction trials, 258 Transepithelial electrical resistance (TEER), 1065 Transfection efficiency, 497e498 Transforming growth factor (TGF), 407e408, 909e910 TGF-b, 3, 52, 75, 222, 280, 339e340, 426, 496, 619, 677, 722, 880, 957, 1172, 1183e1184 superfamily, 6, 75, 1300 TGF-b1, 42, 66, 119, 205, 667, 939, 1298e1299 TGFb-mediated differentiation, 23e24 TGFbR, 18 Transfusable blood components, 925e926 Transgenic mice, 690 induction of hair cell regeneration using, 875e876 Transient receptor potential (TRP), 393e394 Transition wound, 73 Transition-metal coordinated mechanisms, 560e561 Transitional epithelium, 1263 Transitory starch, 548 Translation of cartilage tissue engineering preclinical translation, 945 Translocator protein, 373 Transmembrane heparan sulfate proteoglycans, 618 Transmembrane protein Myomixer, 276 Transmissible Spongiform Encephalitis (TSE), 1358 Transmission electron microscopy, 793e794, 1180

1425 Transplantation regulation of human cells and tissues intended for, 1350e1351 in skeletal muscles, 971e975 Transplantation of Human Embryonic Stem Cell-derived Progenitors in Severe Heart Failure trial, 255 Transplanted cells, 1203, 1253 Transposase, 747 Transposons, 747 Trauma, 405e406, 591, 907, 1263 Traumatic brain injury (TBI), 185e186, 369, 372f, 375f, 1199 classification, 370 clinical trials adipose-derived stem/stromal cells, 384 bone marrow mononuclear cell adult trial, 382e384 bone marrow mononuclear cell pediatric trial, 381e382 collagen scaffoldemesenchymal stromal cell study, 385 mesenchymal stromal cell neurologic stem cell treatment study, 385 mesenchymal stromal cell subacute TIB trial, 383e384 modified SB623 cells trial, 384 multicenter bone marrow mononuclear cell pediatric trial, 384 ongoing clinical trials, 384e385 current TBI management strategies, 376 epidemiology, 369e370 phases of brain injury, 370e376 preclinical data supporting stem cell therapies for, 376e381 Traumatic nerve injuries, 1223 Traumatic spinal cord injury, 1206e1208 biomolecule delivery, 1206e1207 cell transplantation, 1208 guiding axon regrowth, 1207e1208 Trehalose, 527 TRG. See Tissue Reference Group (TRG) Triblock PEGePCLePEG copolymer, 704e705 Tricalcium phosphate (TCP), 594, 703, 706, 708e709 a-Tricalcium phosphate (a-TCP), 592e593 b-Tricalcium phosphate (b-TCP), 540, 551, 577, 591e592, 765 Trigeminal nerve or cranial nerve V, 380 Triggered release systems, 721 Trimethylene carbonate (TMC), 572 TMC-based polycarbonates, 572 Trithorax group complexes, 57 Triton-X, 1241 Triton X-100, 1104 Troglitazone, 780 Trogocytosis process, 1014 Trophic effects, mechanisms of, 186e187 Trophoblast, 133 Tropocollagen, 615 Tropoelastin, 617 tropoelastin/elastin molecules, 643 Troponin, 299

1426 TRP. See Transient receptor potential (TRP) TRP2. See Tyrosinase-related protein-2 (TRP2) Trypsin, 117 TS. See Timothy syndrome (TS) TS iPSC model, 173 TSE. See Transmissible Spongiform Encephalitis (TSE) TSG-6 product. See TNF aestimulated gene 6 product (TSG-6 product) TSL nanoparticles. See Thermoresponsive liposome nanoparticles (TSL nanoparticles) TTCP. See Tetracalcium phosphate (TTCP) Tumor necrosis factor (TNF), 992, 1087 TNF-a, 280, 370, 679, 862, 1187 Tumor-associated macrophages (TAMs), 715 Tumor-draining lymph nodes (tdLNs), 724 Tumor(s), 715e716, 774e775 cell heterogeneity, 108 metastasis-on-a-chip platforms, 776 microphysiological systems, 774e775 models, 839 nanoparticle targeting of tumor microenvironment, 721e722 tumor-on-a-chip modeling, 775 tumor-reactive T cells, 731e733, 733f tumor-specific antigens, 715e716 tumor-targeting liposomes, 722 vessel endothelium, 314 Tumorigenesis, 301e302 21st Century Cure’s Act, 1374 Two-dimension (2D), 454, 769, 787 cell culture, 771 methods, 1243 monolayer cell cultures, 793 to 3D models, progression from, 770e771 Two-step MSC isolation protocols, 208 Tylenol, 778e779 Type 1 diabetes (T1D), 335, 345, 987 Type 1 receptors, 52 Type 2 diabetes (T2D), 281, 335, 987, 1155e1156 Type A gelatin, 643 Type I collagen, 542, 1103, 1118 Type I reaction. See Anaphylactic reaction Type II reaction. See Cytotoxic reaction Type II receptors, 52 Type IIB fibers, 276 Type III reaction. See Immune complex reaction Type IIS restriction endonucleases, 746e747 Type V reaction. See Cell-mediated delayed hypersensitivity reaction Type VI collagen, 279 Tyrosinase-related protein-2 (TRP2), 724 L-Tyrosine, 573 Tyrosine-based polycarbonates, 572 Tyrosinemia, 750e751 type 1, 235, 750e751

INDEX

Tb-4. See Thymosin b-4 (Tb-4)

U

U-curable materials, 811 UBM. See Urinary bladder matrix (UBM) UBM-ECM. See Porcine urinary bladder matrix extracellular matrix (UBMECM) UC-MSCs. See Umbilical cordederived MSCs (UC-MSCs) UCB. See Umbilical cord blood (UCB) UCs. See Urothelial cells (UCs) UCSF. See University of California, San Francisco (UCSF) UF. See Ultrafiltrate (UF) UFS. See Roboticeuniversal force-moment sensor (UFS) Ulcers, 345 Ultracentrifugation, 209 Ultrafiltrate (UF), 1149e1150 Ultralow-fouling zwitterionic hydrogels, 516e517 Ultrasound-guided transhepatic portal venous access, 992 Umbilical cord blood (UCB), 149, 181e182, 996 clinical uses, 156e157 CBT for hematological malignancies, 156 CBT for nonmalignant hematological diseases, 156e157 cord blood expansion technologies, 157 investigations in treatment of acquired brain injuries, 159e162 ASD, 161e162 cerebral palsy, 159e160 HIE, 159 stroke, 160e161 Umbilical cordederived MSCs (UCMSCs), 345, 700e701, 912e913 Umbilical tissue-derived stem cells, 362 Unbound delivery systems, 763e764 Underhealing, 72 Uniaxial tensile testing, 1181e1182 Uniform hydrogel, 637 United States Department of Health, Education and Welfare (DHEW), 1311e1312 United States federal and state stem cell policy, 1311e1316 current US stem cell research policy, 1313e1316 federal policy, 1313 state policy and private funding, 1313e1316 history, 1311e1313 Universal blood generated by modifying red blood cell surface antigens, 924 “Universal donor” RBCs, 926e927 University of California, San Francisco (UCSF), 993e994 Urea cycle defects, 236 Uremic toxins, 1151

Ureter cells, 1264 Urethra reconstruction, 1255, 1256f Urinary bladder, 1263 functions, 1263 Urinary bladder matrix (UBM), 380, 618e619 Urinary diversion, 1271e1273 Urinary tract system, 1264 Urine samples, 1266 Urine-derived stem cells (USCs), 1264 Urodele amphibians, 37 Urodele limbs regeneration blastema formation, 37e42 blastema growth, 42e45 Urothelial cells (UCs), 1264 Urothelial progenitor cells, 1264 Urothelial-specific cell markers, 1267 Urothelium, 1273e1274 US Code (USC), 1264, 1266, 1346 isolation, 1266 US Department of Defense, 907 US Federal Stem Cell Policy, 126b US Food and Drug Administration (FDA), 120, 150, 174, 222e223, 411, 513e514, 561e562, 622, 669, 716, 752, 776, 859, 939, 953e954, 1154, 1188, 1202, 1268e1269, 1338, 1345 advisory committee meetings, 1358e1359 approval mechanisms and clinical studies, 1348e1350 clinical development plan, 1356e1358 coordination efforts, 1361e1362 Critical Path research, 1360 guidelines for additive manufactured devices, 1369t laws, regulations, and guidance, 1346e1347 legislative history, 1345e1346 meetings with industry, professional groups, and sponsors, 1350 organization and jurisdictional issues, 1347e1348 preclinical development plan, 1356 regulations and guidance of special interest, 1350e1356 research and critical path science, 1359e1361 standards development program, 1357e1358 US National Academies, 1310e1311 US National Academies of Science (NAS), 1316e1318 US Public Health Service (PHS), 1353 US Renal Data System (USRDS), 1149 US Stem Cell Research Policy, 1309 ethical, legal, social, and policy questions of stem cell research, 1321e1326 international comparisons, 1318e1321 sources of stem cells, 1309e1311 stem cell research guidelines, 1316e1318

1427

INDEX

United States federal and state stem cell policy, 1311e1316 USC. See US Code (USC) USCs. See Urine-derived stem cells (USCs) USRDS. See US Renal Data System (USRDS) Uterine cervix tissue engineering, 1242 Uterine tissue regeneration, 1240e1242 Uteroglobulin. See Club cell secretory protein, 10 kD (CC10) Uterus, 1240e1242 uterine cervix tissue engineering, 1242 uterine tissue regeneration, 1240e1242 Utricle, 867 UV irradiation, 820e821

V

Vagina(l), 1238e1240 canal, 1238 engineering of functional vaginal tissue, 1239e1240 fibroblasts, 1245e1246 reconstruction surgery, 1239 surgical reconstruction, 1238 vaginal-shaped scaffolds, 1239 Vaginoplasty techniques, 1238 Valproic acid, 297e298, 775e776 Valve interstitial cells (VICs), 1043, 1043f Valvular endothelial cells (VECs), 1043, 1043f Vas deferens, engineering, 1255 Vascular cell adhesion molecule (VCAM), 276 vascular adhesion molecule-1, 308 Vascular disruption, 370e371 Vascular endothelial growth factor (VEGF), 7, 16, 20e21, 66e67, 76, 184, 210, 232, 238e239, 298, 311e313, 473e474, 497e498, 619, 696, 957, 1087, 1138e1139, 1232, 1241, 1271, 1298e1300, 1302 VEGF165, 1088 VEGFR-2, 311 Vascular healing process, 317e318 Vascular networks, 318, 668 Vascular niche, 193 Vascular permeability factor (VPF), 232 Vascular smooth muscle cells (VSMCs), 1087 Vascular system, 773e774 Vascular tissue, 495e496 Vascular-based pathologies, 780 Vascular-like fluidic devices, 773e774 Vascularization, 317e318, 697, 826, 1087e1088 in cell sheets for large-scale tissue construction, 475e477 in vivo bone bioreactors for solving vascularization problem, 797e798, 798f Vascularized tissue regeneration mechanical regulation, 427e432

mechanical stimulation in vitro, 429e432 mechanical stimulation in vivo, 427 Vasculogenesis process, 307e308, 313, 317 Vasogenic edema, 375e376 Vasomotor function, 257 VCAM. See Vascular cell adhesion molecule (VCAM) VCF. See Vertebral compression fracture (VCF) VECs. See Valvular endothelial cells (VECs) VEGF. See Vascular endothelial growth factor (VEGF) VEGF fused with collagen-binding domain (CBD-VEGF), 1271 Venograms, 231e232 Ventral blood island mesoderm, 192 Ventral pancreatic bud, 341 Ventricle-shaped chitosan tissue, 1083 Ventricular remodeling, 252e253 Ventricularis, 1044 VentriGel, 620e621 Vertebral compression fracture (VCF), 211 Vertebroplasty, 607 Vessel-on-a-chip, 773e774 Vestibulocochlear nerve, 867 Vicat needle method, 597e598 VICs. See Valve interstitial cells (VICs) Villi, 133 Villous syncytiotrophoblast, 133e134 Vimentin, 3 Vinyl chemical moieties, 669 N-Vinyl pyrrolidone (NVP), 576e577, 632 Vinylsulfone functionalities, 579 Virus-free human iPSCs, 933 Viruses, 749 Viscosity, 600 Viscosupplementation, 938e939 “Visible indentation”, 597e598 Vision, 1208e1210 Vitrasert, 562 Vitronectin, 393, 523e524 Voigt model, 526e527 Voluntariness, 1323 Von Willebrand factor (vWF), 1016, 1267 gene expression, 636 VPF. See Vascular permeability factor (VPF) Vroman effect of protein adsorption, 676 VSMCs. See Vascular smooth muscle cells (VSMCs) vWF. See Von Willebrand factor (vWF)

W

Wall shear stress (sw), 429 Wallace group, 1212 Wallerian degeneration, 1227e1228 WAT. See White adipose tissue (WAT) Water, 1180 molecules, 440 water-in-oil emulsification, 550

water-soluble polymers, 601 Wdr5 expression, 57 Wearable bioartificial kidney (WeBAK), 1158 in preclinical end-stage renal disease model, 1158 Wearable renal replacement therapies, future advancements for, 1159e1160 WeBAK. See Wearable bioartificial kidney (WeBAK) Wet AMD, 1208e1210 Wet chemical method, 551e552 Wet spinning process, 549e550 WGs. See Working Groups (WGs) White adipose tissue (WAT), 297, 854e855 White matter (WM), 378 Whitlockite, 489 Whole tooth engineering, 912e913 Whole-genome sequencing, 172 Whole-organ scaffolds, 621e622 Wild-type (WT), 276e277 bacterial systems, 742 muscles, 276e277 Williams syndrome, 173 Wilms Tumor 1 (WT1), 251e252 Wilson disease, 235 Wingless type (Wnt), 75e76 pathway, 6e7 signaling, 52e53, 76e77, 874 Wnt-CM, 1300 Wnt1-Cre reporter mice, 250e251 Wnt11, 6 WNt3a, 80e81 WM. See White matter (WM) Wnt. See Wingless type (Wnt) Wolff’s law, 417 Working Groups (WGs), 1358 World Health Organization, 513e514 Wound epidermis, 38 healing, 67e68, 320 cellular heterogeneity in, 106e107 inflammatory phase, 26 in skin, 1284 reepithelialization, 27 repair, 106 response and barriers to regeneration, 1200e1201 BBB, 1200 endogenous stem cells, 1200e1201 reactive astrocytes and glial scar, 1200 WT. See Wild-type (WT) WT1. See Wilms Tumor 1 (WT1)

X

X chromosomes, 240 X-linked severe combined immunodeficiency (X-SCID), 198 X-ray photoelectron spectroscopy (XPS), 525, 527 X-SCID. See X-linked severe combined immunodeficiency (X-SCID) Xanthan, 600

1428 Xanthomona genus, 746 Xeno-free defined media systems, 1372 Xenoderived, decellularized cornea stromas, 1123 Xenogeneic cellular cardiomyoplasty, 141e142 Xenogeneic collagen, 621 Xenogeneic ECMs, 621, 690 Xenogeneic model, 1246 Xenogenic matrices, 1030e1032 Xenogenous cells, 1264 Xenografts, 1042 Xenopus, 38 embryo, 43e44 X. laevis, 37

INDEX

Xenotransplantation, 995, 1353e1354 Xenotransplants, 239 XPS. See X-ray photoelectron spectroscopy (XPS)

Y

Y chromosomes, 240 Y397 autophosphorylation, 19e20 Yamanaka transcription factors, 169e170 Yellow fluorescent protein (YFP), 281e282 Yes-associated protein (YAP), 426 YAP1, 394e395 YGISR peptide, 442

Yolk sacederived CFU-s, 191 Young populations, 1041e1042 Young’s modulus, 280

Z

Zeb homeobox 1 (Zeb1), 4 Zeb homeobox 2 (Zeb2), 4 Zebrafish, 879e880 ZieglereNatta catalysts, 560e561 Zinc, 545 Zinc finger E-box-binding (Zeb), 2 Zinc-finger nucleases (ZFNs), 741, 747, 750 Zona occludens-1 (ZO-1), 449, 1157 Zwitterionic polymer surfaces, 531